3D Bioprinting for Lungs and Hollow Organs - PMC

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Transl Res. Author manuscript; available in PMC 2020 Sep 1.

Published in final edited form as:

Transl Res. 2019 Sep; 211: 19–34.

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The publisher's final edited version of this article is available at Transl Res

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Abstract

Three-dimensional bioprinting has been gaining attention as a potential method for creating biological tissues, supplementing the current arsenal of tissue engineering techniques. 3D bioprinting raises the possibility of reproducibly creating complex macro- and micro-scale architectures using multiple different cell types. This is promising for creation of multilayered hollow organs, which has been challenging using more traditional tissue engineering techniques. In this review, the state of the field in bioprinting of epithelialized hollow and tubular organs is discussed. Most of the progress for the pulmonary system has been restricted to the trachea. Due to the gross structural similarities and common engineering challenges when creating any epithelialized hollow organ, this review also covers current progress in printing within the gastrointestinal and genitourinary systems, as well as applications of traditional plastic printing in engineering these tissues.

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Introduction:

Over the last decade, 3D bioprinting has emerged as a promising technology to rapidly produce patient-specific tissue constructs.1 Bioprinting, as opposed to regular 3D printing, involves the layer-by-layer deposition of biologically-derived materials or cells.2 There are multiple modalities, but often they can be categorized as acellular printing with subsequent recellularization versus direct printing with cells suspended in a “bioink”.3 Compared to cellsheet engineering techniques, 3D bioprinting has an advantage of creating more complex architectures. It also has the potential to address problems of insufficient cell ingrowth into decellularized matrices, since cells can be introduced directly throughout the volume of the bioprinted construct. There is an ever-growing number of methods for material deposition, and each has its limitations. Laser-based printers can achieve single-cell resolution but cannot quickly produce large volumes of material, whereas extrusion-based printers can print quickly but are generally limited to a resolution of approximately 100 microns.4

Possibly the greatest benefit of 3D bioprinters in lung and airway tissue engineering is the ability to create hollow structures with multiple layers of different cells and materials. Although the lung is often classified surgically as a solid organ, for the purposes of tissue engineering it is more beneficial to think of the entire respiratory tree as a branching set of tubes. Thus, the progress and pitfalls discovered in bioprinting of other tubular organs can inform future research into bioprinting of the lung. Organs of the gastrointestinal and urinary systems have a general multilayered structure fairly similar to the large airways in the lungs. Commonalities include an inner lining of epithelium and concentric layers of supporting fibrous and muscular tissue. As the field of bioprinting is very young (and little headway has been made in bioprinting distal lung tissue at this time), this review focuses on current progress in bioprinting of such epithelial tubular organs. Given this nascent state, this review will also cover examples of other extrusion-based or layer-by-layer techniques, which can be considered ‘bioprinting-adjacent’, including more traditional printing-based approaches where the cellular component is added after the structure is 3D printed using a synthetic polymer.

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Lungs:

Treatment for end-stage pulmonary failure remains a significant clinical need. The paucity of donor organs and serious transplant-related complications have compelled researchers to use bioengineering approaches to create a clinically translatable lung graft. The leading approach in the literature has been whole lung decellularization and recellularization with histocompatible or autologously-derived cells to minimize immune-mediated rejection.5–9While this approach depends on a steady supply of donor organs, it might be feasible using scaffolds from human donor lungs that were not suitable for transplant, or from porcine sources. The biggest challenges with this approach have been vascular failure, and the inability to target the proper placement of the many cell types present in the lungs. Tissue engineering approaches to address these problems in many organs are underway, but the lungs present a unique set of challenges.

The lungs consist of 23 generations of bifurcating airways, ranging from 1.5 cm in diameter at the trachea to 0.5 mm at the respiratory bronchioles.10 The large airways from the main bronchi down through the terminal bronchioles are known as the conducting zone of the lungs and are not sites of gas exchange. The trachea and large airways are lined with a pseudostratified epithelium composed of goblet, surface serous, submucosal serous, submucosal mucous, brush, and club (formerly Clara) cells that produce mucus and remove debris via ciliary action. The smaller airways also contain pulmonary neuroendocrine cells (PNE), which secrete growth factors to neighboring epithelium and may help with hypoxia detection. The smallest generations of airways are the respiratory bronchioles, which lead into pulmonary acini made up of multiple, single alveoli and alveolar sacs (Figure 1). These structures are considered the respiratory zone of the lungs and are the sites of gas exchange. Alveoli are often described as small, interconnected spherical sacs with an approximate diameter of 200 microns. One estimate put the total number of alveoli in the human lung at 480×106.11 The alveolar septum, or wall separating two alveolar spaces, is only 4 cell layers thick (0.2 to 4 μm).1213 The air-contacting surface is lined with epithelial cells, with the endothelial cells of the alveolar capillaries lying directly below. In the alveoli, alveolar type I facilitate gas exchange with their associated endothelial cells and type II cells produce surfactants to maintain surface tension in the small volume of the alveolus.14

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Figure 1:

Illustration showing the pulmonary epithelial cells and their locations along the airways; MGC: mucous gland cells; SGC: serous gland cells; SMGs: submucosal glands; GC: goblet cells; PNE: pulmonary neuroendocrine; AI: alveolar type I cells; AII: alveolar type II cells. Reprinted from “Pulmonary Epithelium: Cell Types and Functions,” by M. M.-J. Chang, L. Shih, and R. Wu, 2008, In The Pulmonary Epithelium in Health and Disease, page 2. Copyright 2008 by “John Wiley and Sons”. Reprinted with permission.

