Abstract
Organoid, a 3D structure derived from various cell sources including progenitor and differentiated cells that self-organize through cell-cell and cell-matrix interactions to recapitulate the tissue/organ-specific architecture and function in vitro. The advancement of stem cell culture and the development of hydrogel-based extracellular matrices (ECM) have made it possible to derive self-assembled 3D tissue constructs like organoids. The ability to mimic the actual physiological conditions is the main advantage of organoids, reducing the excessive use of animal models and variability between animal models and humans. However, the complex microenvironment and complex cellular structure of organoids cannot be easily developed only using traditional cell biology. Therefore, several bioengineering approaches, including microfluidics, bioreactors, 3D bioprinting, and organoids-on-a-chip techniques, are extensively used to generate more physiologically relevant organoids. In this review, apart from organoid formation and self-assembly basics, the available bioengineering technologies are extensively discussed as solutions for traditional cell biology-oriented problems in organoid cultures. Also, the natural and synthetic hydrogel systems used in organoid cultures are discussed when necessary to highlight the significance of the stem cell microenvironment. The selected organoid models and their therapeutic applications in drug discovery and disease modeling are also presented.
Keywords: Organoid, Extracellular matrix, Microfluidics, Bioprinting, Self-assembly
1. Introduction
The three-dimensional (3D) cell culture in the form of cell aggregation was attained its first popularity in 1965–1985, according to the number of published materials [1,2]. Despite the name “organoids” was introduced to the scientific community half a century ago, this name was used to describe classical biological experiments that sought to define organogenesis by cell dissociation and reaggregation experiments [1]. Later, the emergence of the 3D cell cultures has fueled the knowledge on the effectiveness of the 3D cell cultures compared to the conventional two-dimensional (2D) cell cultures. The 2D cell cultures do not provide the conditions required for cellular organization and cell interactions in vivo. Also, the cell signaling networks are altered in 2D cell culture versus 3D cell cultures. In contrast to the 2D cell culture, 3D cell culture models exhibit very close properties to the in vivo conditions. The 3D culture models provide more realistic findings similar to in vivo conditions in translational research [3]. The 3D culture models are generally either two models: scaffold-based models or scaffold-free models. In scaffold-based 3D culture, the cells are grown in the substrate, which mimics the extracellular matrix (ECM) properties in either hydrogel or solid scaffolds; while cells are unable to attach to the substrate and force to make cell aggregates or cell spheroids in scaffold-free method [4,5]. The past decade again witnessed the organoid’s resurgence, yet in a somewhat different form due to the parallel development of 3D culture techniques. The breakthrough of the current organoid resurgence was made by Hans Clevers and his colleagues in 2007 after discovering a new type of stem cells [6]. Organoid technology is now on the verge of emerging as an independent research field.
The term organoid has lost accuracy in recent years and has been defined broadly according to the different cell culture techniques and their respective applications. However, this broad applications from small tissue explants to clonally expanding cells that self organize have made it difficult to define or otherwise, have made its meaning ambiguous. Recently, experts from different part of the world came together and defined the organoid as a “3D structure derived from (pluripotent) stem cells, progenitor, and/or differentiated cells that self-organize through cell-cell and cell-matrix interactions to recapitulate aspects of the native tissue architecture and function in vitro” [7]. However, according to the early definitons and studies of the organoid, it has been established that organoids are exclusively derived from the stem cells [1, 8–11]. It is now clear that organoid can generate from differentiated cells such as cholangiocytes [12,13]. Due to this recent expert meeting, previously believed [11] organoid definition and characteristics of the organoid was changed to a newer version and in this review we focused on the latest definition of the organoid. So, recently new catergorization for the organoid was proposed. Organoids cen be devided into three distict groups based on their characteristics; (i) epithelial organoids, (ii) multi-tissue organoids, and (iii) multi-organ organoids [7].
The first category, epithelial organoids are derived from a single germ layer (endoderm, mesoderm, or ectoderm). Epithelial organoid can self-renew under appropriate culture conditions and also represents the most widely studied organoid type. Since epithelial organoids do not contain mesodermal component, in some cases, supporting cells are cocultured with epithelial organoids [14,15]. Multi-tissue organoids at least have two germ layer cells or co-differentiation of PSCs, which are established through co-culture. Interestingly, current studies do not support the self-renewal of multi-tissue organoids in normal pathways, but cells interact to achieve a stable level of maturity and function [16, 17]. The most complex and the least studied organoid type is multi-organ organoids, where these organoid can make a significant contribution for study of organogenesis.
Organoids can be derived from different types of cells: (i) pluripotent embryonic stem cells (PESCs) and their synthetic counterparts, induced pluripotent stem cells (iPSCs); (ii) organ restricted adult stem cells (aSCs) [1]; and (iii) differentiated cell types such as cholangiocytes and hepatocytes, as described in the recent literature [18]. PESCs and iPSCs are simply known as pluripotent stem cells (PSCs) in most literature. Both these PSCs and aSCs have shown an unprecedented capacity to self-organize into structures that mimic the fundamental properties of the tissues which they are supposed to form. As mentioned above, human iPSCs (hiPSCs) play a significant role in the recent development of novel organoid models [10,19–21]. The discovery of the iPSCs has resulted in the development of multiple differentiation protocols in vitro using various endoderm-derived tissue types, including stomach [22], kidney [23], liver [24–26], lung [27–29], pancreas [30], and intestine [31]. The invention of directed differentiation of tissue-specific cell types from iPSCs has resulted in the investigation of different organoids, including the brain [32], kidney, liver, lung [33], pancreas, and stomach. A more detailed review of some organoid types is discussed in the latter part of this paper.
As previously mentioned above, unlike the traditional in vitro cell cultures, organoids have complex cellular composition and architecture, making them a physiologically complex model to study tissue development process, tissue homeostasis, and cellular functions and signaling pathways present in tissues in vitro [10,34]. Theoretically, the organoids made from human stem cells should mimic the actual human organ functions, especially conditions that do not replicate well in experimental animals [6,34]. However, most organoid models do not replicate the actual human organ models for various reasons, including the difficulty of generating full tissue components, difficulty controlling cell types, and complexity in cell-cell and cell-matrix interactions in these systems. The dependence on the limited animal-derived hydrogels, including Matrigel and collagen as an ECM, makes organoids unsuitable for expansion and downstream clinical applications due to their uncontrolled modifications and risk of pathogen and immunogen transfer [35]. Due to these reasons, despite the extensive use of organoids in basic and translational research, organoids are currently restricted to drug screening and initial cell replacement strategies [8].
In this review, we discuss the recent advancement and trends in organoid technology, including different organoid models and their significance to the development of current therapeutic studies. We critically evaluate the use of synthetic or natural polymer as an ECM for organoid formation. Spheroids as an intermediate stage for organoid development are also be addressed by emphasizing recent discoveries. We also highlight the key challenges and limitations of the current organoid technologies, which remain to be addressed in the future.
2. The formation of organoids: self-organization
The formation of an organoid from iPSCs or aSCs or differentiated cells is a controlled process, uses biochemical and physical cues for tissue development and homeostasis. The development of the human body from the zygote is a precisely controlled process, which uses stepwise differentiation and self-organization of the cells. Understanding the self-assembly of organoids solely depends on the classic developmental biology theories. The tissue patterning is essential to understand the self-organization of the organoids. There are two distinct approaches to examine tissue patterning in vitro and in vivo. During tissue morphogenesis, cells segregate into discrete domains in vivo [11]. This phenomenon was observed in the Drosophila wing disc, where the anterior-posterior boundary is established through mutually repressive interactions [36], and a similar process was observed in vertebrate embryonic development at the midbrain-hindbrain boundary [37]. The relative morphogenetic movements of the cells in vitro can be examined through the dissociation and reaggregation of tissues as the second approach of tissue patterning. This method was primarily applied in the development of vertebrate organs [38,39].
The organ’s self-assembly mainly occurs due to the segregation of cells with similar adhesive properties into domains that create a thermodynamically most stable pattern, known as Steinberg’s differential adhesion hypothesis. This phenomenon is also known as cell sorting out. The second mechanism which influences the self-assembly is proper spatially restricted fate decisions. These two mechanisms are schematically represented in Fig. 1(B) and Fig. 1(C). The organoid’s self-assembly is possible because of the growing movement away from the 2D culture, which results in the formation of tubules or ducts when the epithelial cells are embedded into ECM hydrogel [11], but not all the epithelial organoids form tubles when embedded in hydrogel. So, ECM hydrogel is a major driving force of self-assembly of the organoids, which is discussed further in a later section of this paper.
Fig. 1.
(A) Schematic representation of formation of organoids; (B) Cell sorting out describes the movement of cells into different domains. Different cell types (purple or green) sort themselves because of different adhesive properties conferred by their differential expression of distinct cell adhesion molecules (shown as brown or orange bars); (C) Spatially restricted cell-fate decisions also contribute to self-organization in vivo and in organoids. Progenitors (green) give rise to more differentiated progeny (purple), which, because of spatial constraints of the tissue and/or division orientation, are forced into a more superficial position that promotes their differentiation. These cells can sometimes further divide to give rise to more differentiated progeny (pink), which are further displaced [11].
Apart from the above-mentioned natural self-assembly mechanisms, growth factors and small molecules are used to control the self-renewal and differentiation of stem cells by manipulating cell signaling pathways in a tissue-specific manner [34]. The additional supplementation of signaling cues for cell differentiation and self-assembly also supports the ECM materials. The most commonly used ECM material, Matrigel, provides signaling cues through basement membrane ligands to support cell adhesion [40]. Since most of the organoids extensively rely on stem cells, the stem cell microenvironment or niche plays a significant role in self-assembly and directed cell differentiation. Niche functions as a physical anchor for cells, and also extrinsic and intrinsic biochemical signals are generated with the help of different growth factors. Altogether, the organoid formation can be programmed according to the niche, which provides a wide range of biochemical and biophysical signals, cell-cell interactions, and cell-ECM interactions. However, the dependent on cell-autonomous self-organization during organoid formation is not yet easily controlled, and hence more investigations are required to control the organoid formation [11,34,41].
