ENGINEERED SYNTHETIC MATRICES FOR HUMAN INTESTINAL ORGANOID CULTURE AND THERAPEUTIC DELIVERY

Abstract

Human intestinal organoids (HIOs) derived from pluripotent stem cells or adult stem cell biopsies represent a powerful platform to study human development, drug testing, and disease modeling in vitro, and serve as a cell source for tissue regeneration and therapeutic advances in vivo. Synthetic hydrogels can be engineered to serve as analogs of the extracellular matrix to support HIO growth and differentiation. These hydrogels allow for tuning the mechanical and biochemical properties of the matrix, offering an advantage over biologically derived hydrogels such as Matrigel^™. Human intestinal organoids have been used for repopulating transplantable intestinal grafts and for in vivo delivery to an injured intestinal site. The use of synthetic hydrogels for in vitro culture and for in vivo delivery is expected to significantly increase the relevance of human intestinal organoids for drug screening, disease modeling, and therapeutic applications.

Graphical Abstract

Human intestinal organoids (HIOs) derived from pluripotent stem cells or adult stem cell biopsies represent a powerful platform to study human development, drug testing, and disease modeling in vitro, and serve as a cell source for tissue regeneration and therapeutic advances in vivo. Synthetic hydrogels can be engineered to serve as analogs of the extracellular matrix to support HIO growth and differentiation as well as delivery vehicles in therapeutic applications.

Introduction

Human intestinal organoids (HIOs) are defined as in vitro, three dimensional cellular structures that recapitulate the architecture, structure, cellular composition, and function of human intestinal tissue^1–4. Based on their remarkable ability to mimic human intestinal tissue, HIOs provide powerful platforms for regenerative medicine^5–8, personalized medicine⁹, drug testing^10,11, studies of intestinal development^12,13, and disease modeling^14–16 (Fig. 1). These organoids can be derived from intestinal tissue biopsies¹⁷ (also referred to as patient-derived intestinal organoid [PDO], adult stem cell-derived [ASC], intestinal stem cells [ISCs], or enteroids) or from human pluripotent stem cells (PSC)¹² (induced pluripotent stem cells [iPSCs], or embryonic stem cells [ESC]).

Figure 1.

Applications of human intestinal organoids.

HIO growth and maintenance require a 3D matrix. Most commonly used is Matrigel^™, a murine sarcoma cell-derived reconstituted basement membrane extract that transitions from liquid to gel around 37°C. Matrigel^™ exhibits high lot-lot compositional and structural variability, and it is composed of >1000 matrix proteins, proteoglycans, and growth factors¹⁸. The major components found in Matrigel^™ are laminin, collagen IV, heparan sulfate proteoglycans, and nidogen. The matrix elastic modulus values depend on the total protein concentration ranging from 9.1 Pa to 288 Pa for 3 to 19 mg/mL, respectively¹⁹. Although organoid structures thrive and develop in Matrigel^™, concerns over its tumor-derived nature, potential immunogenicity, and lot-to-lot variability and composition limit the clinical translatability of this material. Alternatives to Matrigel^™ have been actively researched for the past several years^20–22, however, Matrigel^™ remains as the gold standard for growing HIOs.

Efforts to translate HIO therapies from the lab into the clinic can be grouped into two broad areas, one focused on synthetic hydrogel matrices sustaining in vitro expansion and differentiation, and the other one centered on transplantable grafts or delivering the organoids in vivo. In this review, we discuss in vitro and in vivo synthetic hydrogel applications for HIOs. The first section includes an overview of the biology in human intestinal organoids from PSC and ASC. Section 2 includes different design considerations when engineering a synthetic matrix. Section 3 discusses different synthetic hydrogels used for intestinal organoid culture and differentiation in vitro. In the last section, current tissue engineering approaches for intestinal regeneration are summarized.

1. Human intestinal organoid culture and growth

Human intestinal organoids can be derived from pluripotent stem cells or resident tissue adult stem cells. In this section we discuss in vitro differentiation methods for both cell types, compare among different cell sources, and discuss cell-ECM interactions relevant to intestinal organoids.

1.1. Human pluripotent stem cell-derived intestinal organoids

For PSC-derived intestinal organoids, iPSC or ESC lines can be used. Alternatively, patient-specific somatic cells²³ (e.g., skin fibroblasts or blood mononuclear cells) can be reprogrammed to iPSCs. Various differentiation protocols to mature HIOs from iPSCs have been reported^12,24–28, but most of them converge on three stages (Fig. 2). Stage 1 encompasses iPSCs expansion and definitive endoderm (DE) differentiation on planar culture. Stage 2 involves hindgut specification, where the resulting DE monolayer self-assembles into 3D structures that spontaneously bud off from the monolayer into the culture media; these structures are referred to as intestinal spheroids. In Stage 3, the spheroids are collected and embedded within Matrigel^™ or an alternative 3D matrix and cultured to mature into HIOs after ~28 days.

Figure 2.

Human pluripotent stem cell (PSC)-derived intestinal organoid differentiation process.

Differentiation is driven by growth factor cocktails introduced to the culture in a stepwise manner to mimic the early intestinal development in vitro²⁹. The most common molecule used during this stage is activin-a to drive definitive endoderm. WNT3A or CHIR99201 in combination with fibroblast growth factor 4 (FGF-4) are used to direct the hindgut and intestinal specification process¹². To promote intestinal maturation in 3D, a combination of epithelial growth factor (EGF), R-spondin 1 and Noggin are used¹⁷. We refer the reader to this excellent review²⁹ highlighting specific signaling pathways activated by the growth factors and small molecules used during the in vitro HIO differentiation and how this process relates to embryonic patterning.

After 28 days of in vitro culture, HIOs form luminized multicellular structures with columnar intestinal epithelium which self-assembles into crypt-like domains and is surrounded by a mesenchymal cell layer³⁰. Although specialized intestinal epithelial cells are present in the HIO structures, including Paneth cells, goblet cells, enterocytes and enteroendocrine cells, various studies have revealed that these HIOs most closely resemble immature fetal tissue^12,31–33. Importantly, HIO maturity is observed after in vivo transplantation with longer microvilli and Paneth cell localization at the bottom of the intestinal crypt-like domains, resembling adult tissue architecture^30,31. Different approaches have been studied to increase tissue maturity including in vivo transplantation under mechanical manipulation³⁴ and evaluating new intestinal niche growth factors³⁵ in vitro. Other groups have explored co-culturing with endothelial³⁶ and neural³⁷ cells to achieve a mature HIO model. Host-microbe interactions have also been studied as an approach to achieve tissue complexity^38,39. Furthermore, Helmrath and colleagues⁴⁰ reported an approach to obtain HIOs structures with human immune cells exhibiting GALT-like structures and M cells after in vivo transplantation.

