Droplet microfluidic devices for organized stem cell differentiation into germ cells: capabilities and challenges

Abstract

Demystifying the mechanisms that underlie germline development and gamete production is critical for expanding advanced therapies for infertile couples who cannot benefit from current infertility treatments. However, the low number of germ cells, particularly in the early stages of development, represents a serious challenge in obtaining sufficient materials required for research purposes. In this regard, pluripotent stem cells (PSCs) have provided an opportunity for producing an unlimited source of germ cells in vitro. Achieving this ambition is highly dependent on accurate stem cell niche reconstitution which is achievable through applying advanced cell engineering approaches. Recently, hydrogel microparticles (HMPs), as either microcarriers or microcapsules, have shown promising potential in providing an excellent 3-dimensional (3D) biomimetic microenvironment alongside the systematic bioactive agent delivery. In this review, recent studies of utilizing various HMP-based cell engineering strategies for appropriate niche reconstitution and efficient in vitro differentiation are highlighted with a special focus on the capabilities of droplet-based microfluidic (DBM) technology. We believe that a deep understanding of the current limitations and potentials of the DBM systems in integration with stem cell biology provides a bright future for germ cell research.

Supplementary Information

The online version contains supplementary material available at 10.1007/s12551-021-00907-5.

Introduction

In vitro stem cell differentiation into the germ cells and healthy gametes is a promising approach for infertile couples who cannot benefit from current infertility treatments such as assisted reproductive techniques (ARTs). These patients have usually lost their ability to produce functional gametes due to disruptions in germline development (Galdon et al. 2016). Therefore, understanding the governing mechanisms in germline development is crucial for the establishment of advanced infertility treatments. However, a low number of germ cells, especially in the early stages of development, has been a serious challenge to obtain sufficient material for research (De Felici 2013; Irie and Surani 2017). Accordingly, the development of specific systems that enable scientists to unveil the molecular mechanism of germ cell development is highly demanded. In this regard, pluripotent stem cells (PSCs) have opened new windows in producing an unlimited source of functional germ cells in the laboratory setting (Sabour and Schöler 2012; Esfandiari et al. 2015). Nevertheless, low efficiency, heterogeneity, and non-reproducibility of PSC differentiation into the germ cells are still the main challenges of conventional methods. This is because of their inability to precisely reconstitute the key features required for efficient in vitro germ cell induction, including systematic, controlled, and localized delivery of bioactive agents (BAs), and true recapitulation of the physicochemical properties of the native niche.

Integrating tissue engineering techniques and advanced biomaterials with stem cell biology have provided excellent opportunities to precisely reconstruct stem cell artificial niches. Today, hydrogel-based biomaterials are known as appropriate candidates for reflecting the physicochemical properties of in vivo microenvironment due to their specific characteristics, i.e., high water content and similarity to the extracellular matrix (ECM). As well known, 3D scaffold-based culture methods, compared to 2D and 3D scaffold-free culture techniques, have greater ability to imitate what the cells experience in the native tissues, including cell-cell and cell-matrix interactions in all three spatial dimensions and BAs receiving in a right dose and spatiotemporally controlled manner. Lately, advanced HMPs with tunable morphology, structure, and characteristics were introduced as a powerful platform for well-organized stem cell differentiation through systematic BA delivery and precise imitation of stem cell niches (Esfandiari et al. 2017; Fang et al. 2018). Compared to the bulk hydrogels with encapsulated cells, utilizing cell-laden HMPs significantly improves the diffusion of nutrient and soluble factors toward the core region. According to the literature, HMPs can operate as either an inside-out soluble factor delivery agent within the cell aggregates, or 3D BA-laden matrices for embedding/encapsulating specified cell numbers to form aggregates and induce differentiation (Siltanen et al. 2016; Esfandiari et al. 2017).