The lung anatomy is further complicated by its dual vasculature. In the respiratory zone, deoxygenated blood coming from the right heart through the pulmonary vein is shuttled into progressively smaller vessels until it reaches the alveolar capillaries that wrap around the alveoli to replenish the blood’s oxygen and release carbon dioxide. This is sufficient blood supply for this region, because the alveolar walls are so thin that oxygen availability is not a concern. However, further up in the conducting airways the tissues are thicker and more metabolically active, and thus require an oxygenated blood supply. For this reason, a separate bronchial circulation is found running alongside the conducting airway system.15 It begins with bronchial arteries arising from the aorta, proceeds to capillaries that nourish the airways, then returns through the systemic circulation through the bronchial veins. The pulmonary and bronchial systems are not completely isolated, and a small amount of deoxygenated bronchial blood can travel into the pulmonary artery through microvascular connections.16 Since vascularization is a central challenge of all tissue engineering, this review will not discuss it at length, instead focusing on the specific challenges associated with alveolar printing. For further discussion, there are a variety of quality reviews specifically on vascularization.17

In all, the multiple cell types, range of size scales and necessity for connectivity both in the airways and in the vasculature pose significant fabrication challenges, requiring microscale resolution to achieve the key features for pulmonary function while still producing large volume grafts capable of holding up to six liters of air.

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Trachea:

The trachea is the major airway connecting the larynx and bronchi, allowing air to enter and exit the lungs. While the trachea is an airway, it is morphologically different than the lower airways of the lungs. It is made of 16–20 cartilage rings that prevent collapse. These rings are surrounded by smooth muscle and connective tissue.18 As the trachea transitions to the primary bronchi, the cartilage rings become a mosaic of cartilage plates in the bronchi.19 In contrast, the subsequent airways rely on the pressure difference between the atmosphere and pleural cavity, and surfactants to maintain patency.20

Current techniques for tracheal graft fabrication are valuable precursors to whole lung bioprinting. Overall, most 3D bioprinting techniques for tracheal construction combine biological hydrogels seeded with cells surrounding a sturdy polymer structure. The following section will highlight these emerging methods and examine their potential in future pulmonary bioprinting.

Dual Headed Printing:

In a recent study by Kaye et al., a dual-headed system was used to print both PCL and an alginate/collagen 1 hydrogel using a partial ring segment design.21 The hydrogel used was a 50/50 blend of 3% alginate and a 2.8 mg/ml collagen type 1 hydrogel, seeded with chondrocytes from rabbit tracheal cartilage. Using a dual-head extruder, a layer of PCL was printed, allowed to cool, then followed by a layer of hydrogel. The completed graft was then implanted in New Zealand white rabbits. The opening of the graft aligned with the trachealis muscle, which runs vertically on the posterior of the trachea. These grafts were harvested at three or six weeks after implantation. The cartilage formation was histologically evaluated using the modified O’Driscoll score, a histological scoring system that evaluates cartilage regeneration.22 While there was not a significant difference in scores between the two time points, the score was higher at either end of the graft where it overlapped native tissue compared to the center, indicating better cartilage growth. Unfortunately, all grafts had significant stenosis. The authors noted decreased inflammation where native tracheal epithelium had grown over the graft, acting as a boundary. The authors suggested that this boundary may decrease the shear stress on the graft from the air or isolate the graft from fluids in the airway. Either of these mechanisms could be replicated by seeding the airway with autologous epithelial cells in future constructs.

One major advantage of bioprinting, compared to other fabrication methods, is the ability to print complex patterns within a larger geometry. Exploring this benefit, Bae et al. produced a tracheal graft made of alternating layers of PCL and cell-laden alginate printed in overlapping patterns.23 A diagram of this method is shown in Figure 2 (concept illustrated in Figure 8 A). PCL was printed in the diagonal grids for the first and fifth layers, providing structural support while leaving negative space for nutrient transport. The center layers were helical form. The second and fourth layers were alginate seeded with epithelial cells and MSCs respectively, with a layer of PCL between them. This resulted in a mechanically robust graft. These grafts were implanted in New Zealand white rabbits and observed for twelve weeks. Over this time course, there were no significant respiratory complications or stenoses. Epithelial mucosa over the entire defect area was observed as well as localized neo-cartilage formation. This work was followed up by Park et al, with a longer in vivo study examining the grafts up to one year post implant.24 After 3 months histology revealed ciliated epithelium on the bioprinted grafts, while the PCL-alone controls had only granulated tissue. Immature cartilage islets were seen after 6 months, but they did not form cartilage rings. These promising survival results exemplify the benefits of 3D bioprinting, based on strategic patterning of both matrix and cells.

Figure 2:

Dual extruded PCL/alginate tracheal graft, A) longitudinal view, B) cross sectional view, C) SEM image with labeled layers (a) outer layer to (e) luminal layer. Reprinted from “3D Bioprinted Artificial Trachea with Epithelial Cells and Chondrogenic-Differentiated Bone Marrow-Derived Mesenchymal Stem Cells,” by S. Bae, K.-W. Lee, J. Park, et. al, 2018, Int. J. Mol. Sci., 19(6), pg. 5. Copyright 2018 by “MDPI”. Reprinted with permission.