3. Organoid systems
Organoid systems offer promising platforms for future therapeutic studies and stem cell applications focusing on tissue and organ regeneration. However, similar to most model systems, there is a significant difference between in vitro and in vivo models, which can be addressed through various bioengineering techniques. Bioengineering approaches can be applied to develop bottom-up synthetic organoid constructs, and facilitate controlled organoid formation through different cell cultures. The natural self-assembly of cells with the help of a suitable microenvironment has limitations due to the lack of biochemical and biophysical cues. The organoid’s niche components are established from either cell siganling as autocrine, paracrine, or juxtracrine, which always has limitations; or exogenously added to the system as ECM substrates, small molecules, or growth factors. Spatial and temporal control of the dynamic environment during the differentiation of the cells can not be controlled properly without the help of exogenous substances during complex organoid formation. There are two approaches to organoid engineering: (i) the cell biology approach modulates cell behavior and programs the self-assembly and cell differentiation; (ii) the bioengineering approach manipulates the microenvironment and the cell organization through different approaches, such as scaffold-free printing and 3D bioprinting. However, to accomplish the best organoid system, a combination of both cell biology and bioengineering approaches is required. Recent advances in biomaterials, hydrogel systems, cell-driven tissue assemblies, micro/nanotechnology, and bioprinting techniques have enabled the possibility of designing dynamic and self-assembled organoids.
3.1. Engineering the cell microenvironment (niche)
Engineering the niche, which is similar to humans or animals, would be a breakthrough for humankind. The complexity of the niche is a primary limitation in organoid system development. Niche consists of soluble and surface-bound cytokines, ECM-cell interaction, cell-cell interaction, mechanical forces, and physicochemical cues, such as oxygen and pH [42]. The cytokines can be delivered as growth factors incorporated with an ECM material, using different chemical modifications such as covalent bonding. The controlled delivery of these growth factors using biodegradable scaffolds or microfluidic devices provides spatially and temporally restricted delivery. Recent studies have also demonstrated that the immobilization of these cues to the ECM plays an important role in mediating their biological effects [42–44]. For example, dexamethasone, a corticosteroid, and a potent modulator (growth factor) for osteogenic differentiation have incorporated into poly(ethylene glycol) (PEG) based hydrogel system to release slowly with the degradation of PEG. At the same time, mesenchymal stem cells (MSCs) differentiated into bone tissue [45]. The manipulation of the ECM material is vital for the growth and differentiation of stem cells into their intended application.
ECM, the noncellular component presents in all tissues and organs, is one of the essential components of the cell microenvironment and provides not only essential structural support for the cellular constituents but also initiates crucial biochemical and biomechanical cues that are required for tissue morphogenesis, differentiation, and homeostasis [46]. ECM components such as collagen, laminin, and fibronectin provide the physical structure of the tissues and also influence cellular behavior by engaging with integrin receptors [44]. ECM-cell interaction can be stimulated through several approaches, including synthesizing novel biomaterials relevant to specific stem cells [47], producing 3D scaffolds with micro or nanoscale topography [48], 2D micropatterning, and high-throughput microarrays [49]. The decellularized matrices also can be used as ECM components. For example, after the confluency in vitro, the bone marrow stromal cell matrix can be decellularized without disrupting the underlying ECM matrix and subsequently used to mimic endosteal and vascular niches with the help of relevant growth factors [34,50].
The most commonly used ECM component of organoid culture is Matrigel. Matrigel, a heterogeneous, complex mixture of ECM proteins including laminin, collagen IV, and enactin, is secreted by Engelbreth-Holm-swarm mouse sarcoma cells [51,52]. The exclusive dependence on this animal-derived hydrogel is complicated by current organoid methods due to the lot-to-lot compositional and structural variability, pose a risk of immunogen and pathogen transfer, and, most importantly, this tumor-derived matrix has limited clinical translational potential [35,52,53]. Moreover, the heterogeneous nature of the Matrigel does not easily allow for manipulating the morphogenetic processes, which are tightly governed in vivo by spatiotemporal cues [34]. However, Matrigel [22,54] and collagen [55–57], are extensively used in organoid systems. As an alternative to the animal-derived ECM, synthetic hydrogels can be tailored to produce the essential signals by incorporating specific biomimetic cues, such as glycosaminoglycans, a major ECM component believed to play an essential role in the cell niche. For example, hyaluronic acid (HA), a glycosaminoglycan highly expressed in bone marrow stromal cells, has played an essential role in regulating hematopoiesis in the hematopoietic stem cell niche modulating neural stem cell niche [44]. The development of synthetic polymer-based ECM, which facilitates cellular infiltration and adequate nutrient transport [58], is still at the primitive stage despite some recent advancements in synthetic hydrogels [34].
The recent advances in chemical methods, such as click chemistry, Michael type addition reactions, and gelation mechanisms, provide more controlled physicochemical properties in synthetic polymers [59]. Also, the bioinert environment of the synthetic polymers can be modified into a biomimetic environment similar to the natural ECM by adding biomimetic cues/signaling proteins through chemical or enzymatic cross-linking. Recently, the PEG-based hydrogel system has been studied for different organoid systems as a biomimetic ECM material. A PEG-based hydrogel, maleimide crosslinked four-arm PEG macromer (PEG-4MAL), has shown superior cytotoxicity and minimal toxicity and inflammation [60,61]. The PEG-4MAL hydrogel system enabled the successful growth of human intestinal organoids (HIO), which normally relies on undefined tumor-derived ECM. PEG-4MAL macromers were functionalized with RGD (Arg-Gly-Asp) adhesive peptides and crosslinked in the presence of hPSCs to generate HIO. The viability and mechanical stability were evaluated using Matrigel-derived HIO as a control model, and different concentrations of PEG-4MAL (3.5–6.0% w/v) were studied to understand polymer density-dependent effects on HIO survival. The viability of 3.5% and 4% PEG-4MAL based HIOs were similar to the Matrigel-based control group, while 5% and 6% PEG-4MAL groups showed lower HIO viability, suggesting the polymer density-dependent effect on survival of HIO as shown in Fig. 2 [62,53]. The ECM hydrogel should provide mechanical stability to the organoid system and the biochemical signals required for cell survival and proliferation, and the stiffness of the polymer is also important to the viability of the organoid structure, as suggested by the above study.
Fig. 2.
(A) (i) Illustration of the designed 384 hanging drop spheroid culture array plate; (ii) Photo and key dimensions of the array plate. (iii) Cartoon of the hanging drop formation process in the array plate. Within hours, individual cells start to aggregate and eventually form into a single spheroid around 1 day; (iv) Photo of the 384 hanging drop array plate operated with liquid handling robot capable of simultaneously pipetting 96 cell culture sites; (e) Cartoon of the final humidification chamber used to culture 3D spheroids in the hanging drop array plate [62]; (B) Synthesis of the PEG-4MAL hydrogel and organoid generation; (i) Floating, hPSC-derived human spheroids are collected and mixed with a solution of functionalized PEG-4MAL macromer. (ii) PEG-4MAL hydrogel is cast by pipetting the mixture of functionalized PEG-4MAL and spheroids into the cross-linker solution. Encapsulated spheroids expand and develop into human organoids [53].
3.2. Engineering organoids
Engineering organoids mainly focus on developing a suitable microenvironment for cells and overcoming the challenges aligned with the complexity of the cell niche. These challenges can be overcome by advancing material chemistry, such as easy and less cytotoxic click chemistry and Michael type addition reactions for novel hydrogel synthesis, and developing the bioengineering techniques that generate the stem cell niche by incorporating bioreactors, microarray systems, and microfluidics. The latter part of this, the bioengineering techniques will be further discussed, including the different techniques such as bioreactors, microfluidics, bioprinting, vascularization, and organoid-on-a-chip models. The summary of each bioengineering technique is shown in Table 1. Most organoid system’s main building block is cell spheroids; therefore, engineering 3D cell spheroids is discussed further, mainly focusing on the material perspective.
Table 1.
Summary of the main Microengineered systems used for organoid formation with the type of application, cell types, and applications.