The cellular niche contributes to stem cell fate decisions through various physicochemical features of the HMPs (Fig. 1). The biophysical features of HMPs including stiffness and elasticity, degradability, porosity, and hydrogel mesh size can be tuned to a favorable state by regulating the cross-linking level, the composition, and polymer concentration (Hogrebe and Gooch 2016; Xin et al. 2018; Truong et al. 2019; Daly et al. 2020). These factors can alter cellular mechanosensing and affect the proliferation and differentiation of PSCs into a specific lineage (Zhang et al. 2017; Lee and Cha 2018). Essential biochemical cues can also be provided for the stem cells through ECM proteins as well as the systematic and controllable delivery of critical BAs including growth factors (GFs), small molecules, morphogenetic factors, and cytokines (Kawaguchi et al. 2005; Zhang et al. 2011). It is well known that these biochemical and biophysical cues are coupled together so that, for instance, the forces generated by cell-cell and ECM-cell interaction result in a mechanical framework which is then transduced to the biochemical signaling within the cells (Vogel and Sheetz 2006; Wozniak and Chen 2009).

Fig. 1.

Schematic of physicochemical features affecting the stem cell differentiation pathways

There are many different strategies to produce HMPs such as DBM systems, batch emulsions, mechanical fragmentation, electrohydrodynamic spraying, and dispersion/precipitation polymerization (Young et al. 2012; Hinton et al. 2015; Ryu et al. 2016; Qayyum et al. 2017; Zhang et al. 2018). Among all, DBM systems have shown greater ability to generate advanced HMPs with desired characteristics, such as high-level monodispersity, size and shape, tailored structures, and integrated functions for systematic BA release, as a result of higher controllability during the droplet-formation process (Zolnik and Burgess 2007; Bock et al. 2011; Li et al. 2018; Campos et al. 2013; Zhang et al. 2017). These abilities enable DBM devices to reconstruct stem cell niche models in microscale and highly accurate fashion.

In this review, first, the key requirements for an efficient in vitro germ cell derivation from PSCs are described. Then we will discuss the abilities and strengths of HMPs in stem cell differentiation into different cell types including germ cells. Thereafter, the impacts of physicochemical properties of hydrogels on the differentiation pathways into a specific lineage are investigated. Afterward, different HMP-based cell engineering strategies and their abilities in stem cell differentiation are reviewed. Next, strengths and restrictions of traditional methods for HMP production are highlighted followed by discussing the superiority of DBM devices in producing advanced HMPs. Also, the obstacles and solutions of developing a sophisticated microfluidic platform for the high-throughput on-chip aggregate formation and long-term aggregate culturing are discussed. Finally, relying on the results of the reviewed papers, we suggest a conceptual design of an integrated device appropriate for high-throughput in vitro germ cell induction and gametogenesis.

Requirements for in vitro germ cell induction improvement

Biophysical cues: a true recapitulation of the native niche

The biophysical properties of the cellular niche can either promote or inhibit certain differentiation pathways through various mechanisms like mechanical traction forces and cell-ECM interactions (Guilak et al. 2009; Tsou et al. 2016; Zhang et al. 2017; Labriola et al. 2018; Lee and Cha 2018). These stem cell-niche interactions in the human body, i.e., cell-cell and cell-matrix interactions in all three spatial dimensions, target intracellular signaling pathways and regulate gene expression patterns (Guilak et al. 2009; Dickinson et al. 2011; Kshitiz et al. 2012). Therefore, accurate imitation of such conditions is in tremendous demand in developmental biology and stem cell research. There are various biophysical factors, including stiffness, degradability, mesh size, and porosity, that can affect the efficiency of in vitro germ cell induction. Therefore, to efficiently differentiate PSCs into the germ cells, all these factors should be considered and optimized during the designing HMPs.