Figure 8:

Schematic summary of hollow tissue fabrication methods. (A) Dual Head Extrusion Printing; multiple printing heads are used to print a construct of multiple materials with distinct regions. (B) Organoid Printing; spherical organoids are grown in a well plate, then assembled in a needle array and fused into a solid construct. (C) Sacrificial molding; a sacrificial mold is made using stereolithography, then the hydrogel mixture is injected and solidifies in the mold. (D) Scaffold prefabrication; a PCL scaffold is printed, then flooded with a cell laden hydrogel. (E) Decelled ECM supported with 3D printed PCL rings; rings are printed out of PCL and sheets of decellularized tissue are wrapped around them. (F) Electrospun PCL supported by 3D printed PCL rings; rings are printed out of PCL and then PCL is electrospun around them. (G) Tissue engineered esophagus made by printing PCL in a cross-hatched pattern, then coated with rabbit MSCs in a fibrin hydrogel. (H) Acellular esophagus graft with 3D printed PCL rings on a rotating mandrel, subsequently covered with PCL electrospinning. (I). 3D bioprinting used to create a two-layered intestinal epithelium model in a transwell plate. The first layer contained fibroblasts, which was covered by a second layer of epithelial cells. The tissue matured and organized in culture. (J) 3D bioprinting of a urethra graft by simultaneously printing PCL/PLCL blend, a SMC bioink outside, and a UC bioink inside.

Organoid Printing:

Shifting away from using polymers, Taniguchi et al. used a removable needle array to organize spherical organoids into a tracheal structure. Rat chondrocytes, endothelial cells, and MSCs were combined in pellet culture to form cell aggregates.25 Aggregated spheroids were then assembled into a tracheal construct in a needle array. After 7 days, the needles were removed, and a stent was placed in the lumen to prevent lumen compaction. The construct was moved to a bioreactor for 28 days. The mature construct was then implanted in an isogenic rat, with the stent included for 23 days. This process is illustrated in Figure 3 (concept illustrated in Figure 8 B). Microvessels with red blood cells were observed in the graft after 8 days in vivo. Further study is needed to verify whether the seeded endothelial cells accelerated in vivo vascularization, and whether the microvessels were host or donor-derived. This approach was expanded by Machino et al, incorporating layers of fibroelastic tissue between the cartilage organoids, more like the native trachea.26 Cartilaginous organoids were made by coculturing normal human articular chondrocytes (NHACs), human umbilical vein endothelial cells (HUVECs), and MSCs. Fibrous organoids were made using HUVECs, MSCs, and normal human dermal fibroblasts (NHDFs). These organoids were stacked in alternating layers in the needle array to mimic the rings of the trachea. These constructs were kept on the needle array for 7 days, then moved to a bioreactor to mature for 28 days. The mature grafts were then implanted in a rat model, supported by a silicone stent, and examined 7 days and 35 days after implant. Sections from 7 days in vivo show aggregates of CD31+ cells, indicating nascent vessel formation. At day 35 CD31+ staining showed more mature vessels with larger lumens. After 35 days in vivo, pan-cytokeratin positive cells were also present, suggesting the spread of the epithelium from the native tissue. This approach is well-suited for larger airways, but it may have challenges in creating smaller airways and vasculature. With small airways and alveoli, the lumens may be on a similar scale as the needles in the array, leaving significant holes in the tissue. To address this challenge, alternative support strategies should be investigated. Other printing methods use support baths of sacrificial materials, such as a slurry of gelatin microparticles in the freeform reversible embedding of suspended hydrogels (FRESH) method for printing low viscosity bioinks.27 While gelatin is not suitable for supporting tissue geometry long term, similar approaches using other materials may be explored.

Figure 3:

Experimental overview starting with isolation of cell types from F344 rat, expansion, spheroid culture, spheroid assembly, maturation, and implantation in another F344 rat. Reprinted from “Scaffold-free trachea regeneration by tissue engineering with bio-3D printing,” by D. Taniguchi, K. Matsumoto, T. Tsuchiya, et. al, 2018 Interact. Cardiovasc. Thorac. Surg., 26(5), pg. 747. Copyright 2018 by “Elsevier Science and Technology Journals”. Reprinted with permission.

Bioprinting-Adjacent Approaches:

In addition to bioprinting approaches, there are a number of tracheal construction strategies that rely on 3D printed plastic scaffolds that could be translated to bioprinting and whose results are applicable to the future expansion of bioprinting. The following section will briefly discusses these approaches and their relevance to bioprinting.

3D Printed Bellows Tracheal Scaffold:

