Mechanobiological engineering strategies for organoid culture

Abstract

Organoid culture systems have emerged as powerful platforms for studying development, disease modeling, and regenerative medicine. However, current models primarily rely on spontaneous self-organization within biomimetic matrices such as Matrigel, which lack precise control over biomechanical cues. Recent advances in mechanobiological engineering highlight the critical role of matrix-derived physical and mechanical properties—such as adhesion presentation, stiffness, viscoelasticity, and geometry—in directing organoid morphogenesis and functional maturation. This review explores how translating in vivo biomechanics into in vitro organoid culture strategies can overcome existing limitations, enhance reproducibility, and enable the development of physiologically relevant organoid systems.

I. BIOMECHANICS OF EXTRACELLULAR MATRIX STEER ORGANOID DEVELOPMENT THROUGH MECHANOTRANSDUCTION AXES

A. Current status of the organoid culture system and limitations

Organoids are three-dimensional multicellular structures that recapitulate key architectural and functional features of their corresponding organs. They can be generated from tissue-specific adult stem cells—often isolated from patient biopsy samples—or from pluripotent stem cells, including embryonic stem cells (ESCs), induced pluripotent stem cells (iPSCs), and tumor cells. Through spontaneous self-organization, organoids generate tissue-specific cell populations arranged in defined spatial domains and adopt relevant structural features, while exhibiting organ-specific functions (e.g., enzyme secretion, barrier formation, or electrophysiological activity). These hallmark features—multicellularity, tissue-specific cellular diversity, spatial organization, and functional readouts—serve as the criteria for organoid validation and maturation.

Over the past decade, organoid and assembloid technologies have emerged as advanced in vitro platforms for studying human development, disease modeling, drug screening, regenerative medicine, and personalized therapeutics.^1,2 Unlike traditional two-dimensional (2D) cultures, organoids retain key aspects of in vivo tissues, including genetic diversity, cellular heterogeneity, and complex cell-cell and cell-matrix interactions. They also reproduce spatial organization and physiological functions of native organs, providing unprecedented insights into organ development, tissue homeostasis, and disease progression.³ For instance, cerebral organoids model neurodevelopmental disorders,⁴ intestinal organoids elucidate gut microbiota-host interactions,^5,6 and liver organoids aid in studying fibrosis and metabolic disorders.^7,8

Despite these advances, the complexity of the in vivo environment makes replicating it in vitro a challenging task. The transition from stem cells to fully developed tissues is orchestrated by tightly regulated biochemical and biomechanical signals within the stem cell niche.⁹ This dynamic microenvironment provides biochemical factors (e.g., growth factors, cytokines) and biomechanical cues (e.g., stiffness, topography, shear stress) that direct cell fate and morphogenesis (Fig. 1).¹⁰ Emerging evidence highlights the critical role of biomechanical cues alongside biochemical factors. For example, mesodermal stiffening during neural crest development facilitates directed cell migration.^10,11 Similarly, the extracellular matrix (ECM), a key component of the niche, undergoes continuous remodeling during development, with its composition, stiffness, and architecture finely tuned in a tissue-specific manner.^10,11 These biomechanical cues are sensed by stem cells through mechanotransduction, translating mechanical signals into biochemical responses that regulate gene expression and cell behavior.

FIG. 1.

Mechanobiological engineering strategies bridge in vivo biomechanics with in vitro organoid culture. By fine-tuning matrix properties—adhesion ligand, stiffness, viscoelasticity, and geometry—we can recapitulate the native stem cell niche, driving more physiologically relevant and functionally mature organoids. Created with Biorender.com.

In organoid culture, recreating an appropriate ECM niche is essential for directing stem cell behavior and promoting organoid maturation. Biomimetic matrices such as Matrigel and decellularized ECM (dECM) hydrogels have been widely adopted to support survival, proliferation, and morphogenesis. However, these matrices exhibit batch-to-batch variability, undefined compositions, and limited tunability in mechanical properties, compromising reproducibility and physiological relevance.¹² Moreover, organoid formation largely depends on stochastic self-organization, leading to heterogeneity in size, shape, and cellular composition. For instance, the number and positioning of crypt-like domains in intestinal organoids remain difficult to control, posing challenges for standardization and scalability in both research and translational applications.

While significant efforts have focused on optimizing biochemical signals in organoid culture systems, the importance of biomechanical cues has only recently gained attention. Given that ECM biomechanics evolve throughout development and profoundly influence stem cell fate, engineering the mechanical properties of the culture microenvironment offers a promising strategy to enhance organoid maturation and functionality. Incorporating precise mechanical stimuli can improve structural organization, reproducibility, and physiological relevance in organoid models.

This review highlights the emerging role of biomechanical cues in organoid development and explores mechanobiological engineering strategies aimed at recreating dynamic tissue-like microenvironments. By leveraging matrix-driven mechanotransduction to guide morphogenesis, these approaches have the potential to overcome current limitations and enable the next generation of robust, translationally relevant organoid systems.

B. ECM-guided mechanobiological regulation in organ development in vivo

During embryogenesis and tissue morphogenesis, cells are regulated not only by biochemical signals but also by dynamic biomechanical cues from the ECM and surrounding tissue environment. Through continuous interactions with the ECM, cells sense and respond to these physical signals, which orchestrate proliferation, differentiation, migration, and morphogenesis. These biomechanical cues are highly context-dependent, varying by tissue type and developmental stage, and are essential for tissue homeostasis and organ formation.

In adult tissues, cells experience organ/tissue-specific mechanical environments, including intrinsic mechanical properties of the ECM such as stiffness, porosity, and viscoelasticity, and extrinsic mechanical forces (for example, compressive loading in cartilage, tensile stretching in skin, and cyclic strain in blood vessels). Such specific mechanical contexts govern force transmission and cellular mechanosensing, enabling tissue homeostasis under normal conditions or, when dysregulated, driving disease. For instance, excessive collagen deposition and ECM stiffening are early hallmarks of fibrosis and cancer, where altered mechanics promote abnormal cell behaviors such as enhanced migration, invasion, and unchecked proliferation.^13,14

Similarly, during development, dynamic intrinsic and extrinsic mechanical forces act as key regulators of organogenesis. For instance, mesodermal stiffening beneath the neural crest during gastrulation triggers epithelial-to-mesenchymal transition (EMT) and collective cell migration.¹⁵ In the intestine, compressive stress from the mesenchyme and smooth muscle drives villus folding, while differential actomyosin contractility and apical constriction are implicated in crypt morphogenesis.^16–22 Epithelial tubulogenesis in organs such as the lung and intestine is likewise regulated by mechanical constraints imposed by differentiating smooth muscle.^23,24 Although the role of biomechanical regulation in in vivo morphogenesis and differentiation is increasingly recognized, the underlying mechanisms remain an active area of research.

At the cellular interface, cells interpret ECM-derived mechanical cues through “mechanotransduction”—the conversion of physical signals into biochemical responses. Transmembrane receptors such as integrins and syndecans link the ECM to the cytoskeleton, initiating focal adhesion assembly via adaptor proteins (e.g., talin, vinculin).^25–27 These complexes facilitate cytoskeletal remodeling and force transmission through the linker of the nucleoskeleton and cytoskeleton (LINC) complex, ultimately influencing nuclear structure and gene expression. Downstream activation of mechanosensitive signaling pathways, including YAP/TAZ,^28,29 Wnt-β-catenin,^30–35 and MAPK/ERK,^36–38 regulates key cellular processes such as proliferation, differentiation, and migration. By responding to mechanical cues, cells fine-tune their behavior to adapt to the changing physical properties of their microenvironment during development.