Biochemical cues: systematic and homogeneous delivery of BAs

Efficient delivery of critical BAs in aggregate-based systems, such as Bone Morphogenetic Protein 4 (BMP4) for germline induction, in a controllable fashion, is a very important factor in efficient in vitro differentiation. Efficient delivery provides the opportunity for the cells located in the middle of aggregates to receive induction factors as well as cells located peripherally. More cells receive induction factor, more cells are differentiated into germ cells. Accordingly, as evidenced by Esfandiari et al., more controllability and homogeneity of BMP4 delivery can significantly improve the efficiency of in vitro germ cell derivation compared to the cases with non-homogeneous BMP4 distribution (Hayashi et al. 2011; Esfandiari et al. 2017). In addition to BMP4, there are more essential BAs required for efficient differentiation that should be presented to the cells in individual time and dose. Conventional systems like 2D and 3D scaffold-free culture systems are not appropriate candidates to meet mentioned requirements, while scaffold-based culture systems, in particular HMPs, as will explain in the following sections, have shown great ability in perfect systematic and homogeneous delivery of multiple critical BAs around the stem cells.

As abovementioned, one of the most critical BAs required for germline induction from SCs is BMP4 (Lawson et al. 1999, p. 4). To be more precise, during embryonic development, secreted BMP4 from the extra-embryonic ectoderm induces the formation of primordial germ cells (PGCs), precursors of the germline, from the most posterior part of the epiblast (Hayashi et al. 2007). BMP4 is an extracellular signal that binds to BMPR-I and BMPR-II on the cell surface, which has a key role in inducing germ cell fate (Aramaki et al. 2013). Knockout studies showed a lack of germ cells in BMP4 homozygous mutants and very few germ cell numbers in heterozygote mutants (Lawson et al. 1999). BMP4 guides the mesodermal factor T (Brachyury) to induce the expression of germline-specific genes such as Blimp1 and Prdm14 (Aramaki et al. 2013). Blimp1 is the first gene expressed in response to BMP4 and it drives two main functions: (i) inhibition of somatic gene expression and (ii) induction of germline-specific genes, including Prdm14 and Stella (Aramaki et al. 2013). It also temporarily inhibits cell proliferation by suppressing the expressions of the somatic genes Hoxa1, Hoxb1, Fgf8, and Snail. Both Blimp1 and Prdm14 promote epigenetic reprogramming in PGCs (Tang et al. 2016) (Fig. 2).

Fig. 2.

In vivo and in vitro primordial germ cell (PGC) formation. Schematic diagram of mouse primordial germ cell (PGC) development and the related signaling pathways. BMP4 plays a key role during germline induction and embryo development. Blimp1 and Prdm14 expressions lead to inhibition of somatic gene expression and induction of germline-specific genes that include and promote epigenetic reprogramming in PGCs

HMPs, a powerful platform for directing stem cell differentiation

It is well known that the most efficient method for in vitro germ cell induction that can lead to a live birth is based on the 3D cultures and aggregate formation (Hayashi et al. 2011, 2012). 3D culture techniques are appropriate platforms for recapitulation of cells’ experience in the native tissues (Friedrich et al. 2009; Kimlin et al. 2013). The current 3D culture systems can generally be categorized into Scaffold-free and Scaffold-based systems. Each of these techniques has exclusive properties and can provide different niches for in vitro germ cell induction. In scaffold-free strategy, the cells are injected into specific containers to establish cell-cell interaction and form compacted aggregates. The most common and popular methods among the scaffold-free systems for 3D cellular aggregate formation are hanging drop, spinner flasks, and ultra-low attachment plates (Friedrich et al. 2009; Moshksayan et al. 2018). However, these approaches suffer from difficult medium exchange and manipulation, being time-consuming, producing non-uniform aggregates, poor efficiency, and generating high shear stresses (Hwang et al. 2013; Chatzinikolaidou 2016; Egger et al. 2018). Further and more important shortcomings associated with these systems are their inability in controlling and modifying biophysical and biochemical characteristics of extracellular matrix (ECM), as well as non-homogenous neither controllable BA delivery throughout the cell aggregates. In these conventional methods, the outside-in transport of the soluble factors is restricted by the diffusion barriers such as tight cell-cell junctions and secreted ECM which leads to non-uniform and gradient delivery of BAs, and thus, heterogeneity of differentiation induction in the aggregates (Bratt-Leal et al. 2013; Esfandiari et al. 2017). This restriction was reflected in Hayashi’s research, as one of the most successful platforms for in vitro germ cell production, where induced germ cells were localized peripherally in cell aggregations caused by the non-uniform distribution of BMP4 throughout the cell aggregates (Hayashi et al. 2011). Alternatively, scaffold-based systems, in particular HMPs, have shown great ability to address the abovementioned drawbacks. HMPs with tunable physicochemical characteristics act as biocompatible, miniaturized, and functional niches for well-organized stem cell differentiation through precise reconstruction of stem cell niche (Tayalia and Mooney 2009). Since the porous structure, excellent biocompatibility, inherent capacity to absorb water, and similar viscoelastic properties to the tissues, the key features of the 3D microenvironment can be precisely imitated by HMPs. Additionally, in comparison with conventional soluble delivery, the HMP-based approach benefits from a significant decrement of required BAs to obtain similar results leading to cost reduction (Bratt-Leal et al. 2013; Heidariyan et al. 2018).