Aiming to mimic the strength and flexibility of the trachea, Park et al. designed a bellows style graft.28 This design was based on native rabbit trachea, 5 mm in diameter with 1.5 mm convolutions every 3 mm along the graft. Using finite element modeling, they compared their bellows style geometry to a regular hollow cylinder. The bellows demonstrated increased resistance to compression while maintaining flexibility. This graft was produced by casting a blend of polycaprolactone (PCL) and gelatin in a 3D printed sacrificial mold. To promote the formation of cartilage, rings of heparinized gelatin sponges loaded with recombinant human transforming growth factor beta 1 (TGF-β1), were seeded with human nasal chondrocytes and manually placed around the graft. A schematic of this process is shown in Figure 4 (concept illustrated in Figure 8 C). These grafts were implanted in a nude mouse model. After 8 weeks they observed sulfated glycosaminoglycan production in the gelatin sponges, similar to native cartilage. Experimental mechanics data agreed with their computer modeling. Radial compression data of the explant were similar to the native trachea. Using a 3-point bend test, the explant and native trachea had a similar stiffness at low strain, but the native trachea became stiffer than the explant after 2 mm of displacement. While this approach uses traditional plastic printing to create a mold for tissue casting, the same graft design could be printed using a multi-head extrusion bioprinter, allowing the fabrication of branching structures. Innovations such as the “bellows” design highlight the potential of using novel geometries to satisfy the tissue function without attempting to fully recapitulate the native organ. Although, this approach has definite strengths, this kind of graft might face functional challenges long-term. As the PCL/gelatin scaffold degrades it would likely be replaced by scar tissue. The collagen dense tissue may surround and secure the cartilage rings, but it will likely be less flexible than native tissue. Following that approach for smaller airways, without cartilage tissue, the resulting fibrotic airway may be too stiff and present other challenges. One future approach to address this would be potentially seeding the bellows graft with pulmonary fibroblasts or another ECM secreting cell type that could produce a wider range of matrix proteins, such as elastin, in addition to the chondrocytes.

Figure 4:

An overview of the bellows graft development. A) 3D printed sacrificial mold, injected with PCL/gelatin, crosslinking and dissolving the outer mold. B) Treating and loading gelatin sponge with TFG-b3 and seeding with chondrocytes. Reprinted from “A novel tissue-engineered trachea with a mechanical behavior similar to native trachea,” by J.H. Park, J.M. Hong, Y.M. Ju, et. al, 2015, Biomaterials, 62, pg. 107. Copyright 2015 by “Elsevier Science and Technology Journals”. Reprinted with permission.

This same geometry was also used to examine epithelium growth for the tracheal lumen. In this approach, the graft was made from PCL alone, but lined with epithelium derived from human turbinate stromal cells (hTMSCs), harvested from the nasal cavity of patients undergoing septoplasties and partial turbinectomies.29 The hTMSCs were grown as sheets on temperature-responsive dishes. These sheets were removed from the culture dish and attached to the graft lumen using a thin layer of atelocollagen, a soluble form of collagen primarily used in the cosmetic industry. These tissue grafts, 5 mm wide by 15 mm long, were then implanted in the anterior tracheal wall of a New Zealand white rabbit model. After four weeks, grafts coated with hTMSC sheets exhibited mature columnar epithelium while grafts with atelocollagen alone showed a more immature epithelium lacking mature cilia. From the staining data presented in the paper, it is unclear if the hTMSCs became part of the columnar epithelium or simply helped accelerate native cell migration and maturation. This could be further examined using fluorescence or gene-marked cells. Using this approach in a larger defect area or in the lower airways would be of interest for further investigation.

3D Printed Polymer Tracheal Scaffolds:

Building on the use of autologous cartilage grafting during laryngotracheal reconstruction, Goldstein et al. designed a 3D printed patch from polylactic acid and type 1 collagen (concept illustrated in Figure 8 D).30 This patch was designed to be sutured directly into the wall of the trachea, providing additional tissue to expand the airway. The oblong patch scaffold, 8 × 3 × 1.2 mm, was printed out of polylactic acid (PLA) with multiple channels. Rabbit chondrocytes were harvested from a donor trachea, combined with the collagen solution and injected into the printed scaffold. The completed patches were then implanted into the anterior trachea of a New Zealand white rabbit model and sacrificed at 4, 8, at 12 weeks. After 12 weeks, histological analysis showed aggrecan production without significant inflammation or granulation. In translating this approach to clinical use, the authors suggest using auricular cartilage, but it may also be worthwhile to explore using mesenchymal stromal cells (MSCs) derived chondrocytes or another donor source site. Building on this approach, more complex 3D printed synthetic scaffolds could combine multiple cell types.

While still relying on 3D printed PCL rings for mechanical strength, Rehmani et al. combined these rings with decellularized dermal tissue (concept illustrated in Figure 8 E).31 Basing their graft on CT scans, they were able to produce personalized rings within 72 hours after the scan. These rings were then sutured to a hydrated sheet of decellularized bovine dermal collagen matrix. These grafts were implanted in a porcine model, replacing the front half the cervical trachea. The graft was monitored for three months using bronchoscopy before animals were euthanized for histological evaluation. Five of the 7 animals survived to the end point, but all of the animals exhibited granulation at the graft interface. The authors surmised this was an effect of the ECM being rapidly resorbed by the host’s cells, eliciting an immune response. They suggested seeding the bovine ECM with autologous MSCs and epithelium, relying on MSC immunomodulatory properties, and the epithelium as a barrier to pathogens. One potential challenge to this approach is its long-term stability. Over time, the PCL will degrade through hydrolysis and, without it, the bovine ECM tissue will likely collapse. To prevent this, MSC seeding could be combined with growth factor delivery to promote localized cartilage differentiation.