Traditionally, mechanotransduction signaling studies have relied on 2D polyacrylamide (PAA) gels,³⁹ which allow precise stiffness tuning but lack the spatial and temporal complexity of native tissues. In 2D, substrate stiffness often dominates cell behavior: increasing stiffness enhances actomyosin contractility, driving cells to spread and flatten—behaviors that inversely correlate with cell volume. Importantly, cell morphology and volume adaptation in 3D environments differ fundamentally from 2D conditions.⁴⁰ In 3D, the surrounding matrix imposes geometric confinement, so cells cannot simply spread but must remodel their surroundings to change shape and volume. In this context, not only stiffness but also matrix viscoelasticity and degradability govern how cells extend protrusions, generate traction forces, and ultimately regulate volume, all of which feedback on downstream signaling pathways.^41–43 Consequently, in 3D, matrix dynamics influence fate specification more strongly than stiffness alone. These differences underscore the necessity of studying mechanotransduction in physiologically relevant 3D contexts to better understand in vivo development and to guide the design of next-generation biomaterials and bioengineering strategies for organoid culture.

Building on insights from PAA substrates, we now turn to specialized mechanomodulatory platforms designed for 3D organoid culture. In this review, we emphasize recent platforms including polyethylene glycol (PEG)-based hydrogels with dynamic presentation of adhesion ligands and tunable stiffness, alginate- and DNA-based hydrogels with programmable viscoelasticity, and photo-responsive hydrogels that enable spatiotemporal control of mechanical properties. While other reviews have surveyed the full landscape of mechanobiology materials,⁴⁴ here we focus exclusively on platforms engineered to recreate the dynamic mechanical cues essential for organoid formation and function.

C. Biomimetic matrices for organoid culture in vitro

Matrigel, a basement membrane extract derived from Engelbreth-Holm-Swarm mouse sarcoma, has long been the gold standard matrix for organoid culture due to its rich composition of ECM proteins, growth factors, and bioactive molecules.^45–47 Its main components, laminin and type IV collagen [Fig. 2(a)], play crucial roles: laminin binds integrin receptors (e.g., α3β1, α6β1, α6β4, α7β1) to regulate adhesion, migration, proliferation, and differentiation, while type IV collagen self-assembles into a supramolecular network that provides mechanical stability and scaffolding,^48,49 thereby driving organoid formation. However, since Matrigel is tumor-derived and biologically undefined, it lacks tissue specificity, potentially compromising the physiological relevance of organoid models.^50,51 Additionally, Matrigel exhibits considerable batch-to-batch variability and limited mechanical tunability, with a narrow stiffness range of ∼20–450 Pa, making it difficult to recapitulate the dynamic and diverse mechanical environments of native tissues.^52,53 These shortcomings hinder the reproducibility and standardization of organoid cultures and limit their applicability in disease modeling and regenerative medicine.^3,10,11,54

FIG. 2.

Tissue specificity of decellularized ECM hydrogels for organoid culture in vitro. Decellularized ECM (dECM) serves as a promising alternative to Matrigel by offering tissue/organ-specific biochemical and biomechanical cues. (a) Proteomic analysis of dECM derived from gastrointestinal tissues [Adapted with permission from Kim et al., Nat. Commun. 13(1), 1692 (2022). Copyright 2022 Nature Publishing Group⁵⁷], revealing distinct protein compositions compared to Matrigel. (b)–(e) Human brain organoids cultured using brain-derived decellularized ECM (BEM) in a microfluidic device [Adapted with permission from Cho et al., Nat. Commun. 12(1), 4730 (2021). Copyright 2021 Nature Publishing Group⁶⁰]. (b) Workflow schematic for generating organoids within the device. (c) and (d) Proteomic analysis reveals that BEM is enriched in brain-specific proteins, closely resembling native brain tissues and distinct from Matrigel. (e) Gene ontology analysis of BEM-enriched proteins highlights their roles in nervous system development and neurogenesis. (f) Immunohistochemical staining of neural progenitor marker (Nestin) and neuronal markers (Tuj1 and MAP2) at day 30 of culture. (g) Image-based quantification of the Tuj1- and MAP2-positive areas in the Matrigel- vs BEM-cultured organoids at day 30, demonstrating that brain-derived decellularized ECM enhances neurogenesis compared to Matrigel.

To overcome these challenges, dECMs have emerged as promising alternatives for organoid culture. dECMs are derived from specific tissues or organs and retain a more defined biochemical composition, including tissue/organ-specific ECM proteins, glycoproteins, and other matrix-bound factors essential for organ-specific function.^27,55,56 Compared to Matrigel, dECMs offer improved biological relevance, reduced variability, and enhanced capacity to support organoid development by providing organ-specific biochemical cues. Moreover, their high collagen content—predominantly fibrillar collagens I - IV dictated by tissue source—yields mechanical microenvironments more akin to native tissues than the low-collagen Matrigel [Fig. 2(a)].^57–59 For example, brain-derived dECMs contain approximately 90 brain-specific matrisome proteins, over 94% of which are also found in normal human brain tissue. In contrast, Matrigel lacks many of these neural-specific ECM components [Figs. 2(b)–2(e)]. Consequently, brain-derived dECM hydrogels have been shown to promote the formation of neural organoids with more accurate cortical organization and improved electrophysiological maturation, outperforming Matrigel in both structural fidelity and functional outcomes [Figs. 2(f) and 2(g)].⁶⁰

Despite their advantages, dECM hydrogels still struggle with precise mechanical tuning, particularly when targeting higher stiffness. The elasticity of dECM hydrogels is typically low (0.2–1.5 kPa),^61,62 and extending their stiffness beyond tens of kPa by simply adjusting concentration or processing conditions remains challenging.^59,61 To enhance mechanical strength, straightforward chemical crosslinkers such as genipin or glutaraldehyde can be applied,^63–65 while photo-reactive modifications (e.g., methacrylation)^66,67 or blending with photo-crosslinkable polymers (e.g., methacrylated gelatin, methacrylated hyaluronic acid)⁶⁹ have been developed to boost bulk stiffness. For example, methacrylated cartilage-derived dECM can reach a compressive modulus of 1070 ± 150 kPa, comparable to that of native cartilage, and thus promote chondrogenic differentiation of stem cells.⁶⁷ However, few studies have yet translated these mechanically reinforced dECM hydrogels into organoid culture, and further optimization of viscoelasticity and degradability will be essential to realize their full potential.

Another important hurdle is the variability and scalability inherent to the decellularization process. Donor-to-donor and batch-to-batch differences can alter matrix composition and compromise mechanical integrity. While dECMs generally exhibit less variability than Matrigel,⁵⁷ establishing standardized and automated decellularization protocols will be critical to producing consistent, high-quality hydrogels at scale.

Together, these challenges highlight the need for next-generation mechanobiological engineering strategies that seamlessly integrate organ-specific biochemical signals with fine-tuned, dynamically adjustable mechanical properties. In Sec. II, we explore emerging mechanomodulatory biomaterials and bioengineering approaches—combining controlled stiffness, viscoelasticity, degradability, and spatial patterning—that aim to overcome these limitations and advance organoid culture toward more reproducible and physiologically relevant models.

II. MECHANOBIOLOGICAL ENGINEERING STRATEGIES FOR ORGANOID CULTURE

Biomechanical cues from the ECM play a crucial role in regulating stem cell behavior and guiding organ development in vivo. These mechanotransductive signals—mediated by ECM composition, stiffness, viscoelasticity, and spatial organization—govern key processes such as proliferation, differentiation, and morphogenesis. Building on these insights, mechanobiological engineering approaches are increasingly being used to create more physiologically relevant and functionally mature organoids in vitro.

This section highlights recent strategies that leverage matrix-driven physical and mechanical cues to control organoid formation. We focus on how modulating adhesive ligand presentation, matrix stiffness, viscoelastic properties, and geometry within 3D matrices can recapitulate critical aspects of the native stem cell niche. By fine-tuning these parameters, researchers have enhanced organoid morphogenesis, improved tissue organization, and promoted functional differentiation. In addition, we discuss the integration of these engineered microenvironments into dynamic culture systems that deliver spatially and temporally controlled mechanical signals, bridging the gap between conventional organoid models and the complex biomechanical conditions of in vivo development.