1.2. Human adult stem cell- (ASC) or patient-derived intestinal organoids

For patient-derived organoids (PDOs), HIOs can be established from small intestine or colon tissue biopsies obtained through endoscopic procedures, or from organoid biobanks at commercial or academic institutions⁴¹. PDOs have been shown to maintain tissue donor identity, including regional intestinal tissue characteristics⁴², donor’s age⁴³, and donor’s sex⁴⁴. Additionally, the donor’s gastrointestinal disease state is also maintained, thus different disease organoids models can be developed^1,16,45.

Depending on the state of the initial tissue source (e.g., fresh biopsy, established PDO lines), either freshly isolated intestinal crypt fragments or single cells are embedded with a 3D basement membrane (e.g., Matrigel^™). These cells will then form epithelial cyst-like structures termed enteroids which can develop intestinal crypt-like structures under appropriate culture conditions. Sato et al¹⁷ demonstrated that LGR5+ crypt stem cells are imperative for the self-renewal capacity of the formed epithelial structures in vitro. A specific combination of growth factors has been optimized to resemble intestinal stem cell niche conditions in vitro and drive the differentiation of epithelial intestinal cysts to villus-crypt like containing structures. Among these are R-spondin, EGF, and Noggin which are associated with WNT signaling activation, intestinal proliferation, and intestinal crypt expansion, respectively⁴⁶. However, to maintain undifferentiated expansion of epithelial cyst structures, the maintenance/expansion culture media often contains WNT3A and R-spondin growth factors^42,47. The addition of the small molecules A83–01 and SB202190 to the culture media significantly improves long term expansion of these intestinal structures⁴⁶. Other groups have studied alternative growth factors to stimulate self-renewal and in vitro expansion of patient-derived intestinal organoids including the use of insulin growth factor 1 and fibroblast growth factor 2 (IGF-1 and FGF-2)⁴⁸. Intestinal epithelial cell types are found on intestinal epithelial organoids containing crypt-like structures including Paneth cells, goblet cells, enterocytes and enteroendocrine cells, while the cellular composition of epithelial organoids with cyst morphology is mostly intestinal stem cells and transit amplifying cells^47,48. PDOs do not contain a mesenchyme layer surrounding the structure^4,17.

1.3. PSC- vs ASC-derived human intestinal organoids

The use of PSC- or ASC-derived human intestinal organoid ultimately depends on the intended application or scientific question of interest. PSC-derived organoids exhibit an immature fetal-like state and rely on in vivo transplantation for full maturation. However, they present a valuable tool to study intestinal development and gastrointestinal disease mechanisms. ASC-derived HIOs are composed primarily of epithelial structures isolated from adult intestinal crypts, while PSC-derived HIOs contain mesenchymal layer, and can be engineered to present immune, neural, and endothelial structures. Patient-derived organoids maintain donor identity, making these excellent in vitro models for drug testing and personalized medicine. HIOs can also be derived using induced pluripotent stem cells derived from human donors, however, this process is laborious and costly when compared to biopsy-derived intestinal tissue. Nevertheless, iPSC-derived HIOs present an alternative for patients who lack healthy intestinal tissue for available for biopsy. When considered for in vivo intestinal regenerative therapy and clinical translation, both iPSC- and ASC-derived are promising candidates, however, manufacturing practices, time and costs, autologous versus allogeneic tissue source, and availability of tissues should be considered⁴⁹.

1.4. HIO cell-matrix interactions

The extracellular matrix (ECM) is a network of structural proteins, such as laminin, collagen and fibronectin, and proteoglycans that serves as a reservoir of growth factors and cytokines⁵⁰. The protein composition of the ECM varies from tissue to tissue, and it influences cell behaviors through mechanical and biochemical signaling^51,52. The main mechanism for cells to interact with their surrounding ECM is through integrin transmembrane receptors which drive mechanotransduction through various signaling pathways^52,53. Integrins exist as dimers, consisting of an alpha and a beta subunit and bind to specific peptide sequences found in the different ECM proteins. Additionally, cells actively remodel their ECM, a process crucial to organogenesis, wound healing, and tissue regeneration⁵⁴. This process is guided by the cleavage of different ECM components through proteolysis, mainly by matrix metalloproteinases and cathepsins^55,56.

For epithelial tissues, the basement membrane is a specialized type of ECM serving as physical support for the epithelial tissue layer and separating the epithelium from the mesenchymal tissue layer⁵⁷. Within the intestinal stem cell niche, the basement membrane is composed of laminins, collagen IV, and fibronectin and plays essential roles regulating epithelial cell homeostasis⁵⁸. Integrins regulating the stem cell niche are spatially distributed through the crypt-villi axis highlighting the importance of ECM composition dictating intestinal epithelial cell signaling⁵⁹. An important mechanosensing pathway is Hippo-YAP/TAZ signaling⁶⁰. Intestinal organoids have been used to study in vitro the activation of the transcription factors and their implications in tissue repair⁶¹. Additionally, the nuclear translocation of yes associated protein (YAP) has been showed be a regulator of intestinal spheroid survival⁸ and intestinal stem cell expansion⁶². We refer the reader to this detailed review⁶³ discussing experimental strategies to study the mechanobiology of the intestinal epithelium, including intestinal organoids.

2. Rational design of synthetic hydrogel matrices for in vitro and in vivo applications

Synthetic hydrogels are water-swollen crosslinked polymer networks which can be functionalized with biological motifs/signals depending on the intended application⁶⁴. Synthetic hydrogels represent a highly tunable alternative to natural matrices to support in vitro growth for 2D and 3D single and multi-cellular structures⁶⁵, and as cell delivery vehicles for in vivo applications⁶⁶. Synthetic hydrogel systems have been also used as in vitro platforms to study key developmental and pathophysiological processes such as cell differentiation^67,68, cell migration^69,70, and cell attachment^71,72. Common synthetic polymers that have been used in the tissue engineering field are listed in Table 1. Notably, poly(ethylene glycol) (PEG)-based hydrogels represent the most widely used system. Engineered hydrogel matrices offer the advantage of tuning the biochemical and biophysical properties of the matrix. This modular approach allows for manufacturing designer matrices for particular applications (e.g., in vitro or in vivo). In this section we will discuss the different design considerations for synthetic hydrogel systems, with a special focus on PEG-based hydrogels.

Table 1.

Common synthetic hydrogel polymer backbones in tissue engineering applications^73–77

Polymer	Abbreviation
Poly(ethylene glycol)	PEG
Poly(acrylamide)	PAAm
Poly(vinyl alcohol)	PVA
Poly(acrylic acid)	PAA
poly (lactic acid)	PLA
poly(lactic-co-glycolic acid)	PLGA
poly(caprolactone)	PCL

2.1. Crosslinking chemistries for synthetic hydrogels

An important consideration when engineering synthetic hydrogels is the crosslinking reaction for the particular hydrogel system. Hydrogels, natural or synthetic, can be prepared using one or several crosslinking mechanisms involving physical, covalent, or ionic crosslinks⁷³. For synthetic hydrogels, covalent crosslinking generally offers a higher control over the mechanical properties of the matrix when compared to other crosslinking mechanisms⁷⁸. A critical requirement is that the crosslinking reaction must be cytocompatible and without toxicities.