| Micro engineered system | Type of organoid | Cell types | Hydrogel type | Applications |
|---|---|---|---|---|
| Bioreactors | Brain | Human iPSCs | Matrigel | Studying Zika virus-induced microcephaly [19] Exploring development of forebrain, midbrain, and hypothalamus organoids [87] Providing largest molecular map of the diversity of cell types that generated in the brain organoid [148] |
| Retinal | Human iPSCs | Generating retinal organoids that contains retina-like architecture with enhanced generation of photoreceptors [85] | ||
| Mouse PSCs | Developing a RWV bioreactor to promote the growth of retinal organoid that is closely recapitulating the spatiotemporal development of mouse retina in vivo [83] | |||
| Kidney | Human iPSCs | Exploring inexpensive and highly efficient method to grow kidney organoid that contains podocytes, proximal and distal tubles segments in developed nephrons [149] | ||
| Hepatic/ liver | Human PSCs | Exploring the impact of dissolved oxygen concentration on hPSCs hepatic differentiation fate and differentiation efficacy [86] | ||
| Adult upcyte® cells | Generating liver organoids from adult differentiated cells and studying long-term liver function [150] | |||
| Microfluidics | Gastric/ Intestinal | Gastrointestinal cells | Alginate/ Gelatin | Exploring feasibility and biocompatibility of nonfluorinated oils for application in gastrointestinal cells to study intercellular chemical interaction [96] |
| Brain | Human NESCs/ hiPSCs | Modelling a millifluidic culture system that supports continuous media supply by laminar flow which reduces the “dead core” of the brain organoid [151] | ||
| Liver | HepG2/ NIH3T3 | Collagen | Developing early stage drug screening system for hepatotoxicity [152] | |
| Induced hepatic cells | Decellularized porcine liver tissue | Developing a highly effective, functional 3D hepatic culture system that is suitable as a drug screening platform using a decellularized LEM hydrogel [91] | ||
| Human iPSCs | Fibrin/ chitosan/ Alginate | Synthesizing composite hydrogel capsules that enable the engineering of liver organoids in oil-free droplet microfluidic system [95] | ||
| Lung cancer | Patient-derived small-cell lung cancer | Matrigel | Exploring the drug responses under physiologically relevant flow conditions using lung cancer organoid culture platform [93] | |
| Bioprinting | Vascularized ossification | Human UVECs | Alginate | Developing a novel airflow-assisted bioprinting method to produce spiral-based spheroids with excellent resolution and complex microenvironment [153] |
| Colon | aSCs | GelMA | Introducing bovine colon organoid for agri-biotechnological applications [125] | |
| Mammary gland | Epithelial/cancer cells | Human mammary ECM | Exploring a novel mammary ECM hydrogel in 3D bioprinting to develop sustained and prolonged organoid formation and differentiation [126] | |
| Mamary epithelial | hMECs | Collagen | Exploring a 3D bioprinting platform to control the 3D formation of organoids through the “self-assembly” of hMECs [127] | |
| Salivary glands | hDPSCs | Validating a magnetic 3D bioprinting system to generate innervated secretory epithelial organoids [154] | ||
| Brain (cancer derived) | Human glioblastoma cells | PEGDA BPADMA | Developing a 4D printed programmable cell-culture arrays using thermo-responsive shape memory polymer for cancer drug testing [155] | |
| Organoid-on-a-chip | Brain | hiPSCs | Matrigel | Investigating the adverse effect of nicotine on cortical development [138] |
| Stomach/ gastric | hPSCs | Developing a gastric organoid that reminiscent of gastric motility by introducing a rhythmic stretch and contraction through a luminal flow [145] | ||
| Kidney | hPSCs/Metanephric mesenchyme cells | Exploring the flow generated microenvironment as a pivotal contributor to the structural and functional development of kidney organoids [146] | ||
| Intestine | Collagen | Introducing a new system to develop the crypt and villus architecture that is distinct to intestine epithelium using a diffusion gradient [147] | ||
| Cortical | hiPSCs | Matrigel | Investigating the autism-like neurodevelopment disorder due to the prenatal antiepileptic drug valproic acid (VPA) exposure [156] | |
| Liver | hPSCs | Investigating the improved cell viability and higher expression of endodermal genes and mature hepatic genes under perfused culture conditions to investigate hepatotoxic response after exposure to acetaminophen [157] | ||
| hiPSCs | Developing a nonalcoholic fatty liver model to explore the underlying conditions of steatosis and study the effective therapies [158] |
3.2.1. Engineering 3D cell spheroids
Most currently available organoid development protocols utilize the encapsulation of 3D spheroids, using biologically active materials, including natural and synthetic polymers [53]. Therefore, we introduce the 3D cell spheroids and their fabrication techniques with the recently published studies. 3D spheroid is also a 3D cell culturing method, as explained in the introduction section. In general, cells are cultured as aggregates in 3D culturing methods. The spheroid culture methods are among the most well-investigated areas due to their simplicity, reproducibility, and ability to make tissues similar to physiological tissues [62]. The 3D spheroids also can be considered as an organoid if they contain organ-specific cell types and satisfy the definition of an organoid [7].
Cells cannot attach to the biomaterial-based ECM surface, therefore, in spheroid formation, cells are forced to attach and form aggregates [62,63]. Numerous fabrication techniques have been demonstrated successful spheroid formation, including hanging-drop cultures [64–66], microfluidic devices [67], liquid overlay/cell suspension system [68], microwell arrays [69–71], and cultures on low-adhesive substrates [72,73]. From the methods mentioned above, hanging drop cultures, microwell arrays, and microfluidic systems are able to generate highly controlled aggregates with uniform size [74].
The development of robot-assisted hanging drop culture systems allows the production of large-scale spheroids in a single array at a low cost. The robot-assisted hanging drop system enabled the production of 384 spheroids in a single array, as shown in Fig. 2(A) [62]. The spheroid size is more critical for cell viability because oxygen and nutrients are transferred to cells using diffusion-driven mechanisms. Therefore, the generation of controlled-sized spheroids, which do not exceed a few hundred micrometers, is vital to avoid necrotic damage to the spheroid’s core. Moreover, these approaches enable incorporating the cells into the biomimetic hydrogel systems while maintaining the nutrient supply and interspheroid metabolic communication [63]. Due to the 3D structure, spheroids offer a wide variety of improved biological properties compared to the 2D single-cell culture systems, including enhanced cell viability, increased proliferation, differentiation, stable metabolic function, and stable morphology. Therefore, spheroids models are extensively used in cancer research as a biomimetic in vitro model, and also, they are extensively used in drug discovery and toxicology screening. 3D spheroid systems are also gained popularity due to the enhanced differentiation capability of multipotent MSCs into osteogenic, chondrogenic, neurogenic, adipogenic, and hepatogenic lineage compared to the 2D culture models. The insufficient oxygen diffusion to the core of spheroids triggers the upregulation of hypoxia-induced survival factors, such as hypoxia-inducible factor (HIF)-1α. These types of reactions result in increased secretion of angiogenic factors. Also, the spheroid diameter >100 μm exhibits more prominent upregulation of HIF-1α and vascular endothelial growth factor (VEGF) secretion and thus, support the vascularization. However, as previously mentioned, too low oxygen levels in larger spheroids may significantly compromise the cell viability, and therefore, it is required to determine the ideal size of spheroids for angiogenesis.
The current organoid formation methods mostly use 3D spheroids as an intermediate building box to form the final organoid. The progress of 3D spheroid technologies enables new paths for organoid engineering as it allows us to generate more complex cell aggregates, including rods, tori, and honeycombs, with a combination of hydrogels. However, these areas must be exploited further. Still, in organoid systems, Matrigel-based cell spheroids are widely used due to their excellent biomimetic capabilities in vitro. Matrigel-based hiPSC spheroids were used in HIO formation [54] and gastric organoid formation [22]. Also, a synthetic hydrogel, PEG-4MAL, was used to make hPSC-derived spheroid to form organoids after crosslink with a functionalized peptide, as shown in Fig. 2(B) [53]
3.2.2. Bioreactors
The bioreactor is a dynamic 3D cell culture platform that can be used in various distinct organoid-related applications, including the expansion of diverse cell types such as PSCs, generation of 3D organoids from PSCs, and tissue-specific differentiation of stem cell-derived organoids [75,76]. The dynamic nature of the medium around the stem cells in the bioreactor system facilitates organoids’ maintenance and differentiation. The flow around the cells further improves medium exchange, facilitating the higher mass transfer and more physiologically relevant gradients of nutrients and waste products between the surface and the core [77,78]. The gradient acts similar to the natural developmental environment during cell differentiation, enabling the correct orientation and lamination of different cell phenotypes [34,79]. The induced mechanotransducive responses from cells have resulted from the shear generated by the flowing medium, enabling better structural and transcriptional outcomes [80]. Different bioreactors have been used extensively in organoid, and regenerative and tissue engineering approaches, including rotating wall vessel (RWV), stirred-tank bioreactors, spin bioreactors, and more described elsewhere [81,82]. The most recent studies of bioreactor-based organoids generating systems are summarized in this review.
DiStefano et al., reported RMV bioreactor to develop retinal organoid using mouse PSCs. The efficacy of retinal differentiation in RMV-cultured organoids was higher than static suspension-cultured organoids, as demonstrated by immunohistochemistry and transcriptome profiling. This system recapitulated the spatiotemporal development of mouse retina in vivo, but further differentiation of retinal organoids in vitro was impaired and required additional functional maturation conditions [83]. Phelan and the group developed a system capable of constantly removing air bubbles from the conventional RWV reactors without interfering with the fluid dynamics enabling optimum cell culture conditions. The air bubble formation of the conventional RWV reactors during operation negates the critical features of the RMV environment, such as low shear, zero headspace, and simulated micro-gravity. This modified bubble capturing bioreactor showed an enhanced organoid formation in alginate beads and spheroids containing A549 human lung adenocarcinoma cells [84].
A stirred-tank bioreactor was used to develop the retinal organoids to address the visual impairment caused by inherited retinal degeneration, retinal diseases, age-related macular degeneration, and glaucoma. The developed organoids using hESCs and hiPSCs showed improved laminar stratification and yield of photoreceptor cells bearing cilia and nascent outer-segment-like structure [85]. In another study, Farzaneh et al., reported the 20–40% of dissolved oxygen (DO) concentration significantly influences the differentiation of hPSC aggregates by enhancing the mesoderm induction. The resulted human fetal-like hepatic organoids contained red blood cells and functional hepatocytes, exhibiting improved liver-specific gene expression, key metabolic functions, and higher cytochrome P450 activity. Further, the implanted hepatic organoids in an acute liver injury mouse model showed albumin production after implanting, suggesting the effect of DO in the liver tissues [86]. Qian et al., developed a cost-effective, simple-to-use miniaturized multiwall spinning bioreactor to generate forebrain-specific organoids from hiPSCs for model ZIKA Virus (ZIKV) exposure. These generated organoids recapitulate human cortical development’s key features, including progenitor zone organization, neurogenesis, gene expression, and notably, a distinct human-specific outer radial glia cell layer. Also, the team has developed protocols for midbrain and hypothalamic organoids [87].