Townsend et al. also used a ring-supported mesh technique using 3D printed PCL rings embedded in an electrospun PCL patch (concept illustrated in Figure 8 F).32 To create this graft, they spun a base layer of isotropic PCL nanofibers on a mandrel matching the size of their target airway. The PCL rings were added to the mandrel and the spinning continued until the target thickness was reached. The PCL mesh was trimmed to be slightly larger than the tracheal defect in their animal model; a 1.5 × 2.5 cm elliptical opening on the anterior side of Rambouillet-Columbia sheep. The PCL rings extended past the defect site and were sutured into the surrounding tracheal wall. Trachea were retrieved 10 weeks after implantation due to respiratory concerns. Tissue had grown over the luminal side of the graft causing stenosis. While native epithelium grew in from the edges of the construct, it did not reach the center region. Overall, the 3D printed and electrospun PCL graft was able to create a liquid/air barrier in the ovine model. This approach has great potential as an off-the-shelf treatment for emergency surgery. However, since it has no cellular components, the long-term healing process will likely result in scar tissue formation in the defect leading to a weaker region in the tracheal wall. Therefore, the lifetime of the graft is limited by degradation of the PCL unless the graft can recruit surrounding cells. This could be attempted by including factors into the acellular scaffold or 3D printing biologically active matrix components.

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Alveoli:

Bioprinting the alveoli remains a technical challenge given their size scale and complexity. Current extrusion-based printers are limited by the shear forces applied on the cells as they pass through the needle, making their maximum resolution around 100 um.2

Stereolithography can produce high resolution constructs (100 um) but limits the composition complexity since most use only a single bath. Laser induced forward transfer (LIFT) (10 um) and some droplet-based printers (50 um) are capable of single cell precision with multiple cell types, but the current production speed is unreasonable for producing clinically relevant sized tissues and further incorporating them into the hierarchical structure of the lung would be a follow up challenge.33 In light of these resolution challenges, creating the necessary void space geometries and blood-air barrier is no easy task. There has been little progress on fully recapitulating the alveolar epithelial/endothelial barrier at the 3D alveolar scale. Most research in developing alveolar models has focused on recreating the air/cell/fluid interface. This area has grown significantly using ‘organ-on-a-chip’ approaches, as done by Huh et al. 2010 who created a breathing alveolar model using a poly(dimethylsiloxane) (PDMS) microdevice. Horváth et al. 2015 expanded on this approach by printing alternating layers of Matrigel® with endothelial and epithelial cells over a porous membrane.34 They showed an increased cell coverage and cell/cell interactions using printing compared to manual techniques. Focusing instead on renal function, Lin et al 2019 has been able to create a proximal tubule model using epithelium and endothelium lined vessels in a permeable ECM to examine renal reabsorption.35 This work demonstrated mass transport between confluent cellularized vessels made using a 3D printed sacrificial hydrogel. While these approaches have great potential for in vitro toxicology and pharmaceutical development, they do not translate easily to an implantable engineered tissue.

Working towards a more complex geometry, Grigoryan et al 2019, published a distal lung model with vascular and airway spaces (Figure 5).36 Using poly(ethylene glycol) diacrylate (PEGDA) and a stereolithographic printer, they created a “breathing model” with tidal air ventilation and blood flow. Using this model, they were able to demonstrate pulmonary transport by measuring blood oxygenation entering and leaving the model. While the printed vasculature was limited to 300 um, it is still a large step towards developing clinically relevant lung tissues. Other research has attempted to recapitulate the scale and morphology of the alveolus, focusing on the epithelial side. Lewis et al. used photodegradable microspheres to create hollow epithelial cysts on the scale of alveolar structures.37 These microspheres were made from poly(ethylene glycol) (PEG) crosslinked by a photolabile, nitrobenzyl ether crosslinker. This allowed the crosslinks to cleave when exposed to 365 nm light. The surface of these microspheres was seeded with alveolar epithelial cells and embedded in a second hydrogel. Degradation of the PEG hydrogel with light resulted in the alveolar cells being seeded on the walls of a 120 micron cyst. While these epithelial cysts were not a hierarchical structure, this method could be combined with previously discussed methods for printing airways.

Figure 5:

Distal lung model with red blood cell perfusion and air sac ventilation, scale bar 1mm. Reprinted from “Multivascular networks and functional intravascular topologies within biocompatible hydrogels,” by B. Grigoryan., S.J. Paulsen, D.C. Corbett, et. al, 2019, Science, 364, pg 461. Copyright 2019 by “The American Association for the Advancement of Science”. Adapted with permission.

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Esophagus:

The esophagus connects the pharynx to the stomach and is the first point along the gastrointestinal tract to exhibit peristalsis. It is a soft tube approximately 25 cm long in adult humans and has a wall thickness between 3 and 5 mm, depending on whether the smooth muscle in the esophageal wall is contracted.38 The lumen is lined with non-keratinizing stratified squamous epithelium, which is sparsely perforated by the orifices of esophageal glands that produce lubricating mucin. Below the epithelium lie the lamina propria and muscularis mucosa. All of these components combined form the mucosal layer. Deeper still lies the submucosa, which contains the submucosal plexus. This plexus is part of the enteric nervous system (ENS), and it controls secretory functions of the glands. The muscularis externa is a thick layer of muscle lying deep to the submucosa. The upper third of the muscularis is mostly skeletal muscle, with the middle third being a transition into smooth muscle. The lower third of the human esophagus has only smooth muscle. It is divided into an inner, circumferentially-oriented layer and an outer, longitudinally-oriented layer. Between these layers lies the myenteric plexus, the tissue of the enteric nervous system which coordinates peristalsis (Figure 6). The outermost layer of the esophagus is the adventitia, which is made primarily of fibroblasts and adipocytes. The large blood vessels and nerves of the esophagus run in this adventitia layer.