One attractive covalent crosslinking reaction for biomedical hydrogels is the Michael-type addition reaction (Fig. 3, i). This type of ‘click’ chemistry involves a Michael donor, typically a thiol group, reacting with a Michael acceptor, to produce a Michael adduct by creating a carbon-carbon bond at the acceptor’s β-carbon⁷⁸. A commonly used reactive end group in PEG hydrogels is maleimide⁷⁹, which acts a Michael acceptor, and dithiol crosslinker molecules as the Michael donor. This reaction involves neutrophilic attack of the reactive thiolate ion on the crosslinker to the double C-C bond on the maleimide group forming a thioether covalent bond⁸⁰. This reaction occurs at physiological pH and temperature making it widely used conjugation for in vitro and in vivo applications. A common strategy is to react macromer units with crosslinker molecules to yield a crosslinked network via step-growth polymerization^78,81.

Figure 3.

Common crosslinking reactions for hydrogels in tissue engineering applications (i) Michael type addition reaction (adapted from Phelps et al)⁷⁹, (ii) photo-activated thiol-ene radical mediated step-growth reaction (adapted from Lin et al)⁸³, and (iii) strain-promoted azide-alkyne cycloaddition (SPAAC) (adapted from Agard et al⁸⁶ and Macdougall et al⁹⁰).

Another common type of covalent crosslinking reaction is photo-activated thiol-ene radical mediated step-growth reaction such as PEG-norbornene⁸² (Fig. 3, ii). This type of click reaction occurs between the norbornene groups and a dithiol crosslinker molecule in the presence of a photoinitiator (e.g., lithium phenyl-2,4,6-trimehtylbenzoylphosphinate, LAP). After exposure to ultraviolet (UV) light, radicals from the photoinitiator deprotonate the sulfhydryl group on the crosslinker forming an activated thyl radical. The thyl radical attacks the C-C double bond on the norbornene group forming a thioether covalent bond. As a result, a carbon center radical is formed on the norborne group which regenerates another thyl radical. The reaction will terminate once the limited moiety is depleted (either thiol or norbornene)^82,83. This step-growth reaction provides a higher spatiotemporal control of the crosslinking reaction compared to radical-initiated chain growth polymerization as seen in PEG diacrylate hydrogels (PEGDA).

Finally, another “click” chemistry crosslinking reaction used in the tissue engineering field is strain-promoted azide alkyne cycloaddition known as SPAAC (Fig. 3, iii). In this reaction, a covalent bond is formed between a crosslinking molecule containing an azide and the macromer conjugated with a cyclooctyne containing-functional group⁸⁴. A commonly used PEG macromer to conduct this reaction is PEG-dibenzocycloocytne (PEG-DBCO)⁸⁵. This covalent reaction occurs as a [3+2] cycloaddition of azides and cyclooctyne under physiological conditions in the absence of a catalyzing agent⁸⁶. A specific subset of hydrogels formed through SPAAC crosslinking are used in photodegradation strategies, conducted after UV exposure of nitrobenzyl ether moiety⁸⁷ introduced into the covalent crosslinks or using an allyl sulfide bis(azide) crosslinker in the presence of a photoinitiator such as LAP⁸⁸.

For in-depth discussion of other functional groups involved in Michael-type addition reactions, step-growth reactions, SPAAC reactions and other types of crosslinking reactions, the reader is referred elsewhere^73,78,88,89.

2.2. Mechanical properties of synthetic hydrogels

Single cells and multicellular structures respond to diverse cues provided by their microenvironment. An important property of the microenvironment is matrix stiffness (more generally, viscoelasticity). The stiffness of biological extracellular matrices (ECM) is controlled by the composition, architecture, and crosslinking of ECM proteins present in the specific tissue. Matrix stiffness controls how cells migrate and regulate signaling pathways important in tissue repair, proliferation, and survival⁵⁰.

For hydrogel systems, the bulk mechanical properties are often measured by rheology⁹¹ and described by the complex modulus consisting of the storage (elastic) modulus (G’) and loss (viscous) modulus (G”). Alternatively, the elastic or Young’s modulus (E) can be measured using other testing modalities such as atomic force microscopy^79,92 and microrheology^93,94. For synthetic hydrogels, the mechanical properties are heavily determined by the crosslink density, often described as the mesh size defined as the distance between crosslinks⁹⁵. For a synthetic hydrogel exhibiting rubber-like elastic behavior, the mathematical relationship between mesh size (ξ) and storage modulus (G’) is provided by the following equation^96,97, where G’ is the storage modulus, A is Avogadro’s number, R is the gas constant, and T is temperature. For synthetic hydrogels mesh size is the nanometer range⁹⁷.

$$\xi ={\left(\frac{G^{\prime }A}{RT}\right)}^{-1/3}$$

At the hydrogel synthesis level, different approaches are available to change the mechanical properties of the resulting crosslinked gel (Fig. 4). For hydrogels based on the crosslinking of pre-synthesized macromer (e.g., multi-armed or ‘star’ building blocks) units, these include (1) polymer macromer amount in solution (polymer density, w/v%), (2) macromer size, and (3) the number of reactive arms in the macromer (e.g., 4-arm or 8-arm). The higher the polymer density, the stiffer the matrix. Conversely, the smaller the macromer, the stiffer the matrix will be. Finally, a higher number of arms on the macromer will result in a stiffer matrix by increasing the crosslink density. It is important to note that these increase the stiffness of the matrix by decreasing the mesh size of the network. Changes in mesh size can impact other properties of the matrix, such as the diffusion of molecules through the network.

Figure 4.

Methods to control the mechanical properties of synthetic hydrogel by tuning: (i) the polymer density, (ii) macromer size, and (iii) number of reactive arms.

Another factor related to the mechanical properties is how the hydrogel responds to applied stress. Most PEG hydrogels are considered to exhibit elastic behavior, observed when there is almost instantaneous and reversible deformation when stress is applied. In contrast, viscoelastic materials exhibit time-dependent stress-strain responses^22,96,98. The relationship between elastic modulus (E), stress (σ) and strain (ε) for linearly-elastic material and viscoelastic material can be described with the following equations⁹⁹. For a linearly elastic material, there is a linear relationship to stress described below.

$$\sigma =\text{E}\epsilon$$

On the other hand, for viscoelastic materials the modulus values depends on time and is reflected by this equation.

$$\sigma =\text{E}\left(\text{t}\right)\epsilon$$

Two common viscoelastic behaviors are stress relaxation and stress stiffening. During stress relaxation, the stress on the material decreases over time when exposed to a constant strain¹⁰⁰. In the case of stress stiffening, the material will become stiffer when exposed to a constant stress greater than the critical stress point for the particular material²². Physiological ECMs typically exhibit viscoelastic behavior, thus efforts to modify PEG hydrogels or engineer viscoelastic materials have been investigated^100,101.