3.2.3. Microfluidics
Nearly two decades ago, the word “microfluidics” was introduced to the biology and cellular engineering research field as a technology that can significantly change the perception of modern biology. Microfluidics is characterized as the science and technology of systems that manipulate a small amount of fluid at the submillimeter length scale [88,89]. The field of microfluidics has evolved from four different areas: (i) molecular analysis; (ii) biodefense; (iii) molecular biology; and (iv) microelectronics. (i) Molecular analysis including high-pressure liquid chromatography (HPLC), gas-pressure chromatography (GPC), uses a minimal amount of samples to achieve high sensitivity and high resolution and, therefore, there was a necessity to develop a more versatile and more compact microscale method in chemistry and biochemistry. (ii) The development of the field-deployable microfluidic systems for detecting biological and chemical threats in 1990 by the United States department of defense paved the stimulus for the rapid growth of microfluidic technology. (iii) After the explosion of genomics in the 1980s, the molecular biology field required analytical methods with higher sensitivity and higher throughput and resolution. (iv) The most influential contribution was coming from microelectronics. Photolithography and related technologies that had been first developed for the semiconductor industry and micro-electromechanical systems (MEMS) are directly applicable for microfluidic systems. Indeed, the earlier microfluidic devices were made using silicon and glass but later replaced with plastics [88]. In recent years, the field of microfluidics has under-gone rapid growth and has become an integral part of many biological and biomedical engineering applications. The simultaneous development of materials and microfluidic platforms is vital to meet the requirement of complex microenvironments [90]. Especially, organoid cultures require a dynamic culture environment to replicate the actual biological conditions in vivo. The microfluidic systems provide dynamic culture conditions with the help of automation by supplying continuous inflow and outflow of medium and nutrition through microchannels and further facilitate the long-term expansion of organoids and uniformity. The microfluidic systems extensively used in drug screening research and integrating microfluidic components into a single chip led to the advent of lab-on-a-chip devices. The recent studies of microfluidic-based organoids generating systems will be summarized in this review, while the more recent organoids-on-a-chip model will be discussed in the later section.
Jin et al., reported a 3D microgel liver organoid as a drug screening platform that showed higher sensitivity and reduced hepatotoxic response variation. Liver extracellular matrix (LEM) hydrogel was synthesized using decellularized porcine liver, and induced hepatic (iHep) cells were differentiated from primary mouse embryonic fibroblasts. The encapsulated iHep cells cocultured with the human umbilical vein endothelial cells (hUVECs) in LEM hydrogel (or 3D cell culture) in a dynamic state (continuous flow) using a simple rocker system showed reduced apoptosis and improved cell-cell adhesion in 3D iHep tissue. These organoids also showed enhanced hepatocyte-specific functions, including urea secretion and serum albumin secretion [91]. Most current organoid protocols involve several steps of cell seeding, differentiation, and maturation before evaluating the drug responses. Therefore, recently several studies were carried on to study the one-stop organoid formation for drug screening. Ao et al., reported a one-stop brain organoid formation method using a microfluidic platform containing perfusable chambers made out of poly (dimethylsiloxane) (PDMS) and poly(tetrafluoroethylene) (PTFE) coated steel mesh, as shown in Fig. 3A. An in situ air-liquid interface was used to reduce the hypoxic core. After day 10, the top chamber medium was removed to facilitate the air-liquid interface culture condition, as shown in Fig. 3A. This microfluidic system was used to analyze the effect of prenatal cannabis exposure on brain development [92]. Jung et al., and the team demonstrated a one-stop microfluidic platform that can reproduce 3D lung cancer organoids in size controllable manner. The lung cancer organoids derived from the patient-specific small-cell lung cancer (SCLC) were demonstrated to preserve SCLC’s morphological and genetic characteristics during the culture period. Also, this microphysiological system can provide insight into the efficacy of the anti-cancer drugs and can further study the mechanisms of possible chemotherapy resistance, which is common in SCLC [93].
Fig. 3.
(A) Schematics of brain organoid fabrication and cannabis exposure within the microfluidic device [92]. (B) Printing process of particular tissue constructs with dECM bioink; hHeart tissue construct was printed with only heart dECM (hdECM). Cartilage and adipose tissues were printed with cartilage dECM (cdECM) and adipose dECM (adECM), respectively, and in combination with PCL framework (scale bar, 5 mm) [119]. (C) Bioprinting of aortic valve conduit (i) Aortic valve model reconstructed from micro-CT images. The root and leaflet regions were identified with intensity thresholds and rendered separately into 3D geometries into STL format (green color indicates valve root and red color indicates valve leaflets); (ii) schematic illustration of the bioprinting process with dual cell types and dual syringes; (iii) fluorescent image of first printed two layers of aortic valve conduit; SMC for valve root were labeled by cell tracker green and VIC for valve leaflet were labeled by cell tracker red. (iv) as-printed aortic valve conduit [122].
In another study, Schuster et al., developed an automated, high-throughput 3D gel-based microfluidic platform capable of culturing organoids for continuous monitoring of growth, morphology, and biochemical analysis. This system consists of two integrated devices; a 3D culture chamber and multiplex fluid control device, automated and programmable experimental control with custom-made software, and live-cell time-lapse fluorescence microscopy. The patient-derived tumor organoids developed in this work showed significant differences in individual patient’s drug responses, suggesting the importance of studying drug response at the individual level for cancer patients [94]. Apart from long-term organoid cultures using microfluidic devices, microbeads or composite hydrogel-based microbeads were recently synthesized using microfluidic devices while culturing conventional culture plates. Wang et al., reported a one-step synthesis of composite hydrogel capsules using fibrin hydrogel as the core and alginate-chitosan composite shell. These microbeads were differentiated into liver organoids with encapsulated hiPSCs and human liver cell lines (HepaRG). The formed liver organoids showed urea secretion and albumin secretion, suggesting a proper liver function in vitro [95]. Pajoumshariati et al. used a drop-based microfluidic system to control the size and shape of the droplets using alginate-Ca-EDTA complex with gelatin hydrogel. This microencapsulation system was used to encapsulate crypt and Peyer’s patch immune cells to model the gastrointestinal niche [96].
3.2.4. Bioprinting
Bioprinting is a unique tool, which can be used to pattern the living cell incorporated into biomimetic hydrogel matrices and other biophysical cues. The computer-aided patterning of the complex structures allows the formation of organoids with controlled cell-cell and cell-matrix interactions. The advancement of the 3D bioprinters creates a new era of printing, which can print different cell-encapsulated hydrogels according to their mechanical and biochemical requirements. These complex 3D co-cultures of customized geometries enable bioprinting as a tool for complex organoid formation, which has not yet been appropriately investigated. The imaging data via magnetic resonance imaging (MRI) or computed tomography (CT) can be directly used to develop the complex geometries of the desired tissue or organ using 3D modeling software [97]. Then, those generated models can be precisely positioned in an additive layer-by-layer approach to create 3D tissue or organs using an inkjet or extrusion-based printing. The current printing technology enables to generate a plethora of complex geometries that attempts to resemble native tissues; thus, their biological structure that mimics the structure and function of their native tissue or organ depends on the hydrogel-based bioinks, which contain the nutrients and biologically active sites (ligands, growth factors). Several types of natural and synthetic hydrogel materials have been extensively investigated in the last decade, and many numbers of reviews can be found elsewhere [59,98,99]. In this review, we are not going to discuss the hydrogel properties and their formulation since it is out of the scope of this paper. However, some of the significant studies related to 3D bioprintable hydrogels are discussed here.
3D bioprinting extensively depends on several natural hydrogel systems, including alginate [100–102], collagen [103–105], chitosan [106,107], gelatin [108–110], hyaluronic acid [108,111], as well as synthetic hydrogel systems, such as gelatin methacrylate (GelMA) [112–115], and PEG-based materials [116–118]. The use of decellularized matrices has attracted increasing attention due to their ability to apply materials from the same tissue of interest, which provides well-matched compositional complexity and architectural fidelity between printed biological structure and the target tissue [119–121]. Falguni et al., were able to develop decellularized ECM (dECM) bioinks from three different tissues, namely cartilage, heart, and adipose. The lyophilized ECM was subjected to gelation using acetic acid and pepsin, and the cell encapsulated gel was printed, as shown in Fig. 3(A). Also, polycaprolactone (PCL) was used to support the gel structure where the printing was done with dECM gel and PCL in an alternative layer design only with adipose and cartilage dECM [119]. GelMA and collagen were recently used as a hydrogel while powdered dECM and cells were encapsulated with the gel for liver and heart tissue regeneration study using a digital micromirror device (DMD) [121]. Duan et al., bioprinted an aortic valve conduit with encapsulation of sinus smooth muscle cells (SMCs) in the valve root and human aortic vascular interstitial cells (hAVICs) in the leaflet as shown in Fig. 3(B). The cells were encapsulated in alginate/gelatin hydrogel, and after 7 days, hydrogel showed 83.2% and 81.4% viability for hAVICs and SMCs, respectively [122]. The aortic valve was 3D modeled using the micro CT scan of the freshly harvest porcine aortic valve (Fig. 3B(i)). Also, using a 3D bioprinting approach, it is possible to recapitulate the complex trabecular structure of a whole organ. Hinton et al., were able to print a day five embryonic chick heart using a 3D generated optical image from confocal microscopy. The printed heart was then imaged using a multiphoton microscope to generate a cross-section of the structure, which showed the internal trabeculation similar to the actual chick heart [123]. However, the generation of functional heart from the 3D printed heart was not studied in this study, yet showed the ability to print complex structures using 3D printing. In another study, Schoneberg et al., were able to develop an in vitro blood vessel model using a drop-on-demand bioprinting technique. This vessel model consists of continuous endothelium intimating the tunica intima, an elastic smooth muscle layer, mimicking tunica media, and the surrounding fibrous and collagenous matrix of fibroblasts mimicking the tunica adventitia. The necessary biochemical and biomechanical functionality were provided through the series of natural hydrogel matrixes, including gelatin, collagen, and custom-made bioreactor to mimic the ECM microenvironment [124].
Few recent studies directly focused on organoid printing can be found in the literature. However, most of the studies demonstrated simple bioprinting rather than complex organ printing. Töper et al., demonstrated the 3D bioprinting of colonoid-colon organoids - encapsulated GelMA hydrogel. The crypts were isolated from bovine colon tissue, and the crypts were seeded into wells with Matrigel. GelMA supported the colonoid culture, and the encapsulated colonoid was viable and proliferative, while GelMA maintained structural integrity during the culture period [125]. In another two studies, mammary organoids were generated using simple cell bioprinting on top of the hydrogel layer. A decellularized mammary tissue was used as an ECM hydrogel in one study, and the common cancer cell line and mammary epithelial cells were bioprinted on the hydrogel [126]. In the second study, human mammary epithelial cells were printed on the Collagen type I hydrogel matrix [127]. Alternatively, 3D cell spheroids can also be used in 3D bioprinting [128]. The use of liver spheroids can protect the cells from shear stress exerted on the individual hepatocytes during the printing process.