Physicochemical properties of HMPs affecting stem cell differentiation

The cellular niche contributes to stem cell fate decisions through physicochemical features of the HMPs including stiffness, degradability, porosity, and hydrogel mesh size as the biophysical features, and accessibility of various BAs and proteins for the stem cells in a systematic and controlled fashion as biochemical cues. Almost all the abovementioned physicochemical properties are more or less controllable via three parameters of cross-linking level, polymer concentration and molecular weight, and the material itself. In the following sections, we investigate these physicochemical properties, their related controlling factors, and how they can influence stem cell differentiation toward a specific lineage.

Biomaterial type

The biomaterial type used in HMP fabrication is important because its properties can directly affect the differentiation signaling pathway (Murphy et al. 2014; Donnelly et al. 2018). HMPs can be generally categorized as natural or synthetic biomaterials (Lee et al. 2013; Wang et al. 2017b; Pina et al. 2019; Vijayan et al. 2019). Natural hydrogels intrinsically provide proper cell attachment and biodegradability without requiring biochemical modifications (Allazetta and Lutolf 2015). They are highly biocompatible and provide natural binding sites for the improvement of cell-matrix interactions (Floren et al. 2016; Yang et al. 2017a). However, several drawbacks are associated with these natural polymers such as low controllability of their gelation and degradation, weak stability, and poor mechanical properties (Huang et al. 2015a; Utech et al. 2015). Alginate, collagen, fibrin, hyaluronic acid, and chitosan are among the most common natural polymers utilized in both HMP-incorporated aggregates and cell-encapsulated microgels (Hoque et al. 2015; Caballero Aguilar et al. 2019). In contrast to the natural hydrogels, physicochemical properties of synthetic hydrogels like porosity, stability, stiffness, degradation, and release rate can be tightly and reproducibly tuned by specifying the polymer molecular weight and concentration, cross-linking level, and functional sites at the molecular scale (Rossow et al. 2012; Higuchi et al. 2015; Tsou et al. 2016). To resolve their poor interaction with cells, synthetic hydrogels are potential to be decorated with various BAs (DeForest and Anseth 2012; Seliktar 2012; Allazetta and Lutolf 2015). However, incorporation of BAs into the synthetic polymers exposes them to harsh fabrication processing such as high temperature and organic solvents which can lead to BA denaturation (Chu et al. 2011). Despite the low loading capacity through surface adsorption, post-fabrication loading of BAs is currently used to circumvent this problem (Wang et al. 2017b; Buie et al. 2020). Alternatively, natural polymers benefit from a high capacity for in-line loading of BAs due to simplicity and mild fabrication processing. Therefore, regarding the importance of material properties in stem cell fate decisions, recognition of the most appropriate biomaterial for efficient germ cell and gamete induction is critical. This can be a natural hydrogel, a synthetic one, or even their combinations. Between the various biomaterials applied for BA delivery, those which provide controllable release rates are more important for germ cell induction. This is because the germ cell induction is a long-lasting process while it requires to be exposed to the critical GFs entire differentiation process.