2.3. Matrix degradability for synthetic hydrogels

The degradability of the matrix is another important design consideration when engineering a synthetic matrix. Matrix remodeling through matrix metalloproteinases (MMP) and cathepsins is essential for cell survival, proliferation, growth, and differentiation¹⁰². Natural ECMs are mostly degraded by cell-secreted proteases, targeting specific amino acid sequences present on the ECM proteins such as collagen, laminin, fibronectin, and vitronectin¹⁰³. For synthetic hydrogels, the degree of degradability can be controlled by enzymatic digestion, hydrolytic degradation, or photoreactivity. For thiol-based crosslinking of synthetic hydrogels (e.g., PEG-maleimide and PEG-norbornene), a chemically stable crosslinker consisting of short di-thiol molecules is used such as dithiothreitol (DTT) or PEG-dithiol (PEGDT) to generate non-degradable networks. To target protease degradability, a short peptide crosslinker can be designed to the specific protease secretome of the desired cell or tissue of interest. A list of commonly used protease-degradable crosslinkers used in PEG hydrogels is included in Table 2. Other approaches to degrade synthetic hydrogel matrix include photocleavable crosslinkers^104,105 and crosslinkers or macromers incorporating groups susceptible to hydrolytic degradation^106,107. Controlling the degradation reaction kinetics and studying the degradation by-products are important factors to consider, especially when designing for in vivo applications. The rate at which the matrix degrades not only depends on the type of crosslinker used, but also on the density of crosslinks in the matrix. For example, a hydrogel with higher polymer density or higher number of arms will take longer to degrade than a hydrogel with less polymer and fewer arms, respectively. The degradation rate for degradable hydrogels can vary between hours to weeks depending on the specific design considerations^107–109. Combining different degradation mechanisms in the same matrix may provide additional control over the degradation rate.

Table 2.

Degradable peptide crosslinkers used in PEG hydrogels

Protease sequence	Degradation mechanism	References
GCRDGPQGIWGQDRCG	MMP cleavable	8,110,111
GCRDVPMSMRGGDRCG	MMP cleavable	112,113
GCRDIPESLRAGDRCG	MMP cleavable	67
GPQGIAGQ	MMP cleavable	69,111
LPRTG	Sortase	111,114
GGLGPAGGK	Collagenase	115
AAAAAAAAAK	Elastase	115
AAPVR	Elastase	115

MMP – Matrix metalloproteinases

2.4. Functionalizing synthetic matrices with bioactive ligands

A key design parameter is functionalizing the synthetic hydrogel to present bioactive signaling molecules. Cells interact with the ECM mainly through integrins, transmembrane proteins dictating essential cellular processes through cell signaling pathways^52,53. Synthetic polymers lack biological motifs that interact with cells, thus including short peptide sequences to promote integrin-mediated cell adhesion is a strategy to overcome this challenge^73,116,117. These short peptide motifs mimic sites in ECM proteins that are recognized by integrin and other adhesion receptors. The most commonly used cell-adhesive peptide contains the arginine-glycine-aspartic acid (RGD) motif found in fibronectin and other ECM proteins^116,118. We provide a list of commonly used bioadhesive peptides on Table 3. While most naturally derived matrices contain bioactive signals, a synthetic hydrogel system offers the opportunity to evaluate the effect of a particular peptide sequence or density, or to explore the synergy between a combination of peptides in the same matrix. It is recommended to design the matrix in accordance to the integrin profile for the targeted application or cell type. In addition to integrins, cells communicate with neighboring cells through different membrane receptors such as cadherins¹¹⁹. The hydrogel can be designed to incorporate peptide sequences targeting these receptors to activate cell-cell mediated adhesion^120,121.

Table 3.

Bioactive peptides with integrin-binding sequences used in PEG hydrogels

Peptide Name	Peptide sequence	ECM mimic	References
RGD	GRGDSPC	Fibronectin	8,62,118,126
GFOGER	GYGGGP(GPP)5GFOGER(GPP)5GPC	Collagen- type I	111,127,128
IKVAV	CGGAASIKVAVSADR	Laminin α1	8
AG73	CGGRKRLQVQLSIRT	Laminin α1	129
RKG	GKKQRFRHRNRKG	Vitronectin	130
YIGSR	CGGEGYGEGYIGSR	Laminin β1	131
A5G81	AGQWHRVSVRWG	Laminin	132
A5G84	TWSQKALHHRVP	Laminin	133
PHSRN	PHSRN	Fibronectin	111,118

Other biological signals of interest to incorporate in synthetic hydrogels include growth factors and cytokines (Fig. 5). In naturally occurring ECMs, growth factor sequestration and presentation are important functions of the matrix controlling important cellular processes. The matrix can be engineered to present matrix bound growth factors^112,122,123, but also can be engineered to incorporate growth factor binding domains into the synthetic matrix^124,125.

Figure 5.

PEG hydrogel presenting bioactive ligands, growth factors, and growth factor-binding sequences.

2.5. Biological tissue-derived hydrogels

In addition to synthetic matrices, hydrogels engineered from naturally derived materials are also used in tissue engineering applications. These matrices are derived from purified naturally occurring ECM proteins (collagen, fibrin, laminin, gelatin, elastin) or polysaccharides (alginate, hyaluronic acid)¹³⁴. Hydrogels derived from decellularized tissue are another example of naturally derived materials. In this case, a particular tissue or organ of interest undergoes a series of decellularization steps to remove the cellular contents and isolate the native ECM and structure for the particular tissue or organ¹³⁵. An advantage to these natural material-based hydrogels is that they contain innate biochemical and mechanical properties, including integrin adhesion sites and protease degradable sequences⁶⁵. Despite these advantages, several limitations hinder their use for fundamental studies, drug screening, and clinical application. These materials exhibit lot-to-lot compositional and structural variability, limited processability and manufacturing, and potential for contaminants and pathogen transfer, especially for animal tissue-derived matrices^18,127. For these reasons, synthetic hydrogels offer promising alternatives as in vitro culture platforms and in vivo delivery vehicles.

3. Engineered matrices for HIO culture

HIOs offer a disruptive tool to study cell development, disease progression, and identify drug targets. Engineering a synthetic matrix to support in vitro human intestinal organoid development is a strategy to overcome dependence on Matrigel^™, by increasing reproducibility in scale-up, manufacturing, and screening. Table 4 summarizes the advancements in synthetic and natural derived hydrogels used for human intestinal organoid culture.

Table 4.