A. Adhesive ligand

A fundamental aspect of cell-ECM communication is the interaction between adhesion receptors—most prominently integrins—on the cell surface and specific peptide or glycan motifs within the ECM, collectively referred to as cell-adhesive ligands. These interactions are central to mechanotransduction—the process by which cells sense and respond to the mechanical properties of their microenvironment. Mechanistically, integrins undergo conformational changes through inside-out signaling, increasing their affinity for ECM motifs such as RGD (Arg-Gly-Asp) in fibronectin and fibrinogen, GFOGER in collagen, and IKVAV in laminin.⁶⁸ Once bound to these motifs, integrins cluster to form nascent adhesion, which matures into focal adhesions that link to the actin cytoskeleton.^69,70 This physical coupling facilitates the transmission of mechanical forces across the plasma membrane and into the nucleus via the LINC complex, ultimately modulating gene expression at both transcriptional and epigenetic levels.

In vivo, adhesive ligands are presented in highly specific spatiotemporal patterns that mirror the evolving demands of stem cell niches. For example, during embryonic implantation, integrin engagement shifts in a coordinated sequence: early on, αvβ3 mediates blastocyst attachment to vitronectin, laminin, and fibronectin in the uterine stroma; as implantation proceeds, integrins α1β1, α2β1, α6β1, and α7β1 become predominant, facilitating trophoblast migration and invasion through laminin-rich environments.^71–73 This dynamic integrin switching highlights how adhesive ligand specificity guides cell behavior in response to microenvironmental changes.

Remarkably, similar spatiotemporal precision emerges in organoid systems through cell-mediated ECM secretion and deposition. In intestinal organoids, Paneth cells assemble their own basement-membrane ECM, localizing integrin β4 to the basal surface and stabilizing epithelial attachment.⁷⁴ Likewise, in pancreatic cancer organoids, tumor cells gradually increase laminin secretion, driving a temporal shift in integrin usage from fibronectin-binding integrins (αVβ3 and αVβ5) to laminin-binding integrins (α6β1, α6β4, and α3β) over time.^75,76 In some cases, cells encapsulated in biologically inert, non-adhesive hydrogel platforms (e.g., alginate- or PEG-based hydrogels) deposit their own ECM, which alone can support organoid formation.^77,78 These examples demonstrate that organoid cultures can intrinsically recapitulate the in vivo-like dynamics of adhesive ligand-integrin interactions.

By harnessing the specificity and avidity of integrin-ligand interactions, organoid development can be actively guided and fine-tuned. For example, pancreatic niche-inspired ECM enriched in type V collagen (Col V) binds integrins α1β1 and α2β1 via its WWASKS peptide motif, thereby activating the canonical Wnt/β-catenin pathway to drive endocrine cell differentiation and spatial organization of α, β, δ, and PP cells.^56,79 This targeted signaling enhances the functional maturation of pancreatic organoids, resulting in glucose-responsive insulin and glucagon secretion. Integrin engagement also directs cell polarity and overall organoid architecture. In epithelial organoids, activation of β-integrin receptors by ECM ligands at the basal cell surface promotes laminin secretion and initiates intracellular signaling cascades that establish apical-basal orientation.⁸⁰ Conversely, removal of the matrix or pharmacological blockade of β-integrin receptor signaling in biliary organoids leads to an inversion of apical-basal polarity—exposing the apical domain externally, which enables the study of apical events that are otherwise inaccessible in conventional Matrigel-embedded cultures.⁸⁰

Recently, leveraging the modularity of PEG-based hydrogels has enabled systemic dissection of the cell-ECM parameters—ligand specificity, presentation, and density—that underpin robust organoid formation. Because PEG resists nonspecific protein adsorption, defined adhesive cues must be deliberately introduced: PEG backbones are functionalized with either full-length ECM proteins or minimal peptide motifs to probe which ligands best drive organoid initiation and expansion. For example, in liver organoid cultures, PEG hydrogels presenting fibronectin- or laminin-111 supported 3D growth and gene expression profiles comparable to Matrigel, whereas type IV collagen or RGD motifs performed less well (Fig. 3).⁸¹ Notably, substituting full-length fibronectin with its minimal RGDSPG peptide produced similar outcomes, confirming the sufficiency of minimal adhesive ligands in promoting progenitor cell expansion and cystic organoid morphogenesis.^81,82 In intestinal organoid cultures, RGD-functionalized PEG hydrogels effectively supported stem cell colony formation but failed to induce complex budding. Only hydrogels containing laminin-111, rather than isolated laminin peptides, induced the full budding morphogenesis, suggesting that intact ECM proteins may provide additional structural or biochemical cues essential for intestinal organoid development.⁸² More precisely, collagen-derived GFOGER peptides (which bind α2β1 integrins) markedly outperformed fibronectin-derived PHSRN-K-RGD peptides (α5β1 ligand) in human intestinal and endometrial organoid cultures.^83,84 GFOGER-functionalized PEG hydrogels yielded higher enteroid-forming efficiency, larger organoids, and sustained viability, polarity, and differentiation across donors. By contrast, PHSRN-K-RGD-functionalized hydrogels exhibited only limited support. Hydrogels incorporating the GFOGDR variant—a peptide with a single amino acid substitution that abolishes α2β1 integrin binding—failed to sustain cell survival and enteroid formation. These findings pinpoint that α2β1 integrin engagement is a critical determinant of epithelial organoid development.⁸⁴

FIG. 3.

Adhesion ligand-dependent organoid formation. (a) and (b) Representative bright-field images showing liver organoid formation in Matrigel (MG), plain PEG (PEG), and PEG functionalized with various ECM components (COL. IV, collagen IV; FN, fibronectin; LAM-1, lamin-1; RGD, RGD peptide). (c) Hematoxylin and eosin staining of liver organoids showing cystic structures with a central lumen surrounded by epithelial cells in both Matrigel- and RGD-functionalized PEG (PEG-RGD) hydrogels. (d) Heatmap depicting the expression levels of liver-specific genes. (e) Immunostaining for liver-specific markers in organoids cultured in Matrigel- and PEG-RGD hydrogels. These findings suggest that organoids cultured in PEG-RGD hydrogels display a progenitor phenotype expressing stem/ductal markers and exhibit morphology and gene expression profiles comparable to those grown in Matrigel [Panels (a)–(e) adapted with permission from Sorrentino et al., Nat. Commun. 11(1), 3416 (2020). Copyright 2020 Nature Publishing Group⁸¹]. (f) and (g) Formation of lumens and morphology of human induced pluripotent stem cell (hiPSC) clusters are influenced by the density of RGD peptide [Adapted with permission from Indana et al., Adv. Mater. 33(43), 2101966 (2021). Copyright 2021 John Wiley & Sons, Ltd.⁸⁵]. (f) Representative images of hiPSC clusters stained with the membrane dye R18 on day 7 of culture in hydrogels with varying RGD densities (scale bar = 100 μm). (g) Quantification of clusters with and without lumens on day 7 across different RGD conditions, indicating that higher RGD densities and faster stress relaxation promote lumen formation.

In addition to ligand specificity, ligand density profoundly impacts mechanosensation and downstream cellular behavior.^86,87 Ligand density influences integrin clustering, focal adhesion formation, cytoskeletal tension, and ultimately stem cell fate. Studies have shown that increasing ligand density can modulate the nuclear localization of Yes-associated protein (YAP), a key mechanotransducer. At low ligand density, YAP remains cytoplasmic (Nuc./Cyt. YAP of around 0.6 for both soft and stiff substrates, with no statistically significant difference between the groups), while intermediate densities promote YAP nuclear translocation in response to matrix stiffness (Nuc./Cyt. YAP of around 0.8 for soft was increased to around 1.8 for stiff substrate).⁸⁶ At high ligand densities, YAP translocated to the nucleus regardless of substrate stiffness (Nuc./Cyt. YAP of around 2 for both soft and stiff substrates, with no statistically significant difference between the groups), demonstrating that ligand density can override mechanical inputs to regulate cell behavior. Furthermore, higher ligand densities enhance cell spreading, promote robust F-actin stress fiber formation, and facilitate osteogenic differentiation—even on substrates with lower stiffness. These findings emphasize that optimizing ligand density is essential for achieving desired mechanobiological outcomes in organoid cultures.