Despite the current progress in 3D bioprinting, several challenges are still remaining, such as the ability to print in high resolution, the ability to make the vascular structure to support long-term cell viability, and controllable cell densities. The development of new hydrogel systems parallel to the printing technologies will provide the possibility to build organoids with controlled stem cell niches and spatiotemporal ECM with multiple tissues.
3.2.5. Vascularization
In real organs, hierarchically branched vascular networks are available around the organ, and therefore, all the cells are within a few hundred microns away from the capillaries, which facilitate sufficient oxygen and nutrients through diffusion. The formation of vascular networks in the larger organoid system is still a major challenge. The one strategy is a scaffold-based system, where the vascular network is generated through a 3D modeling process, such as bioprinting. The most investigated such system is the coaxial nozzle method, where hollow tubular type structures are printed using biomimetic, cell encapsulated hydrogel systems. Also, the advancement of the microfluidic systems allows the distribution of flow and mass transfer more consistently, supporting the multiple types of vascular cells in vitro to form vascular tissues.
Khademhosseini and the group investigated agarose, a polysaccharide, as a sacrificial material to bioprint hollow microchannels within a hydrogel construct. The cell-laden hydrogel (GelMA) contained SMCs and fibroblasts. This photocurable material later photocured into a stable structure, and agarose fibers were removed using a mild vacuum. These microchannels were further coated with endothelial cells to promote angiogenesis (Fig. 4(A and B)) [129]. However, this direct retraction of the sacrificial template compromised the structural integrity, and therefore, another alternative bioink, Pluronic 127, which can simply remove reducing the temperature to 4°C, was introduced [130]. Alternatively, the coaxial nozzle strategy or core/shell model has been developed by Ozbolt and the team to print vessel-like microchannels in the cell-laden structure [131]. This system is mainly used in a calcium alginate hydrogel system, where Ca2+ acts as a crosslinking agent for alginate. This system consists of a coaxial nozzle, where Ca2+ solution flows through the inner core nozzle and alginate hydrogel flow through the outer core. When the alginate crosslinks with Ca2+, it makes a hollow tubular structure Fig. 4(C and D). The vascular endothelial cells were incorporated into the Ca2+ solution, while SMCs were incorporated into the alginate solution to promote vascular tissue formation. This method successfully made microchannels in vitro in numerous recent studies, but the use of this technique in organoid formation has yet to be discovered further. Another printing technique, scaffold-free bioprinting, is also utilized to make vascular structures in vitro. Norotte et al., investigated the scaffold-free vascular construct using layer-by-layer printing of agarose rods as a non-adhering molding template. Human skin fibroblasts (HSF) and human umbilical vein smooth muscle cells (hUVSMCs) were aggregated into discreet units either as multicellular spheroids or as cylinders of controlled diameter and printed with agarose rods in an alternative pattern as shown in Fig. 4(E) [132]. However, all these strategies will need to be modified to integrate within the 3D macroscale organ or tissue construct. Apart from these techniques, scaffolds can be functionalized with proangiogenic biomolecules, such as platelet-derived growth factors (PDGFs), VEGFs, and basic fibroblast growth factors (bFGFs) for rapid generation of vascular networks or angiogenesis [133,134].
Fig. 4.
(A) Schematic representation of bioprinting of agarose template fibers and subsequent formation of microchannels via template micromolding; (i) A bioprinter equipped with a piston fitted inside a glass capillary aspirates the agarose (inset). After gelation in 4°C, agarose fibers are bioprinted at predefined locations; (ii) A hydrogel precursor is casted over the bioprinted mold and photocrosslinked; (iii) The template is removed from the surrounding photocrosslinked gel; (iv) Fully perfusable microchannels are formed; (B) Photographs of the bioprinted templates (green) enclosed in GelMA hydrogels and the respective microchannels perfused with a fluorescent microbead suspension (pink); (i) bioprinted templates in a GelMA hydrogel construct and (i-a) respective network after perfusion; (ii) 3d branching agarose templates embedded in a GelMA hydrogel construct and (ii-a) resulting 3d branching network [129]; (C) The experimental setup: (i) 3D model of the coaxial nozzle, (ii) cross-sectional view of the coaxial nozzle assembly model with fluid flow paths for hydrogel and crosslinker solutions; (D) (i) Perfusion of oxygenized cell type media, (ii) media flow with intentionally generated air bubbles showing flow [131]; (E) Bioprinting tubular structures with cellular cylinders; (i) Design template; (ii) Layer-by-layer deposition of agarose cylinders (stained here in blue for better visualization) and multicellular pig SMC cylinders; (iii) The bioprinter (see Materials and methods) outfitted with two vertically moving print heads; (iv) The printed construct; (v) Engineered pig SMC tubes of distinct diameters resulted after 3 days of post-printed fusion [132].
3.2.6. Organoids-on-a-chip systems
Organoids-on-a-chip is a combination of organoids and organ-on-a-chip models, which are fundamentally different approaches. Organoids are cell-derived miniature tissue construct that replicates the fundamentally same functional characteristics to their in vivo counterparts. On the other hand, the organ-on-a-chip model relies on the human-made constructs in which cells and their microenvironment are precisely controlled. However, the development of the perfect organoid system is still far from achieving. Therefore, researchers are now exploring the possibility of combining the best features of each approach [135]. The complexity of the stem cell microenvironment is the main barrier in organoid development. The control of the biochemical and biophysical microenvironment only from the self-formation of cell aggregates is not sufficient due to the lack of timely activated signaling pathways to cell fate specification, physical aggregation of different cell types, and mechanical forces that guide the process of self-organization. Organ-on-chip models are explored to achieve not only the microenvironmental control but also model the tissue-tissue and multiorgan interaction and reduce the variability. The development of the organ-on-chip system is a combination of all the previously discussed techniques, including microarray systems, bioreactors, 3D bioprinting, and vascularization.
Demers et al., were able to develop a microfluidic system capable of recreating the microenvironment of the diffusion-based patterning of the neural tube in a tightly regulated manner. This system contains a pair of microchannels that serve as a nutrient supplier and waste product remover of the cell chamber. The cell chamber was filled with Matrigel to provide a 3D culturing matrix for embryonic stem cells and also created a stable morphogen gradient via diffusion across the microchannels and the cell matrix [136]. In another study, iPSCs derived perfused human liver model was developed using cell aggregates encapsulated PEG microtissues. These microtissues were cultured in a range of flow rates by adopting the C-trap chip architecture that enabled robust loading with encapsulated organoids (cell aggregates). This microfluidic system enables an opportunity to query patient-specific liver response in vitro while maintaining the microenvironment [137]. The same microfluidic system was used to develop a human brain organoid-on-chip platform. The multicellular cell aggregates (or embryoid bodies (EB)) were formed using hiPSCs and then suspended in Matrigel. Those EB suspended Matrigel were then pipetted on hydrogel channels of the microfluidic chip. Due to the gelation of Matrigel, those cell suspensions were immobilized on the chip, and the microchannels were connected to the neural differentiation media channel, allowing the formation of brain organoids, as shown in Fig. 5(A). The sufficient nutrient exchange within the brain organoid enabled the formation of large and polarized neuroepithelial surrounding a fluid-filled cavity resembling the ventricular zone. This brain organoid-on-chip model successfully investigated the adverse effect of nicotine on cortical development [138].
Fig. 5.
(A) The conFiguration of the brain organoid-on-a-chip device and enlarged view of the procedures for brain organoid generation on a chip. The brain organoid-on-a-chip system was established for EBs culture, neural differentiation and formation of brain organoids by integrating 3D Matrigel and fluid flow [138]; (B) Development of vascularized HSEs. (i) Schematic description of the protocol to develop vHSEs; (ii) Two different vasculature patterns were used in our studies and generated using fluorescently tagged alginate. Scale bar: 600 μm [142]; (C) 3D convoluted renal proximal tubule on chip; (a) Schematic of a nephron highlighting the convoluted proximal tubule; (b,c) corresponding schematics and images of different steps in the fabrication of 3D convoluted, perfusable proximal tubules, in which a fugitive ink is first printed on a gelatin-fibrinogen extracellular matrix (ECM) [143].
Another vital aspect of organ-on-chip models is the ability to mimic perfusable blood vessels due to the recent advancement of 3D bioprinting techniques [139–141]. Abaci et al., were able to develop human skin construct using a bioengineering approach, which allows printing spatially controlled perfusable 3D vascular network on a precast mold using sacrificial sodium alginate hydrogel, as shown in Fig. 5(B). The vascular network was perfused with endothelial cells (iPSC derived or human umbilical vein derived) while surrounding dermal fibroblasts. This model promoted and guided neovascularization during wound healing in mice [142]. This type of vasculature or microchannel formation is beneficial for the prolonged viability of thick organoids. Similarly, Homan et al., have reported a bioprinting method to create human renal proximal tubules in a perfusable tissue chip embedded in an ECM. As shown in Fig. 5(C), the perfusion chip gasket was first made using silicone ink, and then the ECM base layer was printed using a multi-material 3D printing head using gelatin/fibrinogen gel mixture. The proximal tubule structure was printed on base ECM using fugitive ink pluronic F127. This 3D tubule structure promoted the formation of the tissue-like epithelium with improved phenotypic and functional properties compared to the same cell (proximal tubule epithelial cells) grown in 2D controls [143].