Synthetic and natural derived hydrogels for human intestinal organoid culture

Engineered Matrix	Polymer content (stiffness)	Ligands /concentration	Degradation mechanism	Cell source	Pros and cons	Refs
Synthetic hydrogels
8arm PEG-vinyl sulfone + 8arm PEG-acrylate	13.33% w/v PEG precursor (G’=1.3kPa to 190Pa)	RGD [1 mM] and laminin-111 [0.1 mg/mL]	Hydrolytic degradation	Mouse ISC	Pros: Tunable matrix properties, low batch to batch variability, user designed degradation mechanisms Cons: Needs adhesive ligand functionalization	62
LDTM 8 arm PEG-vinyl sulfone LDTM 4 arm PEG-acrylate	2.5% w/v (shear modulus =1kPa) 3% w/v (shear modulus=400Pa to 100Pa)	RGD [1 mM] RGD [1 mM]	MMP degradation Hydrolytic degradation	Mouse ISC Human intestinal PDO		129
4 arm PEG-maleimide	4% w/v (G’=100Pa)	RGD [2 mM]	MMP degradation	Human PSC-HIOs		8
8 arm PEG-vinyl sulfone	5% w/v (G’=926Pa)	GFOGER [1.5 mM] ECM binding peptides	MMP degradation	Human intestinal PDO		111
Natural derived hydrogels
HELP matrix	(G’=1kPa)	RGD [1 mM] hyaluronic acid elastic like protein	Enzymatic degradation	Human intestinal PDO	Pros: Control over protein sequences dictating the hydrogel mechanical and chemical properties Cons: Complex ELP induction and purification process from E. coli cultures	131
Alginate	1% w/v (G’=100Pa)	N/A	Non degradable	Human PSC-HIOs	Pros: naturally found material, no further chemical modifications needed, low cost Cons: lot to lot variability, limited long term stability due to ionic crosslinking	132
Porcine intestine derived ECM	2 mg/mL (G’=40–50Pa for 5 mg/mL)	Collagen proteoglycans glycoproteins	Enzymatic degradation	Mouse ISC, PSC-HIOs	Pros: ECM protein composition matched to the tissue of interest Cons: donor to donor variability, multi-step preparation process	133
Type I collagen	3 mg/mL (G’ not reported)	Collagen	Collagenase degradation	Mouse ISC	Pros: innate bioactive adhesion sites, no further modifications needed Cons: lot to lot variability, structural variability	170

LDTM – low defect thiol Michael addition, MMP-matrix metalloproteinases, G’-storage modulus, PDO- patient derived organoid, PSC-HIO –pluripotent stem cell human intestinal organoid, ECM – extracellular matrix.

3.1. PEG hydrogels

Progress in engineering PEG hydrogels for intestinal organoid culture has focused on identifying matrix mechanical and chemical properties that support viability and development of murine, human patient-derived, or PSC-derived intestinal organoids post-encapsulation. The first group to demonstrate that mouse-derived intestinal stem cell (ISC) colonies can grow within engineered PEG hydrogels was Gjorevski et al⁶². They determined the optimal matrix stiffness (1.3 kPa) supporting ISC survival and expansion when the ISCs were embedded in an 8-arm PEG hydrogel functionalized with RGD peptide. However, when evaluating the organoid differentiation and budding formation, the authors observed a softening matrix (G’ = 190 Pa) in the presence of RGD and laminin-111 was needed for differentiation. The matrix softening, driven by hydrolysis, was achieved by combining 8-arm PEG conjugated with vinyl sulfone (PEG-8VS) and 8-arm PEG conjugated with acrylate (PEG-8Acr) to act as a degradable macromer. Hushka et al⁸⁵ observed softening the matrix during murine ISC colony formation was conducive to longer crypts, suggesting higher levels of differentiation. They designed a 4-arm PEG functionalized with dibenzocyclooctyne (PEG-4DBCO) crosslinked with allyl sulfide bis(azide) molecule. This strain-promoted azide alkyne cycloaddition (SPAAC) crosslinking reaction results in a covalent bond between the cyclooctynes and the azide group. In the presence of soluble glutathione and photoinitiator LAP, after exposure to 365 nm light, the crosslinking bis(azide) containing molecules will form a thiyl radical species that crosslinks in an user-controlled time dependent manner. They observed the highest ISC colony viability with a 5 wt% PEG-4DBCO hydrogel (G’ = 1.3 kPa) functionalized with 0.8 mM RGD and murine full length laminin protein entrapped in the matrix.

For primary human intestinal enteroids, Hernandez-Gordillo et al¹¹¹ identified a synthetic hydrogel composition conducive to the highest enteroid formation efficiency using a 8-arm vinyl sulfone-terminated PEG macromer (PEG-8VS, 20 kDa, 926 Pa) functionalized with a collagen I-mimicking adhesive peptide (GFOGER) at 1.5 mM and crosslinked with an MMP-degradable peptide. In this work, the authors also incorporated peptides with affinity to ECM molecules to promote synthetic matrix interaction with the expected secreted ECM from the encapsulated cells. Rezakhani et al¹²⁶ engineered a low-defect thiol-Michael addition (LDTM) PEG hydrogel to grow human patient-derived organoids. The synthetic matrix formulation that supported colony formation and stemness of human intestinal stem cells was an 8-arm PEG-VS hydrogel (2.5% w/v, shear modulus = 1 kPa) containing 1 mM RGD peptide with MMP degradable crosslinker. Additionally, they observed that a hydrolytically degradable hydrogel which softens overtime prepared with 3% (w/v) 4-arm PEG acrylate macromer presenting 1 mM RGD yielded structures of similar organoid phenotype when compared to Matrigel^™. The hydrolytic degradation resulted in a change in shear modulus of the synthetic hydrogel from 400 Pa to 100 Pa over 4 days. These LDTM hydrogels were synthesized by crosslinking PEG macromers conjugated with bifunctional peptides to unconjugated PEG-VS macromers to avoid potential intramolecular crosslinking between functional groups. Interestingly, there are discrepancies among adhesive ligand requirements (e.g., GFOGER vs. RGD) driving human patient-derived intestinal organoid growth and maturation. This effect may may be due to cell line variability as it known PDOs retain the donor’s identity^42–44. When compared to the mouse intestinal organoids studies discussed in this section, the patient-derived organoids exhibited a cyst-like morphology instead of crypt-like containing structures as observed with the murine organoids. Differences in cell source species may contribute to this observation¹³⁶.

To support the development of PSC-derived HIOs, Cruz-Acuña et al⁸ developed a 4-arm PEG-maleimide terminated (PEG-4MAL) hydrogel system. This group demonstrated that iPSC- and ESC-derived HIOs survive and undergo differentiation from spheroids to mature HIOs for 28 days in the synthetic matrix. The optimal synthetic hydrogel formulation promoting growth was 4% PEG-4MAL (20 kDa) functionalized with 2 mM RGD peptide and GPQW protease degradable crosslinker. This study showed that iPSC-derived HIO survival depends on adhesion to RGD peptide compared to other cell adhesive peptides (laminin derived AG73 and IKVAV, and collagen I-derived GFOGER), and mechanical properties of G’ = 100 Pa. Reduced viability was observed in the organoids grown at the same ligand density for AG73, IKVAV and GFOGER, when compared to RGD, showing adhesion peptide type is an important matrix design factor. Additionally, protease-dependent degradability of the hydrogel matrix was required for survival of these structures. The HIOs grown on this matrix demonstrated organoid growth kinetics, lumen formation, epithelial polarization, and expression of differentiation markers equivalent to Matrigel^™ grown organoids.