Importantly, in a 3D context, adhesive ligand presentation is inherently dynamic—local ligand availability and clustering change as cells remodel their surroundings and cluster integrins. In the case of polymeric hydrogel platforms, because adhesive ligands are covalently tethered to the polymer backbone, the network structure and the polymer conformation critically influence both ligand accessibility and clustering. For instance, in PEG-based hydrogels, the architecture of the PEG network and the conformation of PEG arms in aqueous solution not only govern bulk stiffness but also dictate how readily ligands engage integrins, directly impacting organoid yield.⁸⁸ Likewise, in alginate hydrogels, longer alginate chains reduce ligand accessibility and clustering, which limits ECM remodeling and lumen formation in organoids.⁸⁵ Those same backbone characteristics also modulate bulk mechanics—longer alginate chains slow stress-relaxation, hindering organoid development, although increasing ligand density can partially rescue this effect.⁸⁹ These examples illustrate that adhesive ligands do not act in isolation but synergize with matrix mechanics; accordingly, the co-design of biochemical cues and physical properties is essential to optimize mechanobiological environments for organoid culture.

Together, these findings underscore the critical role of adhesive ligand specificity, density, and distribution in directing organoid development. As our understanding of adhesive ligand-mediated mechanotransduction deepens, future biomaterial designs can incorporate spatiotemporally controlled ligand presentation and dynamic remodeling capabilities, advancing the field of organoid engineering toward more sophisticated and clinically translatable models. Table I presents the latest publications examining how adhesive ligands affect organoid development.

TABLE I.

Adhesion ligands impact organoid development.

Type	ECM or peptide motif	Mechanosensitive pathway signaling	Cellular outcome	Ref.
Pancreatic organoid	WWASKS	Wnt/β-catenin	• Drives endocrine cell differentiation and spatial organization of α, β, δ, and PP cells • Enhances functional maturity	56, 79
Liver organoid	RGDSPG	Integrin/SFK/YAP	Promotes progenitor cell expansion and cystic organoid morphogenesis	81
Intestinal organoid	RGD	YAP	Supports stem cell colony formation	82
	RGD + Laminin-111		Induces budding morphogenesis

B. Matrix stiffness

Matrix stiffness is one of the most fundamental mechanical properties of the ECM, profoundly influencing cellular behavior and tissue development. Defined by Young's modulus, which measures the ratio of applied stress to the resulting strain,⁹⁰ stiffness plays a central role in directing cell fate decisions, including proliferation, differentiation, and morphogenesis.

In native tissues, stiffness varies widely, from the soft brain parenchyma (on the order of a few pascals) to the rigid cortical bone matrix (reaching several gigapascals).⁹¹ These tissue-specific mechanics are not merely passive backdrops but actively guide cellular behavior: cells sense and respond to local rigidity via integrin-cytoskeleton linkages, triggering downstream signaling cascades that control gene expression and ultimately impacting cellular function and tissue homeostasis. Aberrant changes in tissue stiffness, such as softening in keratoconus or stiffening in fibrotic tissues and solid tumors, are often early indicators and drivers of disease progression,⁹² underscoring the critical role of stiffness in maintaining physiological function.

Recent studies have highlighted matrix stiffness as a dynamic and instructive cue during embryonic development and organogenesis. Spatial variations in cell density, coupled with ECM remodeling, generate localized stiffness gradients that guide morphogenic processes in vivo.^{11,15,93–95} For example, during gastrulation in Xenopus embryos, convergent extension movements increase the local stiffness of the head mesoderm through planar cell polarity signaling.¹⁵ Neural crest cells sense this localized stiffening via an integrin–vinculin–talin complex, with integrin β1 serving as the primary receptor. Once mesoderm stiffness surpasses a critical threshold, these mechanosensory inputs trigger the EMT transition, promote Piezo1-mediated microtubule modifications, and facilitate collective cell migration,⁹³ which is further guided by chemotactic signals like SDF-1 to drive morphogenesis.

Beyond cell-density-driven tissue stiffening, cells can actively modulate the stiffness of their microenvironment through contractile forces.⁹⁶ Cells exert traction forces in the nano-newton range^97,98 on ECM components like collagen, which exhibits strain-stiffening behavior. By engaging integrin β1 subunits, cells remodel collagen fibrils to locally enhance ECM stiffness, reinforcing structural integrity and creating stiffness gradients that guide directional migration and morphogenesis.^99–102 These dynamic, reciprocal interactions between cells and the ECM demonstrate that tissue stiffness is not a passive property but an active regulator of development.

Recent advances in biomaterials and mechanobiological engineering have led to the development of stiffness-tunable hydrogels, providing powerful platforms to systemically investigate how matrix stiffness directs organoid development and function.⁴⁴ These engineered materials enable fine control over mechanical properties and allow researchers to decouple stiffness from other matrix factors, offering clearer mechanistic insights. While these materials can achieve stiffnesses ranging from a few pascals to gigapascals, cells are typically only responsive to limit stiffnesses between 0.1 kPa and several tens of kPa.^92,103–105 Consequently, most studies focus on tuning material stiffness within this physiologically relevant range.¹⁰⁶

Polyacrylamide (PAA) hydrogels, for example, have been widely used as 2D culture substrates due to their tunable stiffness, spanning several orders of magnitude. They are synthesized through radical polymerization of acrylamide and bis-acrylamide, with their stiffness finely tuned by adjusting the ratio and concentrations of these monomers. To facilitate cell adhesion, the hydrogel surface is typically functionalized with adhesion ligands, such as RGD peptides or ECM proteins, using a hetero-bifunctional crosslinker like sulfo-SANPAH.

Early work with 2D PAA hydrogels coated with type IV collagen has demonstrated that matrix stiffness directly governs the balance between stemness and differentiation in intestinal organoids.¹⁰⁷ Softer matrices (∼0.6 kPa), resembling the native basement membrane, maintain LGR5⁺ intestinal stem cells (ISCs) and promote crypt formation by facilitating cytoplasmic YAP localization. In contrast, stiffer substrates (∼9.6 kPa) trigger nuclear YAP translocation, suppress Wnt signaling, and shift differentiation toward goblet cells (MUC2⁺), reducing enterocytes (VILLIN⁺) populations. This research demonstrated that YAP was identified as a key mechanotransducer, with its inhibition or expression reversing or mimicking stiffness-induced effects on ISC fate and cryptogenesis.

Additionally, another study revealed that matrix stiffness influences crypt architecture and epithelial compartmentalization using PAA hydrogel systems.¹⁰ In stiffer substrates (∼15 kPa), the stem cell compartment protrudes from the epithelial monolayer, while softer substrates (0.2–0.7 kPa) promote more pronounced folding. These changes are driven by stiffness-dependent radial traction forces regulated by cortical actomyosin. Increased stiffness reduces both the size and number of stem cells, through the stem-to-Paneth cell ratio remains stable. Together, these findings underscore the role of matrix stiffness in coordinating ISC behavior and intestinal tissue morphogenesis.

While PAA hydrogels have provided valuable insights into the role of stiffness in organoid development, they are suboptimal for 3D organoid cultures due to the cytotoxic nature of their monomers, crosslinkers, and cross-linking mechanisms. Additionally, their purely elastic and non-degradable properties restrict cell-mediated matrix remodeling, limiting their ability to replicate the spatiotemporal dynamics of the ECM, and may introduce biases in cellular outcomes.