The control of the biophysical microenvironment is also crucial for organoid development. For example, a developing embryo experiences different types of mechanical forces, including fluid shear stress and single-cell generated traction forces. Some mechanical forces are generated due to the coordinated mechanical activity of a large group of cells, such as fetal breathing movement and heart contraction. However, the generation of biomechanical control from cell aggregates or simple hydrogel systems is not possible as these forces act according to ECM signals and soluble morphogens. To address this issue, recent studies have demonstrated microengineered platforms that can actuate mechanical forces in organoids. Lind et al. were able to fabricate cardiac microphysiological devices that can provide continuous contractile stress within the microarchitecture, allowing the self-assembly of physiomimetic laminar cardiac tissue [144]. The multilayer cantilever consists of a base layer, embedded strain sensor, and iPSCs embedded tissue guiding layer supporting self-assembly. Lee et al., developed a stomach-on-chip model in which hPSCs were differentiated to form stomach organoids in the Matrigel matrix. A luminal flow media was introduced to mimic in vivo gastric function through the lumen using a peristaltic pump. The generated fluid flow through the inner compartment produced the rhythmic stretch and contraction to the organoid reminiscent of gastric motility [145]. Homan et al., recently developed a kidney organoid-on-chip platform as a proof-of-concept to Lee’s approach. Under the static condition, hPSCs-derived kidney organoid shows largely avascular and immature glomerular and tubular-like compartments. As a solution, Homan et al. developed a 3D printed millimeter-scale chamber to examine the vascularization and maturation of hPSCs derived kidney organoids, as shown in Fig. 6(A and B). Notably, high fluid flow stress (FFS) resulted in enhanced vasculature and increased maturity in their glomerular and tubular compartments, as shown in Fig. 6(C). Furthermore, concomitant morphogenesis of tubular epithelial cells and podocytes were also enhanced in a shear stress-dependent manner, suggesting the flow generated microenvironment as a pivotal contributor to structural and functional development of kidney organoid [146].
Fig. 6.
(A) Developing renal organoids are placed on an engineered ECM, housed within a perfusable millifluidic chip, and subjected to controlled FSS; (B) representative phase contrast images of entire organoids, scale bars: 50 μm; (C) Confocal 3D renderings for vascular markers in whole-mount organoids cultured under, (i) static engineered ECM, (ii) low-FSS, and (iii) high-FSS conditions, scale bars: 100 μm [146]; (D) (i) Schematic of micromolded collagen in the modified insert and below a side view of stamp using electron microscopy; (ii) Schematic showing the application of an additional gradient comprised of DAPT (gamma secretase inhibitor); (iii) fluorescence images of a polarized crypt-villus unit under the 3-growth factor gradient and opposing DAPT gradient. Mature enterocytes (red, ALP+) and proliferative cells (green, EdU+) are also marked; (iv) Immunofluorescence staining (Olfm4/KRT20) of in vitro human small intestinal tissues showing tissue polarity under the combine growth factor and DAPT gradient; scale bar: 100 mm [147].
The human small intestine epithelium consists of a distinct crypt and villus architecture, and migration and differentiation of cells on these two distinct cellular compartments are driven in part by biochemical gradient factors that specify the polarity of each cellular compartments. This unique architecture of the epithelium was fabricated using crosslinked collagen on the PTFE membrane with the help of PDMS stamps. A modified Transwell insert was used to establish the diffusion-based gradient across the collagen membrane when lower and upper parts were loaded with two different media. This system enables the growth of crypt-villus architecture, as shown in Fig. 6(D) [147]. By using bioengineering techniques, including microfluidics and bioreactor systems, it is possible to maintain the complex microenvironment with relevant biochemical and biophysical cues.
4. Model organoid systems
4.1. Brain organoids
The development of human brain embodies involves a high degree of coordination between the neural stem cells (NSCs) and the dynamic microenvironment in which they exist [34]. The expansion of the neuroepithelium to generate radial glial stem cells (RGS) is the beginning of human brain development [32]. The human brain is a complex structure consisting of different region-specific cell types and complex cellular arrangements. In most situations, part of the specific interest region of the brain is developed due to the complexity [159]. Also, hPSCs derived brain organoids are considered as a robust model to recapitulate the significant features of brain development at early stages and study the neurodevelopmental disorders [160,161]. Most of the studied brain organoids were generated in the range of micro to millimeter size, and Matrigel was used as a scaffold that can promote neuroepithelial bud expansion and organoids growth [98].
Bangley et al., developed a cerebral organoid model that primarily produces more ventral regions using ventralizing drug treatment to increase the ventral forebrain region. By fusing organoids of dorsal and ventral forebrain identities, a dorsal-ventral axis was generated. The fused ventral dorsal organoid exhibited the typical migration of human cortical interneurons from ventral into forebrain regions, as shown by inhibition of CXCR4, a known regulator of interneuron migration [159]. In another study, Lancaster et al., were able to demonstrate the arrangement of organ-like configuration by using a bioengineered construct. The synthetic polymer, poly(lactide-co-glycolide) copolymer (PLGA), microfiber filaments were used to construct a 3D model as a floating scaffold to generate elongated embryoid bodies (EB) as shown in Fig. 7(A). During the organoid self-assembly process, PLGA microfibers served as a scaffold that could maintain dense cell-cell contact and increase the surface-area-to-volume ratio. This system has displayed enhanced neuroectoderm formation and improved cortical development [162]. Zhu et al., demonstrated the use of calcium alginate hollow fiber system to develop brain organoid using iPSCs Matrigel cell aggregates, thereby further used for modeling prenatal alcohol exposure to explore the mechanisms underlying neural dysfunction [163]. Another natural hydrogel system, a sodium hyaluronan and chitosan polyelectrolyte complex termed Cell-Mate3D, was used to facilitate the cerebral organoid formation without having any additional neural induction component. The generated organoid showed evidence of early cortico-genesis and exhibited neural rosette and neural tube-like structure. Also, this model can be used as an in vitro disease model for adrenoleukodystrophy (ALD) or any other neurodevelopmental disease using patient-derived iPSCs [164].
Fig. 7.
(A) Brain organoid formation (i) Schematic of the method for generating engineered cerebral organoids (enCORs). Timeline and media used are shown at the bottom. (ii) hPSCs attach to and coat the PLGA microfilaments (left panel, arrowheads) as well as fibers derived from sea sponge (middle panel), whereas cellulose fibers with similar dimensions (right panel) fail to form elongated embryoid bodies at day 3 with H9 cells, and instead remain as clumps (arrows) only partially attached to the fibers. (iii) Bioengineered embryoid bodies from H9 cells at two time points, days 8 and 11, during neural induction, showing clearing along the edges and polarized neural ectoderm (arrows) [162]; (B) Alginate Supports HIO Survival In Vitro (i) Representative images of live (Calcein-AM, green) and dead (Ethidium-homodimer-1, red) staining of spheroids in alginate and Matrigel after 7 days in culture. Scale bar, 100 μm; (ii) Rheological characterization of alginate hydrogels. Data shown are the mean ± SD from n ≥ 3 gels per condition [171].
4.2. Intestinal organoids
The intestines, derived from definitive endoderm, are a vital organ that performs the function of digestion, nutrient absorption, and waste elimination, which are further divided into small and large intestines. The complex structure of the intestine consists of several functionally distinct areas, including the duodenum, jejunum, and ilium, and also its surface comprises a highly folded epithelium of villi, microvilli, and intestinal crypts [165,166]. The first-ever modern organoid system reported in the literature was developed by Hans Clevers’s lab based on the self-renewal epithelium of the small intestine. The Lgr5+ stem cell, which can undergo self-renewal of the epithelium, was first reported by this group and thereby presented the single sorted Lgr5+ stem cells that can also initiate the crypt-villus units in the intestine [167]. Recently, Serra et al., reported that the generation of intestinal organoids is not only limited to Lgr5+ cells but also depends on transient YAP1 activation. YAP1 requires to display transient cell-to-cell viability in localization that further initiates a Notch and DLL1 lateral inhibition, which guides the dedifferentiation of Paneth cell and subsequent crypt formation [168].
In another study, Watson et al., generated human intestinal organoid (HIO) in vitro from iPSCs that can engraft in vivo. To establish the in vivo model, the cells cultured in Matrigel were differentiated to form organoids for 35 days and then embedded in collagen type I and transplanted them under the kidney capsule of immunocompromised mice. These transplanted organoids had shown digestive functions similar to the HIO, as shown by permeability and peptide uptake studies [169]. Since ECM mechanical properties influence epithelial cell behavior, several studies are demonstrated that the use of hydrogel matrices can provide the desired mechanical microenvironment for HIO formation [170]. In that aspect, Capeling et al., were able to develop an alginate-based hydrogel ECM to support the growth and differentiation of HIOs. Five different concentrations of alginate (ranged from 0.5% to 4% (w/v)) were used, as it was reported that 1% alginate concentration provided the optimal hydrogel matrix, showing the same viability of HIOs in Matrigel as shown in Fig. 7(B) [171]. The storage modulus of Matrigel was very close to the 1% alginate, suggesting the significance of the mechanical microenvironment for HIO formation. Similarly, the PEG-4MAL synthetic hydrogel system was developed to match the storage modulus of Matrigel, which is further discussed in section 3.1 [52]. The extrinsic mechanical cues provided by chemically modified hydrogels have a significant role in hPSCs expansion, differentiation, and HIO formation.
4.3. Kidney organoids
The human kidney is a highly complex mesoderm organ containing up to 2 million epithelial neurons that regulate blood filtration. The functional kidney requires more than 20 distinct cell types, arranging in a spatial architecture, required for excretion and regulation of pH and electrolyte and fluid balance [23,172]. The human kidney can be generated from four types of progenitor population, including nephron progenitors, renal intestinal progenitors, ureteric epithelial progenitors, and endothelial progenitors [173]. Several studies have been demonstrated the ability to form kidney organoids from the differentiation of hPSCs [174,175]. The formation of kidney organoid through primitive posterior streak and intermediate mesoderm from differentiated human embryonic stem cells was reported using Matrigel-coated monolayer culture [176]. The same group studied the different methods to generate all four types of progenitors from hPSCs. Although both collecting duct and nephrons are originated from the intermediate mesoderm, they have a distinct temporospatial origin. Therefore, in another study, Takasato et al., developed kidney organoids that contained nephrons associated with a collecting duct network surrounded by endothelial cells and renal interstitium. Also, by carefully balancing the anterior-posterior patterning of intermediate mesoderm using small molecules such as growth factors, this multicellular kidney organoid was generated (Fig. 8(A)) [23]. This same group developed a method to form all types of progenitors from hPSCs [173]. Apart from the Matrigel, the fibrinogen/gelatin hydrogel system was used to precisely control the ECM microenvironment [143]. In another study, silk has demonstrated the capability of forming kidney organoids using hPSCs, while vascular endothelial growth factors promoted blood vessel formation [177].