3.2. Other engineered hydrogels for HIO culture

In addition to PEG gels, other engineered matrices that support human patient-derived and iPSC-derived HIOs have been reported. Hunt et al¹³⁷ engineered a hyaluronan-elastin like protein (ELP) matrix to grow human patient-derived intestinal organoids. The recombinant ELP molecule is covalently tethered to hyaluronic acid to generate the hydrogel matrix. The recombinant ELP contains RGD integrin-binding motifs and a hydrazine moiety which reacts with benzaldehyde group incorporated in the HA protein. These investigators observed that enteroids with cyst formation and polarized epithelium develop from single cells after 3 days in culture. After treating the cultures with organoid differentiation media, they observed a change in phenotype and differentiation gene expression at similar levels to Matrigel^™. They determined that HA was an important signaling component in tandem with RGD to promote enteroid formation. Additionally, a stiff matrix (G’ = 1 kPa) resulted in larger spheroids when compared to a softer matrix (G’ = 400 Pa).

Capeling et al¹³⁸ showed that unmodified food grade alginate supports the growth and development of iPSC-derived HIOs. This polysaccharide hydrogel is crosslinked through ionic interactions when exposed to calcium chloride. In this work, they varied the weight percent of alginate in solution to study the effects of stiffness in the viability of immature HIOs. They found that a stiffness of G’ = 100 Pa, similarly to the PEG-4MAL hydrogel⁸, supports high spheroid viability and HIO development. The authors concluded that under the lack of biochemical signals in the matrix, iPSC-derived HIOs can deposit their own ECM niche thus only relied on mechanical support to undergo differentiation when grown in alginate hydrogels.

Another approach to engineer a growth matrix for intestinal organoids is using gastrointestinal (GI) tissue derived hydrogels¹³⁹. In this work, Kim et al described a decellularization protocol to preserve major ECM components from porcine intestinal tissue including glycosaminoglycans (GAGs). After a proteomic analysis of the obtained porcine intestinal ECM hydrogel, they observed collagen to be the principal protein component followed by proteoglycans and glycoproteins. The group identified a 2 mg/mL ECM hydrogel to support mouse ISC-derived intestinal colony formation and promote stemness. They reported the storage modulus of the tissue ECM hydrogel corresponding to 5 mg/mL ECM concentration varied between 40 Pa to 50 Pa depending on the decellularization protocol used. They also reported that their matrix supported the expression of differentiation markers of hPSC-derived intestinal organoids.

Groups have also evaluated the growth of intestinal organoids under viscoelastic hydrogel conditions. Elosegui-Artola et al¹⁴⁰ observed that viscoelastic behavior in alginate-Matrigel gels led to increased growth and symmetry breaking of murine intestinal organoids when compared to intestinal organoids grown in elastic matrices. Additionally, they reported rapid organoid growth and crypt like structure formation in stiffer viscoelastic matrices (G’ = 1500 Pa, relaxation time (τ) = 5 s) when compared to softer viscoelastic matrices (G’ = 500 Pa, τ = 5 s). Similarly, Chrisnandy et al¹⁴¹ reported a synthetic hybrid PEG hydrogel formulation containing both physical and chemical crosslinks exhibiting stress relaxation properties with storage modulus values ranging from 300 Pa to 1000 Pa (τ = 24 s). When compared to only covalent crosslinked gels, the mouse intestinal organoids grown in hybrid PEG formulation showed higher gene expression of symmetry-breaking markers and YAP1 target genes.

3.3. Discussion

When evaluating different engineered matrix approaches for HIO growth and culture, one of the differentiating factors among mouse ISC, patient-derived, and PSC-derived HIOs is the matrix mechanical properties. In terms of matrix stiffness, a stiffer matrix (G’ ~ 1000 Pa) supports patient-derived HIOs and mouse ISC, while for PSC-derived HIOs, a softer matrix (G’ ~ 100 Pa) was preferred. Another key matrix property discussed was matrix degradability, either through UV or hydrolytic matrix softening or protease-driven degradation, which guided HIO survival and differentiation. Cell adhesive ligand presentation is an important property with RGD, full murine laminin, and GFOGER representing the most effective. Hernandez-Gordillo et al¹¹¹ discussed the importance of taking into consideration donor to donor variability when engineering synthetic matrices for patient-derived organoids since they observed changes across donors in enteroid size and growth within the engineered formulations.

For PSC-derived HIOs, a main difference across engineered matrices is the need for solely providing mechanical support (alginate hydrogels) versus the combination of adhesive peptide presentation and mechanical support (PEG-4MAL hydrogels). When grown in an alginate matrix¹³⁸, the investigators observed that the surrounding HIO mesenchyme did not invade the surrounding matrix, whereas in Matrigel^™ or RGD presenting PEG-4MAL gels⁸, the mesenchymal cell layer infiltrated the matrix. A recent study¹⁴² showed that PSC-derived HIOs could grow in suspension culture suggesting that a 3D ECM matrix might not be necessary for growth. These authors hypothesized that in the absence of a 3D matrix, the mesenchymal layer self-organized and gave rise to a serosa mesothelial tissue layer surrounding the HIOs similar to human fetal intestine. One limitation, however, was the low spheroid to HIO maturation yield when grown in suspension when compared to Matrigel^™. Although recent efforts have focused on characterizing the mesenchymal cell population^{31,32,143,144}, it remains unclear how the different culture approaches could influence the mesenchymal and epithelial cell components in the final HIO entity. Overall differences in matrix requirements (stiffness, degradation, and adhesive ligands) across cell source (PDO vs. PSC), and species (human vs. mouse) were observed in the discussed studies. Cell line variability as well as interspecies variability must be taken into account while designing an engineered matrix.

Engineer matrices can be expanded from growth substrate to other in vitro applications, such as drug screening and drug testing applications¹⁴⁵. These also serve as a tool to study how human intestinal organoids respond to spatiotemporal manipulation. Researchers have studied how curvature drives intestinal crypt formation⁹⁴, and the importance of tissue geometry in organoid patterning¹⁴⁵. Other approaches in this space are discussed elsewhere¹⁴⁶.