Figure 6:

Schematic of the structure of the esophageal wall. The innermost layers of the mucosa and submucosa are bordered by the neurons of the submucosal plexus. The next layer is the circular smooth muscle, which is separated from the outer longitudinal smooth muscle by the neurons of the myenteric plexus. Adventitia/serosa and vasculature not shown. Reprinted from “Development, Anatomy, and Physiology of the Esophagus,” by K. Staller, B. Kuo, 2013, In Principles of Deglutition, pg. 283. Copyright 2013 by “ Springer ”. Reprinted with permission

Many tissue-engineered esophageal solutions have been attempted, but so far very few of them have utilized 3D bioprinting technology. Those that have used 3D printers did so indirectly, creating acellular scaffolds out of PCL, then seeding with cells or implanting into the greater omentum for cellularization and vascularization. The first attempt was in 2016, when Park et al. used a melt extrusion 3D printer to lay down a grid structure of PCL (concept in Figure 8 G).39 The resulting scaffold was coated with a solution of fibrin, thrombin, and rabbit Mesenchymal Stem/Stromal Cells (rMSC). After coating, the scaffolds were sutured as an allogeneic implant into a 5×10mm surgical defect in the esophagus of New Zealand white rabbits. The test subjects were tube-fed for 7 days following the surgery to allow the healing process to begin before challenging the barrier and mechanical functions of the esophagus. Both cellular and acellular controls remained patent out to three weeks without leakage or infection. Scaffolds containing rMSCs were completely covered by epithelium and subepithelial tissue, whereas controls did not have coverage on the luminal side. This work is promising for use of PCL as a 3D-printed scaffold and for use of allogeneic MSCs for supporting regeneration. However, no evaluation of the source of regenerated cells was done. MSCs are known for their ability to differentiate into a variety of somatic cell types including fibroblasts and smooth muscle cells.40 Future experiments in fate tracking using reporter genes should be done to determine whether the MSCs are acting as paracrine signaling cells or are differentiating and proliferating to reconstitute the esophageal tissue. Staining with alpha-smooth muscle actin, vimentin, and desmin would allow for delineation of smooth muscle cells from fibroblasts and myofibroblasts.

More recently, Chung et al. used PCL to create a fully-circumferential esophageal replacement.41 The goal was to create a suitable acellular graft that could maintain an open lumen (concept illustrated in Figure 8 H). A 3D printer was used to melt-extrude PCL onto a rotating mandrel, creating multiple rings. PCL was then electrospun over the rings while still on the mandrel, resulting in a structure approximately 5 mm in length and 1.6mm in internal diameter. Evaluation by scanning electron microscopy (SEM) showed an average microfiber thickness of 2.5 microns and an average poresize of 5 microns. Mechanical testing showed that the ultimate tensile stress and the yield stress were significantly higher than that of a natural rat esophagus, and the elastic modulus appeared to be comparable, below 400% strain. These acellular scaffolds were cultured in a rat omentum for 2 weeks for in vivo cellularization and vascularization. Omental tissue was found to have adhered to the outside of the graft and to have obscured the lumen, requiring opening with a needle. The cellularized construct was then orthotopically implanted into a full-circumference surgical defect in the same rat. All rats were euthanized or died by two weeks. Rats that died were found to have luminal obstruction with hair and excrement due to rat grooming behavior and coprophagy. The authors noted that the non-contracting nature of the PCL graft with a small lumen was the most likely problem, as aperistaltic grafts in dogs with larger inner diameters had not resulted in such obstructive problems. Histological evaluation showed loosely-organized, vascularized tissue adherent to the inner and outer surfaces of the graft, but few cells appeared to have infiltrated the nanofibrous structure itself. This may be due in part to the fairly small pore sizes available to the cells, which are generally known to be about 10 microns in length. PCL is also known to be hydrophobic, with poor cellular adhesion if not surface-treated with methods such as plasma deposition or peptide immobilization.4243Some epithelial ingrowth from the proximal and distal anastomoses was noted, but its progression may have been halted by the inner supporting PCL rings. Overall, this approach shows promise as a method for vascularizing an acellular tissue via omental implantation. The inclusion of 3D-printed PCL rings to support the nanofibrous structure maybe more useful in large airway engineering, especially since it mimics the ring cartilage of the trachea.

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Stomach, Intestines, and Bile ducts:

The stomach and intestines follow a very similar radial pattern to that of the esophagus, with an inner mucosa and submucosa surrounded by nerve tissue of the ENS and muscle tissue of the muscularis externa. The stomach has an additional layer of muscle compared to the rest of the gastrointestinal tract, as the muscularis externa has an innermost oblique layer of smooth muscle. There are striking differences in the epithelium, with gastric epithelium designed to secrete acid and protect against acid-damage, and intestinal epithelium specialized to absorb nutrients and secrete waste. Compared to the largely mechanical role of the esophagus, the functional roles of these specialized epithelia present a significant challenge to the tissue engineering field.