Stiffness

Mechanotransductional mechanisms activated by mechanical forces can influence cell proliferation, growth, and differentiation (Sart et al. 2016; Narayanan et al. 2017). According to this, cellular differentiation can be directed by regulating the mechanical properties of cellular aggregates and providing distinct and stable mechanical cues (Labriola et al. 2018). HMP stiffness, as the most important mechanical feature, plays a decisive role in stem cell fate determination. As a proof of concept, Abbasi et al. explored the effect of stiffness on MSC differentiation independent of biochemical parameters and only by focusing on the mechanical characteristics of MPs. To do this, they mixed cell-sized and non-biodegradable polydimethylsiloxane (PDMS) MPs with different elastic modulus to produce cell-MP composite spheroids. They demonstrated that incorporation of such mechanoregulator MPs can be considered a simple and effective strategy to regulate MSC differentiation in 3D aggregations (Abbasi et al. 2018). In a similar study, Labriola et al. incorporated the polymer-based cell mimicking MPs inside the stem cell aggregates and demonstrated the effects of their stiffness on mechanophenotype and stem cell behavior (Labriola et al. 2018). They also showed that the stiffness of MP-contained spheroids is altered in accordance with MP elasticity. Furthermore, Zoldan et al. demonstrated that through regulating the scaffold stiffness, every three germ layers can be derived from hESC (Zoldan et al. 2011). The hydrogel stiffness can be determined depending on parameters like the material itself, cross-linking intensity, and the polymer concentration (Vining and Mooney 2017; Xin et al. 2018; Truong et al. 2019; Mohamed et al. 2020). The stiffness of hydrogel also influences the aggregate formation inside the HMPs, so that solid-like hydrogel with high stiffness impedes the aggregate formation and disturbs cell proliferation, and thus, the cell-cell interaction, as a critical factor for generating compact aggregates, is significantly decreased (Hwang et al. 2013; Moshksayan et al. 2018). As evidence, Yu et al. reported that utilizing a high concentration of alginate reduces cell proliferation and results in the formation of small bumpy shape clusters instead of aggregates due to the high rigidity and transport barrier for cell migration (Yu et al. 2015). However, producing a composition of indigestible hydrogels like alginate and agarose with digestible hydrogels such as collagen and Matrigel can resolve this problem (Hall and Brooks 2014; Wang and Wang 2014). Altogether, it can be claimed that optimizing the stiffness of HMPs is of significant importance for successful aggregate formation, and differentiation of stem cells into functional germ cells.

Biodegradability

The degradation rate of HMPs can differently affect the stem cell fate depending on hydrogel application in stem cell culture systems, i.e., cell encapsulation, or BA delivery. In the case of cell-embedded hydrogels, degradation frees up some space suitable for cell spreading and proliferation (Raeber et al. 2005). In degradable hydrogels, the stem cells can actively reorganize the surrounding matrix and adjust their morphology, as a decisive factor in the differentiation process, through secretion of matrix metalloproteinases (MMPs) and cell-exerted forces (Schwartz and Chen 2013; Schultz et al. 2015). Khetan et al. reported that the stem cells encapsulated in non-degradable hydrogels differentiated into the adipocytes with spherical shapes, while they showed spreading morphology and high traction with an altered fate to osteogenesis when embedded in degradable hydrogels, independent of matrix stiffness (Khetan et al. 2013). The degradability of the hydrogel matrix also enables the cellular condensation with a low cytoskeletal tension to form a compact 3D aggregate instead of multiple dispersed cell clusters (Toh et al. 2012). The degradation rate, in cases that HMPs are applied for BA delivery, can control the release profile and influence the differentiation through it. Unlike the natural hydrogels with uncertain and uncontrollable degradation, the degradation of synthetic hydrogels can be tuned by either utilizing photo-triggered degradation processes or incorporating the proteolytic or hydrolytic cross-links in hydrogels which makes them more appropriate for BA delivery (Rapp et al. 2018). The degradation rate of hydrogels can be controlled by tuning cross-link density so that higher (lower) cross-link density causes lower (higher) degradability (Hogrebe and Gooch 2016; Loebel et al. 2019).