To overcome these limitations, recent studies have focused on developing bioinert synthetic hydrogels that enable precise modulation of matrix stiffness in 3D culture. For example, Lutolf and colleagues engineered PEG hydrogels that are enzymatically crosslinked by activated transglutaminase factor XIIIa (FXIIIa).^81,82 FXIIIa, a transglutaminase enzyme essential for fibrin clot formation during tissue injury,¹⁰⁸ catalyzes acyl-transfer reactions between the α-carboxamide group of protein-bound glutaminyl (Gln) residues and the ε-amino group of lysyl (Lys) residues, forming ε-(α-glutamyl) lysine isopeptide crosslinks.¹⁰⁹ Inspired by this natural chemistry, eight-arm PEG macromers were functionalized with complementary transglutaminase peptide substrates (Gln and Lys), enabling mild cell encapsulation via gelation at 37 °C and neutral pH in the presence of FXIIIa.¹¹⁰

For intestinal organoid formation, mouse ISCs encapsulated within PEG hydrogels modified with RGD ligands demonstrated stiffness-dependent proliferation and morphogenesis.⁸² An optimal stiffness window (∼1.3 kPa) promoted robust ISC expansion and crypt-like structure formation, while softer matrices (∼0.3 kPa) failed to support proliferation. Mechanically, matrix stiffness regulated YAP localization via an integrin-actomyosin-YAP axis. Inhibiting contractility or integrin binding disrupted colony formation, underscoring the requirement for mechanical tension and YAP activity in organoid development.

Crucially, the PEG platform also allows control of matrix degradability, providing insights into how dynamic stiffness remodeling influences organoid development. Organoids formed efficiently only in degradable hydrogels with stiffness around 190 Pa.⁵¹ Rapid degradation of stiff matrices induced inflammation-like responses, impairing stem cell maintenance. These findings suggest that organoid development is confined to a narrow mechanical window, highlighting the critical interplay between matrix stiffness and dynamic remodeling.

Similar mechanosensitive responses to matrix stiffness have been observed in liver organoids cultured in PEG-based hydrogels (Fig. 4). Organoid proliferation was optimized at physiologically relevant stiffness values (1.3–1.7 kPa), closely mimicking native liver tissue mechanics.^81,82 Interestingly, while matrix stiffness strongly influenced progenitor proliferation, it had minimal impact on differentiation potential, as hepatic marker expression remained consistent across conditions. YAP signaling was again identified as a key mediator, with nuclear localization and target gene expression increasing at higher stiffness levels. In liver organoids, stiffness-regulated proliferation occurred via an integrin-dependent, but actomyosin-independent mechanism, mediated instead by Src family kinase (SFK)-driven YAP activation. This alternative pathway demonstrates the tissue-specific diversity of mechanotransduction mechanisms and highlights the importance of tailoring mechanical inputs to different organoid systems.

FIG. 4.

Matrix stiffness-dependent organoid formation. (a) Liver organoid yield increases proportionally with matrix stiffness, with optimal growth observed in matrices mimicking physiological liver stiffness (∼1.3–1.7 kPa) [Adapted with permission from Sorrentino et al., Nat. Commun. 11(1), 3416 (2020). Copyright 2020 Nature Publishing Group⁸¹]. (b) Expression levels of YAP target genes are upregulated in liver organoids cultured in stiffer matrices [Adapted with permission from Sorrentino et al., Nat. Commun. 11(1), 3416 (2020). Copyright 2020 Nature Publishing Group⁸¹]. (c) and (d) The integrin/SFK/YAP signaling pathway mediates liver progenitor proliferation in response to increased matrix stiffness (PF, PF-573228; BLEB, blebbistatin; VP, verteporfin; DAS, dasatinib) [Adapted with permission from Sorrentino et al., Nat. Commun. 11(1), 3416 (2020). Copyright 2020 Nature Publishing Group⁸¹]. (e) and (f) Cyst formation by normal rat cholangiocytes (NRCs) cultured in PEG-RGD hydrogels of varying stiffness, ranging from soft (0.5 kPa) to stiff (7 kPa) [Adapted with permission from Funfak et al., Front. Bioeng. Biotechnol. 7, 417 (2019). Copyright 2019 Frontiers¹¹¹]. (e) Representative bright-field images and (f) quantification of NRC cysts at day 10, cultured in PEG-RGD hydrogels with varied stiffness and adhesion ligand density.

Collectively, these studies underscore the critical role of matrix stiffness as an instructive cue in organoid culture, influencing stem cell expansion, differentiation, and tissue organization. Mechanobiological engineering strategies that precisely tune stiffness and matrix remodeling enable the recreation of physiologically relevant microenvironments that promote organoid maturation and function. However, as an active area of research, the precise molecular mechanisms linking matrix stiffness to organoid morphogenesis in 3D, beyond well-studied YAP signaling pathways, remain incompletely understood, and further in-depth investigation of alternative mechanotransducers, integrin-cytoskeleton crosstalk, and integration with biochemical signals will be essential. Table II summarizes recent work on how biomatrix stiffness influences organoid outcomes.

TABLE II.

Effect of matrix stiffness on organoid development.

Type	Platform	Adhesion ligand	Stiffness	Findings	Ref.
Intestinal organoid	PAA	Collagen IV	0.6 vs 9.6 kPa	YAP regulates stiffness-induced effects on ISC fate and cryotogenesis.	107
Intestinal organoid	PAA	Collagen I + laminin-1	0.2–0.7 vs 15 kPa	Crypt architecture and epithelial compartmentalization are driven by stiffness-dependent radial traction forces regulated by cortical actomyosin.	10
Intestinal organoid	PEG	RGD	0.3 vs 1.3 kPa	The integrin-actomyosin-YAP axis regulates the stiffness-dependent proliferation and morphogenesis of ISCs.	82
Liver organoid	PEG	RGD or RGD/Laminin-111	1.3–1.7 vs 4 kPa	Integrin/SFK/YAP signaling regulates liver progenitor proliferation.	81

C. Matrix viscoelasticity

Biological tissues are inherently viscoelastic, exhibiting both elastic (solid-like) and viscous (fluid-like) mechanical behavior. This time-dependent property allows tissues to deform under mechanical stress and gradually dissipate the applied energy, providing a dynamic mechanical environment that profoundly influences cell behavior. Soft tissues such as the liver, breast, and skin can relax their stress over seconds to minutes, while even stiffer tissues like cartilage, bone, and tendon display measurable viscoelastic responses. Importantly, viscoelasticity is not only a characteristic of mature tissues but also plays a crucial role during embryonic development, where dynamic tissue remodeling and force generation drive morphogenesis.¹¹²

Cells continuously interact with their surrounding ECM by exerting forces ranging from piconewtons to hundreds of nanonewtons, with timescales spanning milliseconds to hours. Given this dynamic interplay, the viscoelastic nature of the ECM plays a crucial role in modulating cell-matrix interactions and mechanotransduction.^113–115 Unlike purely elastic materials, viscoelastic matrices exhibit an initial elastic response to applied stress, followed by gradual stress relaxation over time. This time-dependent mechanical behavior enables the matrix to either store or dissipate cellular forces, directly influencing cell behavior by regulating cytoskeletal tension, focal adhesion dynamics, and nuclear mechanosensing, all of which are essential for guiding cell behavior and fate decisions.

While most mechanobiology research has focused on ECM stiffness as a static parameter, emerging evidence highlights the time-dependent mechanical properties of the ECM (encompassing viscoelasticity, plasticity, and non-linear elasticity) as an equally important regulator of cellular behavior, particularly in 3D environments where physical confinement imposes additional mechanical constraints.¹¹⁶

For example, in 2D culture, intermediate viscoelasticity enhances cell spreading on soft substrates, whereas, on stiff substrates, cell spreading appears less sensitive to changes in viscosity.¹¹⁷ However, the effects of viscoelasticity are even more pronounced in 3D environments, and studies in 3D culture conditions reveal that matrix viscoelasticity can govern cellular behaviors that are not observed in purely elastic matrices.^118,119 A notable example is the proliferation of cancer spheroids, where growth is primarily driven by the viscoelastic properties of the matrix rather than its stiffness.¹²⁰ Moreover, matrix remodeling is essential for cell expansion and viability in 3D cultures. In purely elastic, non-degradable matrices, the buildup of internal elastic stresses creates a mechanically restrictive environment, limiting cell expansion and thus hindering cell viability and proliferation. In contrast, viscoelastic matrices that permit stress relaxation or degradation enable cells to relieve mechanical confinement, create space for division and migration, and support robust organoid development.^121,122 These findings underscore that viscoelasticity is not simply an accessory property of the matrix but a fundamental regulator of 3D cell behavior, particularly within 3D confined microenvironments relevant to organoid culture.