4. HIOs in regenerative medicine

Intestinal failure (IF) occurs when patients do not receive appropriate nutrient and fluid absorption required either by anatomical limitations or loss of intestinal function¹⁴⁷. IF is a serious complication resulting from underlying GI diseases such as inflammatory bowel disease (IBD) or short bowel syndrome (SBS). Intestinal transplants are the most common treatment for severe cases, however, graft failure and low availability of transplantable organs remain significant challenges^148,149. In IBD patients, encompassing Crohn’s disease and ulcerative colitis, patients present with ulcers or wounds in the lining of the intestine. Although current treatments such as tumor necrosis factor alpha (TNF-α) inhibitors reduce inflammation levels, in most cases full closure of the intestinal wounds is not achieved¹⁵⁰. Advances in regenerative medicine utilizing HIOs, either iPSC-derived^6,8,151,152 or patient-derived^7,153, have showed promise in developing therapies to overcome challenging GI diseases. Two main approaches using intestinal organoids in tissue engineering have been pursued: (1) engineering a transplantable graft, and (2) repairing the injured intestine in vivo¹⁵⁴ (Fig. 6). For intestinal organoid therapy, it is important to consider scale-up processes and regulatory aspects to increase the relevance and clinical translation^49,155. In this section, we discuss current work in engineering new intestinal organoid therapies employing regenerative medicine principles.

Figure 6.

Tissue engineering approaches using HIOs: (i) therapeutic delivery of HIOs in vivo, and (ii) engineering a transplantable graft in vitro using decellularized tissue or degradable polymer scaffolds.

4.1. Engineering tissue grafts with HIOs

One repair approach, termed de novo, is when a transplantable graft is engineered in vitro and then transplanted to treat the injured or diseases tissue. In past years, efforts have focused two starting materials. One is engineering a tubular polymeric degradable scaffold populated with human intestinal cells, often termed tissue-engineered small intestine (TESI)^156–158. Another approach centers on using decellularized intestinal tissue repopulated with intestinal cells. The main challenge with engineering a transplantable graft is achieving a full thickness intestinal graft capable of adsorption, secretion, and food bolus movement¹⁵⁹. In addition, vascularization and nerve integration of the graft post-transplant are major hurdles that need to be overcome.

Finkbeiner et al¹⁵² compared both approaches to engineer a transplantable intestinal graft with (1) decellularized porcine small intestine and (2) a synthetic polyglycolic/poly-L-lactic acid (PGA/PLLA) hydrolytically degradable tubular construct, both seeded PSC-derived HIOs. For the first approach, they placed HIOs cut in half inside the luminal side of 3 mm × 3 mm squares of decellularized intestine and maintained them in culture for up to 4 weeks. The HIOs attached to the tissue with both epithelial and mesenchymal cells. However, after transplanting the grafts into the omentum or under the kidney capsule of immunocompromised mice, the areas with human cells were relatively small and expressing low levels of epithelial differentiation markers, suggesting poor graft survival. For the second approach, the authors observed that HIOs were capable of populating the PGA/PLLA tubular graft. After implantation in the omentum, the HIO PGA/PLLA grafts had high survival and intestinal epithelial architecture cells. However, these constructs lacked neuronal function important for complete integration. Another group¹⁶⁰, compared synthetic and natural polymer materials to engineer a TESI construct using mouse ISC organoids. The scaffolds they used were polyglycolic/poly-L-lactic acid (PGA/PLLA), polycaprolactone (PCL), and commercially available collagen based wound dressing material (CollaTape) coated with PLLA. All scaffolds supported in vitro organoid growth and in vivo organoid maturation after transplantation to the peritoneal cavity of immunodeficient rats. They observed PGA/PLLA and PCL scaffolds generated neomucosa with a maintained central lumen and comparable intestinal tissue architecture to the native intestine.

Other work has focused on repopulating segments of decellularized intestine maintained under perfusion in vitro before transplantation. Meran et al⁷ engineered human autologous jejunal mucosal grafts using decellularized human small intestine and colon segments, which were reseeded with intestinal enteroids from pediatric patients. After reseeding with these patient-derived organoids, tissue-engineered constructs were maintained in a combination of static and perfusion culture for 7 days before transplantation either under the kidney capsule or subcutaneously in immunodeficient mice. After 1-week post-implantation, the investigators observed initial signs of vascularization and presence of epithelial cells. The researchers proposed that grafts from cadaveric donors can be reseeded with autologous cell materials to overcome current challenges in the field. Kitano et al⁶ used iPSC-derived midgut spheroids, resulting from the initial stages of HIO differentiation, to engineer a humanized intestinal graft with a segment of decellularized rat intestine. After seeding the spheroids inside the lumen of the decellularized rat intestine, the intestinal segment was maintained in a whole organ bioreactor with media perfusion. They observed formation of a continuous epithelial monolayer and a mesenchyme layer under the regenerated epithelium in vitro. They transplanted this bioengineered tissue construct in a heterotopic position anastomosed to the carotid artery or jugular vein in immunocompromised rats where the graft survived over a period of 4 weeks.

4.2. In vivo delivery of intestinal organoids to injury models

The second approach in intestinal organoid tissue engineering centers on delivering intestinal cells directly to the injured area in vivo. One major hurdle with this approach lies in identifying a delivery method that supports cell survival, engraftment, and retention at the site of administration⁴⁹.

A proof-of-concept study demonstrating that organoids engraft into injured tissue was reported by Yui et al⁵. They transplanted LGR5+ mouse colonic organoids to the colon in a DSS-induced colitis murine model. To test the regenerative capacity of the intestinal organoids, they prepared a single cell suspension with ~500 organoids and delivered these via an intracolonic infusion. To prevent leakage from the single cell suspension, they glued the anal verge of each animal for at least 6 hours. The authors reported that the delivered cells engrafted to the epithelium of the injured colon as observed after 16 days post-delivery. This work motivated future studies using human intestinal organoids for tissue repair.

Sugimoto et al¹⁶¹ generated GFP-expressing human patient-derived colonic organoids and transplanted them to the injured colon of immunodeficient mice. The injury model consisted of an EDTA chemical injury combined with an electric toothbrush to mechanically remove the epithelium of a selected area of the mouse colon. Similarly to Yui et al⁵, the authors glued the anal verge to prevent the contents from spilling out after delivering the organoid suspension in a diluted Matrigel^™ solution. They observed that organoids successfully engrafted in the injured areas and noted that the resulting human crypts were larger than the host murine crypts. They continued monitoring the animals endoscopically for up to 6 months with positive GFP signal indicative of human engrafted cells.

Using a similar injury approach, Khalil et al¹⁵³ transplanted patient-derived intestinal organoids into a small intestine segment of immunodeficient mice following denuding of the epithelium via chemical and mechanical damage. Using clamps, they were able to target only a segment of the small intestine while it was still in continuity. They delivered the organoids in the injured region and used a fibrin-based sealant inside the lumen to maintain the organoids in place. After 1-week post-transplant, they observed cells positive for human markers in the delivered regions, however, they concluded that alternative transplantation techniques need to be developed since the sealant provided insufficient support for the engraftment of the donor cells.