Mesh size and porosity

The polymer mesh size affects both the cell viability and proliferation rate in cell-contained hydrogels so that larger mesh size improves the proliferation and viability as a result of better diffusion of BAs and paracrine signaling throughout the polymer network (Lin et al. 2011; Perera et al. 2018). In BA delivery application, similar to the degradability feature, the mesh size acts as a controller of release profile in such a way that in the same degradation rate, smaller mesh size reduces the BA release rate leading to a more sustained release profile. The cross-link density, polymer concentration, and molecular weight are the three decisive parameters in polymer mesh size specification. The degradability, stiffness, and mesh size are thoroughly interdependent, since all of them can be controlled by cross-linking level in such a way that higher (lower) cross-link density causes lower (higher) degradability, stiffer (softer) hydrogel matrix, and smaller (larger) mesh sizes. Finally, according to the literature, the porosity structure of hydrogels improves cell adhesion, growth, and spreading throughout the hydrogels (Huang et al. 2018). Also, improved differentiation of stem cells has been reported by Chen et al. as a result of culturing stem cells in a microporous hydrogel (Wei et al. 2018). The porosity can also impact the degradation rate as well as BA release rate depending on its pore sizes.

Controlled release of BAs

The importance of controlled and systematic release rate of BAs for a successful in vitro germ cell induction is elucidated in the “Biochemical cues: systematic and homogeneous delivery of BAs” section. Therefore, the parameters affecting the release rate are of utmost importance to be known and considered during the designing HMPs. Generally, the method of BA incorporation into the HMPs, i.e., physical entrapment or immobilization, specifies the BA release rate through different processes such as diffusion, swelling/diffusion, degradation, dissolution, and osmosis wetting phenomena (Ungaro et al. 2006; Tessmar and Göpferich 2007; Esfandiari et al. 2017). The physical entrapment of BAs is the most routine and simple approach which is gained through mixing BAs with the hydrogel prior to the solidification process. In this case, the transport rate of BAs through and out of the HMPs depends on degradation, mesh size, particle size, and loading percent. The affecting mechanism of these biophysical parameters is based on the degradation and diffusion (Drury and Mooney 2003; Kitamura et al. 2011). When the average mesh size is larger (smaller) than the BA hydrodynamic diameter, the release happens through diffusion (degradation). As mentioned in the previous sections, these factors can be controlled by tuning the cross-link density, polymer concentration, and molecular weight. Immobilization is another strategy for BA incorporation which provides better-controlled release compared to the simple physical entrapment. For instance, heparin is a well-known molecule with the ability to reversibly bind heparin-binding GFs to the hydrogel to slow their release through increasing BA-hydrogel affinity (Hettiaratchi et al. 2014, p. 2; Li and Mooney 2016). Immobilization of BAs to HMPs can occur through physical adsorption (ionic complexes or electrostatic interactions between BAs and hydrogels), chemical bonding, and secondary association (BAs association with hydrogels through an intermediate group in a reversible\irreversible fashion) (Tayalia and Mooney 2009). In the case of chemical bonding, unlike the two other methods, agent internalization by the cells is prohibited until the bond is broken. In cases that the chemical bonding is hydrolytically or enzymatically cleavable, the local cellular activities can control the release rate of loaded BAs (Bryant and Anseth 2003). Taken together, when designing the HMPs for in vitro germ cell induction, knowing their strengths\limitations aids to select the most appropriate method of BA incorporation to provide a desired release rate.

Strategies for HMP-based stem cell differentiation

According to the reports, HMPs can be applied in two different strategies for controllable and uniform releasing of critical soluble factors as well as resembling the physicochemical properties of the environment: (i) embedding the cells in BA-laden HMPs (Zhao et al. 2016; Siltanen et al. 2016) and (ii) incorporating BAs-laden HMPs inside the aggregates as the building blocks, through co-culturing the stem cells and HMPs in a specific and efficient ratio (Bratt-Leal et al. 2013; Esfandiari et al. 2017) (Fig. 3). In both these methods, the inside-out delivery of critical soluble factors throughout the aggregates can lead to homogeneous differentiation into the desired lineage. The applications and potentials of each mechanism in stem cell differentiation are comprehensively discussed below.