Alginate-based hydrogels have become a widely used model system for studying viscoelasticity due to their tunable stress-relaxation properties. Reversible ionic cross-linking via divalent cations, such as calcium ions, forms ionic bridges between adjacent polymer chains within alginate hydrogels. By varying alginate molecular weight, calcium ion concentration, and incorporating PEG spacers, researchers can modulate relaxation times independently of stiffness and ligand density.^119,123,124 Importantly, because mammalian cells do not secrete enzymes to degrade alginate, any observed effects on cellular behavior can be directly attributed to the viscoelastic properties of the matrix rather than confounding factors related to matrix degradation.¹²⁴ Since alginate lacks intrinsic cell-binding motifs, adhesion peptides such as RGD are typically conjugated to its backbone. This strategy ensures a consistent density of adhesive ligands across different formulations, allowing researchers to isolate and assess the specific effects of matrix viscoelasticity on cell behavior without confounding influences from variations in cell-matrix adhesion.

Generally, fast stress-relaxing hydrogels use low-molecular-weight alginate (35–45 kDa), whereas high-molecular-weight alginate (∼280 kDa) yields slower stress relaxation.¹²³ Typically, calcium salts (e.g., calcium sulfate) are used because they are straightforward and produce rapid gelation. Studies using these hydrogels demonstrated that faster stress relaxation promotes cell volume expansion and cell cycle progression, mediated by the TRPV4 stretch-activated ion channel and the PI3K/Akt-p27^Kip1 signaling pathway.¹²⁰ In 3D cultures, fast-relaxing hydrogels (with a half-time (τ_1/2) of ∼20 s) enhanced cell proliferation and growth at a matrix stiffness of 20 kPa, while slow-relaxing hydrogels (τ_1/2 ∼ 600 s) resulted in cell cycle arrest. Mechanistically, cells typically increase their volume in preparation for division by exerting outward mechanical forces on the surrounding matrix, which in turn activates stretch-activated ion channels, such as TRPV4, through membrane deformation.^125–127 Consequently, rapid stress relaxation reduces mechanical confinement, thereby enabling more effective cell growth and division. A similar mechanism has been observed in MSCs, where faster stress relaxation enhances osteogenic differentiation by promoting TRPV4 activation and subsequent nuclear localization of RUNX2.⁴³

The influence of viscoelasticity extends beyond individual cells to tissue-scale morphogenetic processes.¹²⁸ In intestinal organoid cultures, fast-relaxing hydrogels (τ_1/2 ∼ 30 s) at a constant stiffness (5 kPa) facilitated rapid lumen expansion and symmetry-breaking events, such as protrusion formation and budding, whereas organoids in slow-relaxing hydrogels (τ_1/2 ∼350 s) exhibited slower expansion and remained morphologically unchanged.¹²¹ This behavior was associated with enhanced phosphorylation of focal adhesion kinase (FAK), nuclear YAP localization, and Arp2/3 complex-mediated cytoskeletal remodeling, which promoted spatially coordinated cell division and morphogenesis.

Beyond alginate-based hydrogels, advanced synthetic viscoelastic materials offer unprecedented control over matrix mechanics and have been increasingly applied to organoid culture. For example, photo-responsive hydrogels developed by Anseth and colleagues enable spatiotemporal modulation of local viscoelasticity through photo-induced softening (Fig. 5).¹²⁹ These systems incorporate allyl sulfide crosslinks, which undergo reversible bond exchange upon light activation, leading to a transient increase in stress relaxation without altering the bulk stiffness of the matrix. This photopatterning approach allows for localized control of matrix viscoelasticity, which in turn regulates epithelial curvature and directs crypt formation by modulating cell shape-induced mechanotransduction. Additionally, these photo-induced morphological changes establish spatial gradients of mechanical tension and membrane depolarization, ultimately promoting symmetry breaking and crypt formation through YAP-mediated mechanotransduction signaling pathways.

FIG. 5.

Photo-responsive hydrogels enable spatiotemporal modulation of local viscoelasticity through photo-induced softening. (a) Schematic illustrating organoid encapsulation in photo-responsive hydrogels, with or without photopatterning (unpatterned; no light activation). Spatially defined softening regions (yellow), created by light activation, direct crypt formation and shape the epithelial architecture. (b) Orientation order parameter based on crypt formation angles, indicating that crypt development is directed and templated by the patterned epithelial shape. (c) Immunostaining of organoids showing lysozyme-positive Paneth cells confined to crypt ends. (d) Quantification of lysozyme-positive crypt numbers demonstrates significantly more crypts in patterned regions compared to unpatterned controls, indicating that local viscoelastic modulation primes these areas for crypt development through cell shape-induced mechanotransduction [Adapted with permission from Yavitt et al., Sci. Adv. 9(3), eadd5668 (2023). Copyright 2016 AAAS¹²⁹].

Notably, Krieg and colleagues developed DyNAtrix, an innovative DNA-based viscoelastic hydrogel platform that offers highly programmable and precise control over stress relaxation for organoid culture.¹³⁰ DyNAtrix is built on a poly(acrylamide-co-acrylic acid) backbone functionalized with anchor DNA strands, which serve as universal attachment sites for DNA-based crosslinkers. By using a combinational crosslinker library (CCL) with diversified overlap domains—engineered through sequence variations and ambiguous nucleotides—the cross-linking kinetics can be finely tuned. This allows the stress relaxation time of the hydrogel to be adjusted from less than one second to several hours, independently of stiffness or other matrix properties. Additionally, heat-activated crosslinkers enable rapid and controlled gelation at physiological temperature (37 °C), creating a homogenous and cytocompatible environment for 3D organoid encapsulation. When functionalized with RGD peptides to promote cell adhesion, DyNAtrix supports the proliferation and long-term maintenance of trophoblast organoids, with organoid number, size, morphology, and marker expression comparable to those cultured in Matrigel. Furthermore, the system allows dynamic matrix remodeling through enzymatic degradation with recombinant DNase I, enabling controlled organoid release and passaging for up to 21 days.

Together, these studies form a cohesive narrative: cells sense not only the magnitude of stiffness but also the timescale over which resistance dissipates. Viscoelastic matrices permit cells to remodel and relieve mechanical stresses, unlocking proliferation, morphogenesis, and lineage-specific programs that purely elastic materials cannot support. Importantly, the relevance of viscoelasticity in 3D systems has only recently come to light and involves far more complex parameters than traditional 2D stiffness measures. Although we now recognize that stress relaxation kinetics can override static mechanical cues, the precise molecular pathways and mechanosensors translating viscoelastic signals into cellular decisions remain largely unexplored, especially within organoid research.

Recent advancements in viscoelastic material development have opened exciting new possibilities for organoid culture. Beyond viscoelasticity, although reports remain limited, emerging studies suggest that other forms of viscous dissipation, such as plasticity,^131,132 can also influence organoid morphogenesis, further underscoring the broader impact of time-dependent mechanical properties on cellular behavior. This growing body of research suggests that the ability to precisely engineer programmable, time-dependent mechanics that align with the dynamic nature of organoid development could unlock unprecedented opportunities for advancing 3D organoid culture systems.