The first demonstration of a synthetic biomaterial for in vivo delivery and engraftment of HIOs was provided by Cruz-Acuña et al^8,151. Using the same engineered hydrogel engineered for in vitro PSC-derived HIO development (PEG-4MAL), they showed effective endoscopic delivery of 10 HIOs per 50 μL hydrogel into focal mechanical injuries in the colonic mucosa of immunodeficient mice. The PEG-4MAL hydrogel crosslinking chemistry made it possible to inject a liquid hydrogel precursor into the wound that polymerized in situ. HIOs delivered with the PEG-4MAL hydrogel significantly increased the wound closure rate at the injury site, whereas HIOs delivered in saline did not improve wound healing. Remarkably, localized delivery of HIOs with PEG-4MAL hydrogel resulted in complete re-epithelialization of the intestinal wounds with human-derived tissue. This result highlights the importance of a biomaterial carrier for successful HIO therapies. This study represents a proof-of-concept study for an injectable synthetic hydrogel for therapeutic delivery and engraftment of HIOs to heal intestinal wounds.

4.3. Discussion

Collectively, these studies support the ability of either PSC-derived or patient-derived HIOs to populate decellularized tissue, polymeric biomaterial grafts, and in vivo injured tissue. As observed throughout the discussed work, cell adhesion to the material and cell engraftment to the intestinal tissue are important cellular events leading to a successful outcome of implanted grafts or in vivo hydrogel localized delivery. This process is mediated by cell-ECM receptors and other cell-cell mediated communication on the engraftment site¹⁶². Soluble biochemical signals, including growth factors and inflammatory cytokines and proteins, present in the host tissue also play a role dictating the cell engraftment process^163,164.

Decellularized grafts maintained in vitro under perfusion may have an advantage over those maintained at static culture since poor graft survival was observed with the later¹⁵². For transplantable grafts, one limitation in the discovery and optimization stage is the size of the animal model since this limits the engineered tissue size and the potential of orthoptic transplantation. Cell sourcing, whether patient-derived or PSC-derived HIOs, is a major consideration for these regenerative medicine approaches. Healthy tissue availability, batch-to-batch variability, and biomanufacturing costs have to be considered. Selecting between patient-derived and PSC-derived HIOs is important for in vivo delivery approach. Whereas patient-derived HIOs could promote epithelial regeneration, the mesenchymal layer in PSC-derived HIOs potentially offers additional tissue regeneration such as smooth muscle serving as a base for epithelial healing^49,153. The use of a biomaterial to retain the delivered HIOs to injured site in vivo is crucial for translating this technology to the clinic. An engineered synthetic biomaterial, such as injectable PEG hydrogels, is attractive since it has been demonstrated to be non-toxic, minimal lot-to-lot variability, and affording direct control over the mechanical and biochemical properties of the matrix. The PEG-4MAL hydrogel^127,151 approach is likely limited to small wounds based on the rapid gelation kinetics of the maleimide-thiol reaction. Delivery systems must be engineered to address clinically relevant wounds covering extensive areas of intestinal tissue.

Hydrogels have also been used in other intestinal applications as drug delivery vehicles^165,166, wound repair adhesives^167,168, and imaging agents¹⁶⁹. Hydrogel adhesion to the intestinal tissue, understanding degradation profile and mechanisms, and non-toxic byproducts are material properties that can be translated to engineering an in vivo intestinal organoid hydrogel vehicle.

5. Conclusion and future directions

PSC-derived and patient-derived HIOs hold great promise in the field of tissue engineering and regenerative medicine. Dependence on Matrigel^™ for HIO growth and delivery limits clinical translatability of the organoid technology due to translational/regulatory barriers, immunogenicity, and lot-to-lot variability. Synthetic hydrogels offer significant advantages in control over the matrix parameters (e.g., mechanical properties, bioactive ligand presentation, and degradability) and have been successfully deployed to generate both patient-derived and PSC-derived HIOs. For PSC-derived HIOs, Matrigel^™ is still used in the culture and differentiation of PSCs to spheroids in the planar culture. Adaptation of completely synthetic substrates to support the entire growth process will significantly increase the translation and manufacturing of the organoid technology. Additionally, increasing HIO culture complexity and evaluating the interplay of mesenchymal, endothelial, immune, and neural components grown within a synthetic hydrogel would provide a platform for robust in vitro HIO models. Recent advancements in tissue engineering with HIOs to treat intestinal failure focus on transplantable tissue grafts and organoid in vivo delivery. The use of a biomaterial is imperative in assessing current cell delivery challenges such as promoting localized cell engraftment. Future studies in the organoid therapy space should evaluate the immune system in the tissue-engineered constructs or in vivo cell engraftment using humanized animal models and relevant functional models Incorporation of synthetic hydrogels and HIOs into microfluidic gut-on-chip will provide additional insights to intestinal physiology and pathologies and serve as an in vitro platform for drug delivery and disease modeling.

Biographies

Andrés J. García is the Executive Director of the Petit Institute for Bioengineering and Bioscience and Regents’ Professor at the Georgia Institute of Technology. Dr. García’s research program integrates innovative engineering, materials science, and cell biology concepts and technologies to create cell-instructive biomaterials for regenerative medicine and generate new knowledge in mechanobiology. This cross-disciplinary effort has resulted in new biomaterial platforms that elicit targeted cellular responses and tissue repair in various biomedical applications, innovative technologies to study and exploit cell adhesive interactions, and new mechanistic insights into the interplay of mechanics and cell biology. In addition, his research has generated intellectual property and licensing agreements with start-up and multi-national companies. He is a co-founder of 3 start-up companies (CellectCell, CorAmi Therapeutics, iTolerance). He has received several distinctions, including the NSF CAREER Award; the Young Investigator Award, the Clemson Award for Basic Science, and the Founders Award from the Society for Biomaterials; the International Award from the European Society for Biomaterials; and Georgia Tech’s Outstanding Interdisciplinary Activities Award and the Class of 1934 Distinguished Professor Award. He is an elected Fellow of Biomaterials Science and Engineering (by the International Union of Societies of Biomaterials Science and Engineering), Fellow of the American Association for the Advancement of Science, Fellow of the American Society of Mechanical Engineers, and Fellow of the American Institute for Medical and Biological Engineering. He served as President for the Society for Biomaterials in 2018–2019. He is an elected member of the National Academy of Engineering, the National Academy of Medicine, and the National Academy of Inventors.

Adriana Mulero-Russe is currently pursuing her doctoral degree in bioengineering at the Georgia Institute of Technology where her research interests focus on engineering biomaterials, tissue engineering, and regenerative medicine. Adriana received the National Institutes of Health (NIH) Ruth L. Kirschstein Predoctoral Fellowship to support her PhD thesis work in Dr. Andrés García’s Lab. She obtained her bachelor of science degree in chemical engineering from the University of Puerto Rico at Mayagüez in 2018.