Fig. 8.
(A) (i) Global bright field observations of self-organizing kidney organoids across a time series. The success rate of organoid differentiation was 94.2%; scale bars: 1 mm. (ii) Tile scan immunofluorescence of a whole kidney organoid displaying structural complexity; scale bar: 1 mm. (iii) High-power immunofluorescence microscopy showing a nephron segmented into 4 compartments, including the collecting duct (CD, GATA3+ECAD+), distal tubule (DT, GATA3−ECAD+LTL−), proximal tubule (PT, ECAD−LTL+) and the glomerulus (G, WT1+). Scale bar, 100 μm [15]. (B) (i) Schematic representation of in vitro liver bud formation; (ii) Self-organization of three-dimensional human iPSC-LBs in co-cultures of human iPSC-HEs with HUVECs and human MSCs. The time-lapse fluorescence imaging of human iPSC-HEs is also shown here; scale bar: 5 mm [184].
4.4. Liver organoids
The liver is an essential organ that regulates many metabolic functions, including the production of hormones, regulation of cholesterol and glucose storage, and the synthesis of some plasma proteins; and also, the liver regulates the detoxication of the body. So, liver injuries can generate dramatic consequences, leading to hepatic pathologies such as liver fibrosis, hepatitis, cirrhosis, or hepatocarcinoma. Fortunately, the liver has some unique potential to maintain its homeostasis by generating healthy tissue because of the proliferation potential of mature hepatocytes. However, due to the severity of the injuries, sometimes the proliferative capacity of the hepatocytes is not sufficient enough to regenerate the lost tissue [178,179]. Since the liver regulates the detoxication of the body, it is easily susceptible to drug-induced liver failures. Two types of endothelial cells: the hepatocytes and biliary epithelial cells (also known as cholangiocytes) are mainly responsible for liver formation. Apart from these cells, several other non-epithelial cell types, including stellate cells, stromal cells, Kupffer cells, and endothelial cells together generate the functional liver structure called the hepatic lobule [180,181]. Adult hepatic progenitor cells (HPCs) are considered a good source for hepatocytes and hepatocytes like cells compared to the popular ESCs or iPSCs. The HPCs possess several advantages over iPSCs, such as more efficient and less time-consuming differentiation, lower risk of spontaneous mutations, chromosomal rearrangements, and teratoma formation [182].
Broutier et al., developed a protocol to develop long-term expandable liver culture combined with Matrigel, acts as an ECM, and hepatocyte growth factors (HGF), epidermal growth factor (EGF), FGF, and Rspo1 using mouse liver tissues and induced progenitor cells for months [180]. Several studies focused on the development of vascularized liver organoids. For example, Baptista et al., were able to develop a scaffold that human liver cells could easily enter and proliferate to repopulate the decellularized liver scaffold. This decellularization process was preserved the macrovascular skeleton of the entire liver. The human fetal liver cells and endothelial cells were perfused through the vascular structure, and those cells were able to repopulate areas throughout the liver in vitro [183]. In another study, the vascularized and functional liver was generated from iPSCs by in vivo transplantation of in vitro derived liver buds, as shown in Fig. 8(B). Specified hepatic cells were self-organized into 3D liver buds by recapitulating organogenetic interactions between human endothelial and human MSCs. The iPSCs derived liver buds were transplanted into mice models, and human vasculature in liver buds became functional within 48 h by connecting to the host vessels [184]. Manipulated iPSCs showed great potential in future regenerative medicine.
Recently, Sorrentino et al., developed a chemically defined mechano-modulatory 3D culture system for basic research and regenerative medicine applications where PEG based synthetic hydrogel system provided niche for liver organoid culture [185]. In another study, cellulose nanofibril hydrogel system was used as an alternative to Matrigel in generation of liver organoids. The niche provided by the cellulose nanofibrils hydrogel system generated functional hepatocyte-like cells similar or superior to the Matrigel. The mechanical properties of cellulose nanofibrils further supported the differentiation of liver organoids [186].
5. Therapeutic potential of organoids
5.1. Drug development and toxicity screening
The current drug toxicity screening methods are heavily dependent on the traditional 2D cell cultures (monolayer) on plastic, which results in unacceptable failure rates [152,187,188]. The new investigational drugs that have gone through phase II and phase III clinical trials were failed to proceed into the next stage of development 66% and 30%, respectively [152]. The current toxicity assays are the main reason for this high failure rate and hence highlight the need for reliable toxicity assessment. Also, the current concept for in vitro-in vivo correlations does not address the drug toxicity mechanisms but emphasized drug dissolution and bioavailability equivalence [187]. Organoid systems demonstrate improved toxicity screening due to their ability to harness the organ-specific cellular/tissue organization, such as embryogenesis, organogenesis, and angiogenesis, and thereby overcome the challenges. Also, inflammatory and toxicity pathways of the organoids are more similar to the actual physiological conditions, whereas none of those activities are possible in cell mono-or-co-culture in 2D monolayers on hard plastic supports. In the current toxicity screening paradigm, after the initial 2D monolayer studies, animal models are the immediate next step for promising new drugs. However, there is a significant gap between organoid models or simplified 3D models (cell spheroids) and the effects seen in animal models due to the complete lack of correlation in the genomic response to acute inflammatory stress between human and murine models. Furthermore, this discrepancy is proved by the years of data indicating that 57% of the rodent studies failed to predict the actual human toxicity [189]. Despite the abundant knowledge on the implementation of organoid systems, organoid models in toxicity screening are still limited. This might be mainly due to the complexity of the functional organoid formation with the correct stem cell niche, as previously addressed in this review.
Patient-derived organoid models also play a significant role in developing personalized treatment regimes. The disease organoid model can be implemented through disease tissues/cells from disease-site biopsies, which deliver sufficient material to identify the mutations or phenotypic profiling to facilitate more customized treatments. Likewise, both healthy and diseased organoids generated from the patient enable the screening of new drug combinations that selectively targets the diseased tissues only, which further facilitates the more effective treatments with minimal side effects. Since most of the side effects of drugs can be attributed to acute liver toxicity, the liver/hepatic organoids can be successfully used for predicting in vivo liver toxicity of experimental drugs before moving to more expensive clinical trials [10,21]. The majority of drug toxicity screening studies were focused on tumor models because the organoid technology provided a unique platform to ongoing extensive cancer research works due to various reasons, including the ability to expand long term, modify genetically, and remain stable genetically and phenotypically.
Hou et al., developed a primary pancreatic organoid tumor model for high-throughput screening (HTS). The 3D pancreatic cell culture was developed using standard flat-bottom well plates with cell repellent surface, where 3D bioprinting technology that relies on the magnetic force was used to make 3D organoids on surface well plates using cell spheroids in the absence of exogenous ECM components. Nearly 3000 approved drugs were tested on these patient-derived organoids and their 2D monolayer cell culture counterparts [190]. According to the results, some of the approved drugs had shown cytotoxicity against 3D models over 2D culture, showing the efficacy of 3D models, as explained before. Recently, Saito et al., were able to find a couple of drugs, which suppressed the organoid derived from biliary tract carcinomas (BTCs). BTCs are one of the most aggressive malignancies and have shown a poor prognosis. Therefore, the team was able to develop in vitro organoids using cancer tissues obtained from a patient who has diagnosed with BTCs. The organoid derived from non-cancer tissues was also developed to find a drug with minimal side effects. Finally, they found that the antifungal drugs amorolfine and fenticonazole significantly suppressed organoid growth derived from BTCs with minimal toxicity to healthy organoid tissues [191]. Also, Yan et al., developed a human gastric cancer organoid biobank that comprises normal, dysplastic, cancer, and lymph node metastases, including detailed whole-exome and transcription analysis. The developed biobank organoid system from patient-derived tissues was tested for different drugs similar to the respective patients, and the results were promising as the tumor response of the patient and the cancer organoid have shown the same response [192]. It is evident that organoid cultures have the potential to predict in vivo drug response.
Liver organoid models also provide insight into drug toxicity and food toxicity studies. The identification of hepatotoxic drugs in the initial drug development phase would save millions of dollars. We look into more recent liver organoid models used for toxicity studies. Leite et al., were able to develop a human hepatic organoid model, which enables the identification of drug-induced liver fibrosis in vitro. The current models for liver fibrosis in vitro do not consider the role of hepatocyte injury, which triggers a plethora of events leading to the accumulation of ECM proteins like collagen known as fibrosis. The 3D spheroids cultured from hepatic stellate cells (HSCs) and HepaRG were grown into functional liver organoids. Interestingly, allyl alcohol, which was not identified as a hepatotoxic compound using HSC monocultures, showed indirect HSC activation that leads to fibrosis. Moreover, this liver organoid model identified the compound that is not hepatotoxic, but the presence of functional hepatocytes indirectly activates the HSC [193]. Apart from drug screening, some studies were focused on environmental liver toxicity due to heavy metals and pesticides. For example, Forsythe et al., were able to develop a liver organoid from HSCs. Using this model, hepatotoxicity values were obtained for different heavy metal compounds, including lead, mercury, thallium, and pesticide glyphosate. The 3D hepatic model showed more sensitive and accurate readings for toxicity values compared to the 2D HepG models [194]. Recently, Mun et al., were able to develop hPSC-derived hepatocyte-like liver organoids, which can be successfully used for drug screening and disease modeling similar to HSC derived liver organoids [195].