D. Matrix geometry

The geometrical properties of the ECM—including porosity, fiber architecture, and topological patterning—play a fundamental role in regulating cellular behaviors and, consequently, organoid development and function. These structural features define the physical landscape in which cells reside, directly impacting cell-matrix interactions, mechanotransduction, and the spatial organization of multicellular structures.

One of the key geometrical parameters in 3D matrices is pore size, which is not only associated with the bulk mechanical properties of matrices at the macroscopic level but also profoundly influences cell behavior at the microscale.^133–136 In general, when pores are smaller than a single cell size, they increase surface area and facilitate initial cell attachment but often lead to overcrowding and restricted proliferation. Conversely, larger pores (≥300 μm) minimize physical confinement but may cause heterogeneous cell distribution and uneven cell clustering.¹³⁵ Critically, in the context of organoid formation, a pore size that balances confinement with sufficient space for cell-cell interactions enables robust cell clustering and tissue-like organization.^137,138 For instance, scaffolds with 140 μm pores supported the clustering of iPSC-derived liver progenitors into liver-bud-like structures with enhanced hepatic gene expression, whereas smaller pores (40–100 μm) hindered organized organoid formation. When pores are large enough to allow individual cells to enter and adhere directly to internal surfaces, cells colonize the scaffold uniformly; in contrast, overly small pores confine cells at the surface, causing them to aggregate in clumps rather than form cohesive organoids. Thus, whether initiating organoids from single cells or from multicellular aggregates, careful selection of matrix pore size is essential to ensure reproducible yields, uniform morphology, and proper tissue function. Furthermore, appropriate porosity enhances the diffusion of gases, nutrients, and metabolic waste, which reduces reliance on vascularization in long-term 3D cultures—an important consideration for sustaining organoid growth over extended periods.^52,136

Organoid cultures typically rely on spontaneous stem cell self-organization, a process that often results in significant variability in organoid size, shape, and cellular composition. This inherent stochasticity not only compromises the reproducibility of organoid formation but also limits its scalability and translational potential. To overcome these challenges, researchers have increasingly turned to precisely engineered geometrical cues as a strategy to provide external spatial guidance. By imposing defined physical constraints and topological patterns, these bioengineering approaches aim to direct organoid morphogenesis in a more controlled, predictable, and reproducible manner, improving consistency in organoid architecture and function.

For example, Warmflash et al. used 2D micropatterned substrates—arrays of adhesive circles of defined diameter—to confine human embryonic stem cell (hESC) colonies within precise geometries.¹³⁹ Colonies grown on these patterns reliably self-organized into concentric gastruloid zones, with distinct trophectoderm, mesendoderm, and ectoderm regions, mirroring early embryonic development. This geometric confinement enhanced cell-cell communication by increasing cell density within defined regions, thereby promoting more consistent differentiation outcomes and offering improved spatial control compared to traditional embryoid body (EB) cultures. More recently, Zheng et al. translated this concept into a 3D microfluidic system.¹⁴⁰ Their device features three parallel microchannels separated by microposts: the central channel is cast with Geltrex, which contracts during gelation to form concave pockets, while adjacent channels supply cells and soluble factors. Single hESCs seeded into the cell-loading channel settle into these gel pockets and cluster, and a BMP4 gradient induced in the induction channel drives dorsal-ventral patterning of epiblast-like cysts. By combining 3D geometric confinement with controlled chemical gradients, this approach recreates dynamic in vivo-like patterning in a scalable, highly controllable platform.

Building on this concept of geometric guidance, free-floating fiber-based scaffolds have been developed to further modulate organoid morphology and enhance differentiation outcomes. For example, poly(lactide-co-glycolide) (PLGA) fiber microfilaments were incorporated into EB cultures to induce elongation of the aggregates, resulting in an increased surface-area-to-volume ratio while preserving dense cell-cell interactions.¹⁴¹ This engineered geometry promoted more efficient neuroectoderm specification and improved cortical organoid organization compared to conventional spherical EBs.

Similarly, carbon fibers have been used as free-floating scaffolds to support midbrain organoid formation.¹⁴² Like PLGA microfilaments, carbon fibers provided physical guidance that modulated EB morphology and promoted organized tissue formation. However, they offered a distinct advantage due to their multi-scale porosity, featuring micro- and nanoscale pores. This hierarchical porosity enhanced serum protein adsorption and cell-scaffold interactions, leading to improved differentiation efficiency of hiPSCs toward midbrain organoids compared to PLGA microfilaments. In addition to geometry, the material composition and surface properties of the fiber scaffold played a critical role. While PLGA, carbon fibers, and collagen-rich fibers supported uniform hiPSC adhesion and EB elongation, cellulose fibers led to the formation of spherical aggregates instead.^141,142 These findings highlight the importance of considering both scaffold architecture and material choice in directing organoid morphogenesis and lineage commitment.

Expanding on the role of scaffold geometry, recent studies have employed advanced microfabrication techniques to precisely control structural parameters such as angle, spacing, and shape in fibrous scaffolds for organoid patterning. One notable example is the use of melt electro-writing (MEW), a high-resolution extrusion-based 3D printing technique, to fabricate centimeter-scale grid scaffolds with defined geometries, including square, rhombus, and triangular patterns [Figs. 6(a)–6(c)].¹⁴³ When hESCs were cultured on these grid-like scaffolds, the geometry was found to direct collective tissue behavior in a curvature-dependent manner. For instance, lumen formation and cell density were influenced by scaffold angles, with 45° grids producing larger lumens and a 26% increase in cell density compared to 90° grids [Figs. 6(d) and 6(e)]. These findings demonstrate that scaffold geometry can quantitatively modulate tissue morphology by altering spatial and mechanical constraints during early organoid formation. Furthermore, grid scaffolds with optimized spacing of 500 μm promoted the formation of interconnected cerebral organoids, resulting in thicker tissue development and enhanced neurogenesis compared to scaffolds with wider spacing of 1000 μm [Figs. 6(f) and 6(g)]. Notably, by adjusting grid geometry, researchers were able to spatially organize discrete cerebral organoids in a high-throughput manner, demonstrating the potential of these engineered platforms to improve the scalability and reproducibility of organoid culture systems. While this study focuses primarily on geometrical parameters, the intrinsic stiffness of MEW fibers—often in the MPa to GPa range—far exceeds physiological levels. Future studies should therefore decouple geometry from mechanics by independently varying fiber stiffness to elucidate how scaffold rigidity, alongside geometric cues, shapes organoid formation and function.

FIG. 6.

Scaffold-guided embryonic body and organoid platform. (a) Microfibrous scaffold formation via melt electrospinning writing (MEW). (b) Schematic of MEW-fabricated scaffolds used to guide lumenogenesis and cerebral organoid growth. The cell-material interface of the scaffold is engineered to enable both interconnected and spatially discrete organoids. (c) Graphical illustrations of different scaffold geometries. (d) Representative fluorescence images showing lumen formation (green) on square vs triangular grid scaffolds. (e) Quantitative analysis indicates that the higher curvature provided by 45° angles increases cell density and yields larger, less circular lumens. (f) Representative bright-field images of cerebral organoids developing on scaffolds with 500 or 1000 μm spacing in square grids, as well as on triangular grids. (g) Immunostaining for the dorsal forebrain marker PAX6 and the forebrain marker FOXG1 [Adapted with permission from Ritzau-Reid et al., Adv. Mater. 35(41), 2300305 (2023). Copyright 2023 John Wiley & Sons, Ltd.¹⁴³]. These findings collectively suggest that scaffold geometry can spatially organize organoid formation by influencing lumen formation and cell distribution.