At this time, 3D bioprinting techniques have not been applied to the stomach. However, a joint research team from Organovo and Merck have used a proprietary 3D bioprinting system to create a flat model of intestinal epithelium sitting on muscle.44 The goal of this work was to create better tissue models for drug and toxicology studies, and it demonstrated the ability to print cell-laden gels to recreate the complex structure of the intestines. Human intestinal myofibroblasts (hIMFs) and human biopsy-derived intestinal epithelial cells (hIECs) were suspended separately in proprietary bioinks. The fibroblast ink was printed as a layer onto a transwell membrane, then the epithelial ink was printed as a layer on top (concept illustrated in Figure 8 I). The bilayer was matured in culture for 10 days, then examined histologically and tested with multiple drug compounds. The resulting tissue demonstrated multiple epithelial subpopulations, including enterochromaffin cells and secretory goblet cells, both of which are vital to mucosal function. Immunolabeling of tight cellular junctions between the epithelial cells, and of cell membrane transporters demonstrated that the epithelium had an appropriate structure and polarization. While analysis of the materials and techniques used in this work is limited due to its proprietary nature, this work is promising as it demonstrates the possibility of 3D bioprinting with patient-derived cells. Replication of the laminar structure of the intestine and achievement of epithelialization without relying on inward growth from the anastomosis will allow for longer intestinal segments to be created. This work did not investigate any mechanical properties of the bilayer, however, and consideration of the muscular structure and peristaltic function must be considered in the next steps towards bioprinting an intestinal graft. Other hollow structures associated with the gastrointestinal tract include the extrahepatic bile ducts and gallbladder, which transport bile salts and acids produced in the liver to the intestinal lumen. The general tissue structure is an inner lumen lined with simple columnar epithelium, surrounded by layers of smooth muscle. Recently, Yan et al used a bioink composed of gelatin, self-assembled nanofibers, and cholangiocytes to demonstrate that bioprinted cholangiocytes were capable of self-organizing into branching tubular structures.45 Part of the novelty of this work was using peptide amphiphiles with a laminin-like amino acid sequence to make self-assembled nanofibers in the bioink. Laminin is known to be a key component of the basement membrane underlying epithelium, and inclusion of these nanofibers led to much better tubular development when compared to other bioinks. The 3D bioprinted structure was a flat, multilayered grid. Combined with the small tubular organization, the most immediate application may be for intrahepatic bile ductules in liver tissue engineering.

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Ureter, Urethra, and Bladder:

Hollow and tubular structures are integral to the function of the urinary system from the Bowman’s capsule and tubules found in the renal nephron down through the urethra. The scale of tubules found in the kidney is very small, leading to classification of the kidney as a solid organ. The ureter connects proximally to the kidney, and passively transports urine to the bladder in the pelvis. It is also capable of using peristaltic motion to shuttle the urine, due to the smooth muscle layer surrounding the epithelium.46 The epithelium of the urinary tract is significantly different than that seen in other regions of the body, as the cells take on different shapes depending on the degree of distension applied to the surface. As such, it is known as transitional epithelium, or urothelium. When relaxed, it appears as a stratified cuboidal epithelium, but takes on a stratified squamous appearance when stretched.47 Urothelium lines the ureters, bladder, and parts of the urethra.48 The upper portion of the ureter has two layers of smooth muscle, with the inner layer being longitudinally oriented and the outer being circumferential.49The distal ureter has a third, outermost muscular layer which is longitudinally oriented. The bladder also has three layers of smooth muscle, which are continuous with the layers of the distal ureters.

Tissue engineering in the urinary tract is one of the first and most successful examples of growing replacement organs, and many promising examples of successful tissue engineering strategies exist.50 3D bioprinting has not been extensively applied yet, but some recent publications show promise of the technique. In 2017, Zhang et al used a bioprinting system developed by Atala’s group at Wake Forest University to print a multi-layered urethra.51 This printing system is able to extrude polymer melts and bioinks. By blending PCL with poly(lactide-co-caprolactone) (PLCL), the group was able to capture the desirable high tensile failure properties of PCL and still replicate the modulus of the native rabbit urethra. The optimal polymer geometry for replicating the tissue mechanics was found to be spiraling in structure (Figure 7; Figure 8 J). Two fibrin bioinks were made, one with rabbit bladder smooth muscle cells (SMCs) and the other with rabbit bladder urothelial cells (UCs). The full tissue-engineered urethra was formed layer-by-layer, printing rings of PCL-PLCL blend, then inner rings of UC bioink and outer rings of SMC bioink. The result was a vertically-standing three-layered tube, with PCL-PLCL occupying the normal position of the lamina propria and providing mechanical support. The fibrin gel was crosslinked with thrombin solution and the constructs were maintained in static culture for 7 days. Both the SMCs and the UCs were found to proliferate, although the percent viability decreased over time. The authors attributed this to lack of vascularization and use of static culture instead of a dynamic bioreactor. Overall, the use of cells from a biopsy was successful in creating a 3D bioprinted urethra. Future studies in vascularization and in vivo function and remodeling will be necessary. This approach could also be applied to the ureters, since their general structure is similar to the urethra.

Figure 7:

3D bioprinting of a urethra. A-D) printing process, crosslinking, and immersion in culture medium. E-F) views of bioprinter before and during printing. Reprinted from “3D bioprinting of urethra with PCL/PLCL blend and dual autologous cells in fibrin hydrogel: An in vitro evaluation of biomimetic mechanical property and cell growth environment,” by K. Zhang, Q. Fu, J. Yoo, et. al, 2017, Acta Biomater., 50, pg. 158. Copyright 2017 by “Elsevier BV “. Reprinted with permission.