Fig. 3.

HMP-based strategies for cell niche reconstitution and stem cell differentiation. BA-laden HMPs and cell-BA-laden HMPs generated by DBM devices with different operational strategies for cell aggregate formation and stem cell differentiation

Cell-BA-laden HMPs

A new trend in DBM technology is generating cell-BA-laden HMPs that can precisely mimic the surrounding 3D ECM for stem cells. Compared to the macroscale scaffolds, cell encapsulation in HMPs benefits from higher cell viability due to the more efficient exchange of oxygen and nutrients resulted by the higher surface-area-to-volume ratio (Choi et al. 2016b; Tsou et al. 2016). In this approach, a mixture of stem cells, BAs, and the hydrogel precursor is injected into the DBM device and emulsified in a continuous oil phase to separately produce cell-BA-laden droplets on-chip, which are then gelled by adding the cross-linking agent to increase their stability and strength (Fig. 3) (Alessandri et al. 2016). Unlike the cell-free HMPs, for producing cell-encapsulated HMPs, the gelation process must be implemented in a mild and gentle fashion to avoid cellular damages. In this strategy, the hydrogel acts as a 3D ECM that encompasses stem cells and exposes them to a suitable concentration of critical soluble factors in a long-term period. Encapsulated cells can experience similar physicochemical properties of the native tissue by tuning different characteristics of HMPs explained in the previous sections.

Paracrine/autocrine signaling is one of the effective parameters in stem cell fate specification. Regarding this, stem cell aggregation within the HMPs is a more preferable method in terms of preserving, instead of immediately washing, the paracrine and autocrine factors in a more physiologically realistic period of time (Wozniak and Chen 2009; Lesher-Perez et al. 2013). Many different papers have proved the competence of this method for in vitro stem cell differentiation into specific lineages. For instance, in a recent study, Hou and colleagues co-encapsulated bone morphogenetic protein-2 (BMP-2) and hMSCs into the HMPs generated by the DBM device and observed enhanced osteogenic differentiation of hMSCs as a result of controlled release profile of the GFs (Hou et al. 2018). In another study, Siltanen et al. demonstrated that encapsulated-mESCs within the heparin-containing hydrogel particles loaded by Nodal and FGF-2 could differentiate into the endoderm lineage (Siltanen et al. 2016). Furthermore, Mansouri et al. (2017) utilized a similar technique to derive primordial germ cells from the stem cells encapsulated in a mixture of alginate-collagen IV microspheres.