Recent advances in bioengineering have focused on replicating in vivo-like tissue geometries through the development of topologically mimetic ex vivo platforms.^144,145 These strategies aim to recapitulate the stereotypical spatial organization and morphological patterning observed in native tissues by leveraging precise geometric engineering techniques. By establishing reproducible local differences in cell packing, density, and morphology, these platforms provide a deterministic framework for guiding organoid development and morphogenesis, addressing the inherent stochasticity of conventional organoid models. For example, Gjorevski et al. developed bioengineered hydrogel substrates that mimic the crypt-villus architecture of the native intestine.¹⁴⁵ These 3D micropatterned substrates guided the self-organization of ISCs and their progeny into organoids with well-defined tissue polarity and regionalization. Spatial differences in cell packing between the crypt-like and villus-like regions localized ISCs to the crypt base while promoting the differentiation of progenitor cells along the villus structures, thereby establishing a functional crypt-villus axis analogous to the native intestinal epithelium.

Recently, Lutolf's group¹⁴⁶ developed a microfluidic-based hydrogel platform, called Transgels, by using silicone elastomer stamps to micropattern hydrogel scaffold surfaces for a stem-cell-derived epithelial monolayer with bilateral accessibility. Cells derived from 3D organoids were seeded onto these patterned scaffolds and cultured for several days to promote spatial organization. Within just one day, cells showed preferentially. Immunostaining for Sox9 identified stem cells concentrated within crypt-like regions of the scaffold, closely mimicking the native intestinal stem cell niche. RNA sequencing results revealed that the Transgel organoid models exhibit a strong correlation with their in vivo tissue counterparts and closely resemble 3D organoids. This finding confirms that engineered organoids based on Transgels effectively preserve the physiological relevance of traditional 3D organoid models.

In another study, Lorenzo-Martín et al. successfully integrated microfabrication and tissue-engineering technologies to create patient-specific colorectal cancer (CRC) organoid models that closely mimic the architecture of native colon tissue.¹⁴⁴ Using a microfluidic organ-on-a-chip platform combined with scaffold-guided organoid morphogenesis, they generated miniaturized colon structures featuring crypt-like invaginations and lumen-like domains. When seeded with patient-derived CRC cells, these engineered constructs recapitulated tumor growth dynamics with greater precision than conventional organoid systems.

Mechanistically, these studies revealed that variations in tissue geometry and cell packing density generate heterogeneities in YAP signaling activity, which regulate stem cell maintenance and differentiation. High YAP activity in villus-like structures promoted proliferation and suppressed stem cell maintenance, while low YAP activity in crypt-like regions maintained ISC identity and supported Notch-mediated Paneth cell differentiation. Notably, these spatial patterning effects were achieved without exogenous biochemical gradients, highlighting the instructive role of geometry and mechanical cues in directing cell fate decisions.

Collectively, these findings underscore the potential of engineered tissue geometry as a deterministic cue for symmetry breaking and regionalization in organoid systems. By mimicking native topological features, such platforms enable greater control over organoid morphogenesis, facilitating the recapitulation of physiologically relevant tissue architectures and functional compartmentalization. These insights suggest that spatial variations in cell morphology and packing, driven purely by geometric design, can orchestrate complex morphogenetic events that are typically stochastic in conventional organoid cultures. Moving forward, it will be critical to dissect how geometric cues interact with other mechanical parameters, such as matrix stiffness, viscoelasticity, and degradability, to jointly regulate cell behavior and organoid formation.

III. REMAINING CHALLENGES AND FUTURE PERSPECTIVE

A comprehensive understanding of mechanobiological principles underlying organ development in vivo provides a valuable framework for engineering next-generation organoid culture platforms. Translating these insights into in vitro systems—by mimicking the spatiotemporal dynamics of matrix adhesion, stiffness, viscoelasticity, and geometry—offers promising strategies to better direct stem cell fate, enhance tissue organization, and promote the functional maturation of organoids. Such approaches address key limitations of conventional organoid systems, improving their reproducibility, scalability, and physiological relevance.

Despite recent advances in mechano-modulatory biomaterials, several challenges remain. Synthetic matrices such as PEG-based hydrogels offer tunable biomechanics and improved reproducibility compared to Matrigel. However, they lack the complex tissue-specific biochemical cues present in native ECM. While cell-adhesive peptides and bioactive molecules can be incorporated to partially compensate, replicating the full biochemical landscape of native tissues remains difficult. Furthermore, biomechanical and biochemical signals are deeply intertwined in vivo—mechanical cues modulate biochemical signaling, and vice versa^147–149—highlighting the need for integrated chemo-mechanoregulation strategies in organoid engineering. For example, the addition of bone-specific hydroxyapatite in bone organoid cultures has been shown to enhance mineralization and functional maturation,¹⁵⁰ underscoring the importance of combined biochemical and biomechanical regulation.

While this review has primarily focused on matrix-driven intrinsic biomechanical cues, emerging evidence suggests that external mechanical stimulation can also play a pivotal role in guiding organoid development. Extrinsic mechanical forces such as gravity,^151,152 compression,^153–155 stretch,^156,157 shear stress,^158,159 and electromagnetic fields^160–163 have been shown to influence stem cell fate and subsequently organoid morphogenesis.^164,165 Microfluidic systems, in particular, offer precise control over mechanical stimulation, enabling spatiotemporal modulation of mechanical inputs, automated nutrient delivery, and real-time monitoring of organoid development.¹⁶⁶ Future studies should focus on integrating mechano-modulatory materials with microfluidic platforms to deliver dynamic, programmable mechanical stimuli for enhanced organoid engineering.

In parallel, advanced fabrication techniques, including 3D bioprinting, present opportunities to overcome the stochastic nature of organoid self-organization. By enabling the spatially controlled placement of cells and the simultaneous modulation of scaffold mechanical properties, 3D printing facilitates the generation of complex organoid structures beyond the traditional spherical morphology. This approach allows for the fabrication of larger organoids spanning millimeter- to centimeter-scale, with improved control over tissue architecture and multi-lineage cell patterning. Additionally, increasing the surface area and porosity of printed constructs can enhance nutrient diffusion and promote long-term organoid viability. Moving forward, the development of versatile, mechanistically modulatory bioinks will be critical for advancing 3D printing-assisted organoid culture.

Finally, artificial intelligence (AI)-assisted organoid engineering is poised to transform the field by overcoming the limitations of conventional scaffold and matrix design, which largely depend on empirical, trial-and-error approaches.¹⁶⁷ These traditional methods often lack precision, scalability, and consistency, hindering the ability to systematically optimize biomaterial properties for organoid culture.¹⁶⁸ In contrast, AI and machine learning offer powerful data-driven strategies capable of analyzing large, complex datasets to identify optimal biomaterial designs and predict organoid responses to specific biomechanical environments.¹⁶⁹ For example, AI-driven platforms have successfully reduced heterogeneity in organoid cultures and uncovered previously unrecognized relationships between matrix geometry and organoid function without human bias insights that are often missed using conventional analytical methods.¹⁷⁰ By processing high-dimensional datasets derived from organoid libraries cultured under defined biomechanical conditions, AI frameworks can rapidly evaluate functional outcomes and quantify how specific engineering parameters influence organoid development and physiology. This predictive capability enables the rational design of scaffolds and biomaterials tailored to support desired organoid structures and functions.

Looking ahead, the integration of advanced technologies into organoid engineering, particularly mechanobiological strategies, will be pivotal in developing dynamic, in vivo-like matrices that more accurately replicate the native tissue environment. By precisely controlling intrinsic and extrinsic biomechanical cues, these approaches offer the potential to enhance the reproducibility of organoid cultures, minimize inter-organoid heterogeneity, and promote more consistent tissue maturation. Ultimately, such innovations will accelerate the translation of organoid technologies into personalized medicine, disease modeling, and regenerative therapies.

AUTHOR DECLARATIONS

Conflict of Interest

The authors have no conflicts to disclose.

Author Contributions

Mohsen Taghizadeh: Conceptualization (lead); Writing – original draft (lead). Ali Taghizadeh: Writing – original draft (equal). Hye Sung Kim: Conceptualization (lead); Funding acquisition (lead); Supervision (lead); Writing – review & editing (lead).

DATA AVAILABILITY

Data sharing is not applicable to this article as no new data were created or analyzed in this study.