5.2. Organoid as a model to study diseases
The commitment of the organoids as a disease model for various therapeutic potentials is steadily being established. Compared to the traditional cell cultures of a single cell type, the organoid cultures have a clear advantage for disease modeling due to their ability to mimic pathologies at the organ level [9,196]. One of the caveats in the use of animal models when used for disease modeling is the interspecies differences, which can be addressed by using human-derived organoids. A wide variety of organoid models have already been developed that reproduce genetic diseases, cancer, or host-pathogen interactions. For example, one of the most cited studies in recent years was focused on the reproduction of pathological features in the human gastric organoid. The gastric organoid developed from human gastric corpus tissue was microinjected with Helicobacter pylori bacteria, and the resultant organoid successfully reproduced the signs of bacterial infection. This model plays an essential role in studying gastric pathologies in humans because the species-specific features of the stomach make current animal models unsuitable. Also, H. pylori infection in mice is usually mild gastritis and does not progress to ulceration or cancers as it does in humans [197]. In another study, the same type of human gastric organoid model was developed using the direct differentiation of hPSCs instead of gastric cells. After microinjection of H. pylori to the organoid, the progression of the infection mainly depends on the virulence factor CagA with the association of the c-Met receptor [22]. Together, these types of studies help to elucidate the mechanisms underlying human stomach development and diseases.
Another best example of the host-pathogen interaction is the ZIKV incorporated brain or cerebral organoid systems [9,196]. ZIKV has a robust teratogenic effect: infection in pregnant women is associated with microcephaly in newborns. The development of the human brain organoids by mimicking fetal brain formation has allowed us to discover the deficiencies caused by the ZIKV in the brain. Garcez et al., were able to use hiPSCs cultured as neural stem cells, neurospheres, and brain organoids to explore the consequences of ZIKV infection during neurogenesis [19,160,198]. The ZIKV targets the brain cells as it reduces the viability of neurospheres and brain organoids that further abrogates neurogenesis. Furthermore, the main phenotypic effect of ZIKV is premature differentiation of neural progenitors associated with centrosome perturbation, leading to progenitor depletion, disruption of the ventricular zone, and cortical thinning [199]. The effect of different ZIKV strains was also studied to understand the pathogenic activities of different forms [200], and these models were successfully used for anti-ZIKV drug development [201].
Recently, Boretto et al., were able to develop a disease model to capture endometrial diseases. Apart from these disease models, different cancer organoid models were extensively studied to understand the progression of the tumor, and their potential therapeutic procedures, including pancreatic [202], lung [203], liver [204], intestinal [205], and prostate [206] cancer models. The cancer organoids as a disease model for pre-clinical trials and drug screening were extensively reviewed elsewhere [207–210], and therefore, we do not review further on this topic.
6. Future perspectives and limitations
Organoids are one of the most versatile and physiologically relevant models to study human diseases and drug toxicity screening. So far, the traditional cell biology principles, together with the novel bioengineering approaches, have evolved organoid into a new era of translational research platform instead of being a basic research tool. Most importantly, the development of miniature organ-like structures using available microengineering techniques eventually curtail the time to develop autologous tissue/organ for transplantation. Although the development of physiologically similar human organoid for transplantation is far from reality, Taguchi et al., succeeded in inducing metanephric nephron progenitors that capable of reconstructing 3D nephric tubules and glomeruli, the two components of critical kidney function using both human and mouse PSCs in adult mice, thereby prove the capability of kidney transplantation [211]. In this respect, kidney organoid shows enormous therapeutic potential since kidneys have the highest rate of end-stage organ failure, leading to the highest demand for transplant [11].
Most of the current organoid systems were primarily developed via exogenous signals, in which using numerous cell types and bioengineering approaches were able to create the sophisticated stem cell niche. Nonetheless, the targeted genome editing, primarily via CRISPR (clustered regularly interspaced short palindromic repeats) or TALEN (transcription activator-like effector nucleases) technology, enables the manipulation of endogenous genes in clinically relevant cells and organisms. This method can be used to enhance the microenvironment by reprogramming the internal decision-making component of cells. The reprogramming of the cells in organoids would widen the understanding of the organoid formation and function via self-regulation and self-assembly. Most importantly, this technique can be used to correct genetic defects in humans. The first proof-of-concept demonstration of genome editing in the organoid was done by Schwank et al., in which the genome editing of human stem cells in intestinal organoid was demonstrated via CRISPR/Cas9. The intestinal organoids isolated from two cystic fibrosis (CF) patients were used to correct the CF transmembrane conductor receptor (CFTR) mutation by deleting the phenylalanine at position 508. The corrected organoid showed the restoration of CFTR function in the swelling assay [212]. CRISPR/Cas9 genome editing has been recently used in different cancer organoid models [213,214] and kidney organoids [215]. The genome editing of the organoid systems would be an exciting area to explore further to develop treatments for genetic disorders, including CF, Alzheimer’s, and autism.
The preparation of organoid cultures is generally considered a tedious laboratory procedure because of the small size, fragility, and dynamic culture requirements. Therefore, organoids generated through routine procedures have shown substantial variability in size, functional capacity, structural organization, and gene expression. As a solution for this issue, microengineered systems are uniquely suited because of their ability to handle fluids and dynamic cell culture more precisely. For example, Au et al., developed a digital microfluidic platform capable of generating arrays of individually addressable, free-floating, 3D hydrogel-based hepatic organoids. The hepatic organoids were formed in a programmed manner by activating the electrodes in a defined sequence. This programmed actuation system was also used to analyze the acetaminophen-induced hepatotoxicity without manual intervention [152]. Another way of reducing organoid culture variability is by using biosensing elements in culture platforms to allow continuous uninterrupted in situ monitoring. The microfluidic system developed by Zhang et al., had shown the capability to monitor drug-induced organ toxicity in a dual-organ human liver-and-heart-on-a-chip platform. The microfluidics controlling breadboard consists of physical sensors for monitoring microenvironmental parameters (e.g., O2, pH, temperature), electrochemical immunobiosensors for measuring soluble biomarkers, and a miniature microscope for observation of the organoid morphologies [216].
Finally, the lack of vascularization is an apparent reason for the limited lifespan of organoids because of limitations in nutrient supply. This also affects the size of the organoid in which this limited growth affects their maturation. We have discussed various solutions from other emerging bioengineering approaches such as microfluidics, 3D bioprinting, and organ-on-chip models to address this issue. Fabrication of channels that can host endothelial cells and form perfusable vascular units is a promising solution for this issue [217]. However, the current vascularization approaches are mainly restricted to simplified cellular systems.
7. Conclusion
Organoids have the capacity to mimic the specific cellular functions, 3D architecture, and cell-type compositions similar to the actual organ and thereby hold great promise for a range of biological and biomedical applications. In the study of drug discovery, organoid cultures may reduce millions of dollars spent on unnecessary drug developments due to the variability in animal models and humans. Since the organoids can resemble actual human organs, the number of animal models and clinical trials can be reduced, thereby addressing some ethical issues. However, the path towards translating organoid technology into real-life preclinical and clinical applications is still considerably more challenging. The demonstrated results in organoids in drug discovery, personalized medicine, and gene therapy suggested that more widespread acceptance of organoid technology in these fields will become a reality in the coming decades.
The complexity of the cell microenvironment (niche) is still a significant challenge to be addressed in organoid technology. The development of the novel hydrogel system except Matrigel is crucial for regulating the self-organization organoids in a more controlled manner. Increasing the organoid’s life span, which is crucial to create mature and functional tissues that reach homeostasis, is a vital challenge to be addressed in the future. The current 3D bioprinting techniques, microfluidics, bioreactors, and organ-on-a-chip platforms will help facilitate biochemical and biomechanical cues necessary for self-assembly and the prolonged life span of the organoids. However, these techniques must also be improved to meet the complexity of organoid formation and culture. Overcoming these challenges will require a multidisciplinary approach, combining both cell biology and bioengineering.
Acknowledgments
The authors acknowledge funding support from the National Institutes of Health (R01DE023356) and the University of Toledo.
Abbreviations:
- aSCs
Adult stem cells
- ALD
Adrenoleukodystrophy
- BTCs
Biliary tract carcinomas
- CFTR
CF transmembrane conductor receptor
- CRISPR
Clustered regularly interspaces short palindromic repeats
- CT
Computed tomography
- CF
Cystic fibrosis
- dECM
Decellularized ECM
- DMD
Digital micromirror device
- EB
Embryoid bodie
- EGF
Epidermal growth factors
- ECM
Extracellular matrix
- FGFs
Fibroblast growth factors
- FFS
Fluid flow stress
- GPC
Gas pressure chromatography
- GelMA
Gelatin methacrylate
- HPCs
Hepatic progenitor cells
- HGF
Hepatocyte growth factors
- HPLC
High-pressure liquid chromatography
- HTS
High-throughput screening
- hAVICs
Human aortic vascular interstitial cells
- hDPSCs
Human dental pulp stem cells
- HIO
Human intestinal organoids
- hMECs
Human mammary epithelial cells
- hNESCs
Human neuroepithelial stem cells
- hPSCs
Human pluripotent stem cells
- HSF
Human skin fibroblasts
- hUVSMCs
Human umbilical vein smooth muscle cells
- hUVECs
Human umbilical vein endothelial cells
- HA
Hyaluronic acid
- HIF
Hypoxia-inducible factor
- iHep Cells
Induced hepatic cells
- iPSCs
Induced pluripotent stem cells
- LEM
Liver extracellular matrix
- MRI
Magnetic resonance imaging
- PEG-4MAL
Maleimide crosslinked four-arm PEG macromere
- MSCs
Mesenchymal stem cells
- MEMS
Micro-electromechanical systems
- NSCs
Neural stem cells
- PDGFs
Platelet-derived growth factors
- PLGA
Poly (lactide-co-glycolide) copolymer
- PESCs
Pluripotent embryonic stem cells
- PSCs
Pluripotent stem cells
- PCL
Polycaprolactone
Footnotes
Declaration of Competing Interest
The authors declare no conflict of interest
Data availability
Processed data can be obtained from the cited references.
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