Another, more recent approach to creating tubular urothelial tissues used a coaxial extrusion technique.52 A bioink was made using a blend of gelatin methacrylate (GelMA), alginate, and poly(ethylene glycol) acrylate (PEGOA) with a tripentaerythritol core. This allowed optimization of fluid properties for printability, and multiple crosslinking options for mechanical tuning after printing. Alginate was used as a thickening agent in the bioink and was crosslinked by calcium ions to support later post-processing steps but was then de-crosslinked by chelation of the calcium with Ethylenediaminetetraacetic acid (EDTA). A coaxial needle with three concentric components made it possible to create hollow tubes having a wall with two concentric layers. To create a urethra, human bladder SMCs were loaded into the outer layer bioink and human UCs were loaded into the inner bioink. These tissues were cultured out to 14 days and showed proliferation and appropriate tight junction proteins in the developing urothelium. Printing tissues in this way is convenient for quickly creating hollow tubes with multiple layers, but limitations to the approach will need further exploration. For example, the inner diameter of these urethras was approximately 650 microns with a total wall thickness of roughly 150 microns. Considering that the normal human urethra diameter is closer to 8 mm inner diameter, it is not known if the structure will support itself or if it will require additional mechanical elements. Application to larger organs, such as the trachea, esophagus, and intestines will present an even greater challenge.

In 2018, Imamura et al. constructed bladder tissue by applying a similar needle-based technique previously used by the same group for tracheal engineering.53 Bone marrow mesenchymal cells (STRO-1 positive) from GFP-positive rats were isolated by adherent cell culture, and then formed into spheroids by gravity-induced accumulation in wells. The organoids were bioprinted onto an array of needles, to create a rectangular cuboid with dimensions of 3×3×1 mm. After maturation for 7 days in perfusion culture, the spheroids fused to form a solid tissue that was removed and implanted into a surgical defect in a radiation-injury bladder model in a rat. Four weeks after implantation, vascularization and innervation from the wound edges was seen growing into the tissue-engineered construct. Smooth muscle actin staining also indicated myogenic differentiation of the GFP-labeled donor cells. Compared to sham surgery controls, the treated rats were found to have significantly better bladder function in terms of micturition frequency and volumes. While this biofabrication approach was used to make a patch, it could theoretically be used to make a whole hollow bladder, as the needle array structure allows for free positioning of spheroids in space without need for printing supports.

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Summary and Future Perspectives:

The current work on 3D bioprinted hollow organs spans the pulmonary, digestive, and urinary systems with data obtained both in vitro and in vivo, summarized in Table 1. Most of these approaches rely on a polymer scaffold to provide the tissue geometry and mechanical stability, visually summarized in Figure 8. One of the core challenges of tissue engineering has been matching the degradation rate of the scaffold with the production of new matrix, determining the longevity and integration of the graft. A few approaches focus solely on cell aggregation and their products. Without using a synthetic graft, these approaches avoid degradation challenges and limitations. These wholly biological grafts may also have more active benefits, such as the potential to grow and adapt with patient. However, these become more challenging to validate and regulate. The current industry is suited for fully synthetic or acellular grafts that can be off-the-shelf products. 3D bioprinting lends itself to the rapid production of clinical scale and patient-specific tissues, which may be the technical shift needed to translate biological grafts. This shift will create new supply pipeline needs and require a new paradigm for clinical translation. During this transition, complex solid organs, such as the lungs, will depend on the development and translation of other hollow organs to pave the way into the clinic. In general, progress in bioprinting of muscular hollow organs has demonstrated gross preservation of cell-specific functions and an ability to create multi-layered tissues. So far, most of these implanted tissues have been built around biodegradable polymers such as PCL for mechanical stability. A notable exception is the scaffold-free bladder tissue created by Imamura, et al. Long-term studies will be necessary to investigate the mechanical stability of these tissues as the supporting polymers are degraded and cells proliferate and remodel the construct. Furthermore, preserving of the excitable, contractile phenotype of smooth muscle cells must be addressed. Creating contractile muscular walls and developing a strategy for innervation or other stimulation in a physiological context will move the field closer to clinically useful therapies.

Table 1:

Summary of Hollow Organ Tissue Bioprinting and Adjacent Methods

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The future of lung bioprinting will depend on overcoming several key challenges. The first challenge will be reliable cell sourcing for each layer and generation of the airways, from the ciliated epithelium to the alveolar epithelium. This may come from histocompatible, allogeneic donor populations or derived from host induced pluripotent stem cells (iPSCs). The next will be maintaining lumen patency at all scales of the lung, including the vasculature. Alveoli require creation of a hollow structure with single cell precision of multiple cell types, closely associated with convoluting microvessels, while maintaining a sealed air/fluid interface to avoid a gas embolism. Finally, the 3D bioprinted lung will need a dynamic bioreactor capable of providing fluid and gas perfusion while mechanically stimulating the tissue.54

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Acknowledgments:

We would like to acknowledge Del Reed from the University of Minnesota Bio-Medical Library for assistance with the literature search. This work was supported by the National Institutes of Health NHLBI F31 (5F31HL142313) and a CTSI Translational Research Development Program grant awarded to ZPG, and the T32 (5T32HL007741) training grant for CDV (Training in Pulmonary Science, PI: D. Ingbar). All authors have read the authorship agreement and have reviewed and approved this manuscript. The authors declare no conflict of interest for this work, having read the journal’s policy on conflicts of interest.

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Abbreviations:

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Footnotes

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