Cell-free BA-laden HMPs

Generally, cell-cell interactions can be divided into two categories: (i) homotypic interactions between the same stem cells; (ii) heterotypic interactions between the stem cells and niche cells. These niche cells secrete signaling molecules such as Wnt and BMP that are transferred toward the stem cells through the specific channels to regulate stem cell behavior (Scadden 2006; Yeung et al. 2011). Therefore, emulating the activities of niche cells seems essential for directing stem cell differentiation. In this strategy, the output HMPs from the microfluidic system are the cell-free BA-contained compositions that are then mixed with stem cells in a specific ratio and are placed in a specific container to generate cellular aggregates (Fig. 3). The smart HMPs incorporated within the stem cell aggregates can faithfully reconstitute the role of mentioned niche cells, and regulate stem cell differentiation into the target lineage by organizing natural-like spatial and controlling the release rate of GFs in a long term (Sundararaghavan et al. 2011; Lienemann et al. 2012). More specifically, these incorporated BA-laden HMPs can successfully recreate the role of the cells that secrete the soluble factors, regulate mechanical properties of the aggregates, and reconstruct the ECM-cell interactions while maintaining the cell-cell interaction at a high level (Scadden 2006; Yeung et al. 2011; Abbasi et al. 2018). Further, through applying mixed particle populations with different cargoes and various releasing profiles, multiple BAs can be delivered to the stem cell aggregates with individual scheduling and dose. MP-mediated delivery of BAs can increase their lifetime in the laboratory by sustaining release over time (Esfandiari et al. 2017). For instance, Esfandiari et al. incorporated BMP4-laden alginate sulfate microparticles into the stem cell aggregates for in vitro germ cell induction from ESCs. They showed that the sulfated groups of their synthetic hydrogel act as specific binding sites for GFs and lead to a sustained, localized, and long-lasting delivery of loaded bioactive agents during ESC differentiation into GC. Their results revealed that compared to the soluble delivery approach, incorporation of BMP4-laden MPs incorporated within the cell aggregate increases both the germ cell differentiation efficiency twofold and their developmental potential as shown by a significant enhancement in MP-induced germ cells for meiosis entry (Esfandiari et al. 2017). In 2014, Hayashi et al. could homogeneously incorporate gelatin hydrogel microspheres into the MSC aggregates in order to improve the viability and differentiation. Expectedly, they observed that the viability and the differentiation of the cells have been significantly increased in HMP-incorporated aggregates compared to the aggregates without HMPs as a result of better oxygen, and nutrient supply (Hayashi and Tabata 2011). In another study, Dang and colleagues incorporated two different BA-laden HMPs with different release rates to improve the efficiency of differentiation (Dang et al. 2016). They loaded TGF-b1, and BMP-2 in gelatin and mineral-coated hydroxyapatite microparticles for rapid and more sustained release rates, respectively. They reported that incorporating the GF-loaded MPs into the aggregates enhances the differentiation efficiency compared to the cell-only aggregates. Carpenedo et al. (2009) have reported that HMP-mediated delivery of morphogenic factors directly within embryoid bodies (EBs) results in homogeneous, synchronous, and organized ESC differentiation while applying soluble morphogen treatment culminates in disordered and heterogeneous cell differentiation.

The mechanical properties of the cellular aggregates, as a key factor in specifying the stem cell behavior, are affected by incorporated HMPs through three different mechanisms (Abbasi et al. 2018). First, as a result of HMP incorporation, some of the cadherin-mediated cell-cell interactions can be replaced with integrin-mediated cell-MP contacts which affect the interfacial tensions and contractile forces, cytoskeletal organization, and, thus, cell’s fate depending on the MP stiffness (Guilak et al. 2009; Treiser et al. 2010; Chen et al. 2014; Turlier and Maître 2015). Secondly, incorporated HMPs, as the localized scaffolds with tunable elastic modulus, can activate mechanotransduction pathways and influence the cell fate through determining the mechanical characteristics of the whole composite aggregate (Engler et al. 2006). Finally, since the rigidity and resistance against the condensation force of the aggregate formation, incorporated HMPs can operate as the stress concentration points and distribute the stresses throughout the aggregates. The internal condensation force, as a specific characteristic of 3D cell aggregates, is generated during the stem cell self-assembling and condensation processes to generate condensed aggregates. This force is exerted on the incorporated HMPs through cellular interactions with HMPs such as integrin-based bindings. The magnitude of this force has been measured by analyzing the deformation of the incorporated microdroplets from the initial spherical shape (Campàs et al. 2014; Móczó and Pukánszky 2016; Abbasi et al. 2018; Gutierrez et al. 2021). Taken together, through the incorporation of HMPs inside the aggregates, three major advantages will be obtained: (i) The BAs can be homogenously delivered throughout the aggregates in an inside-out, localized, sustained, and controllable manner; (ii) the mechanical properties of the cell aggregate microenvironment can be tuned similar to the native tissue to more accurately mimic the in vivo condition (mechanical regulation); (iii) the incorporated cell-sized HMPs can operate as the spacers between the cells to enhance the delivery of soluble factors like oxygen and nutrients to the central parts of the aggregates (mass transfer regulator) (Hayashi and Tabata 2011) (Fig. 4).

Fig. 4.

Biochemical and biophysical effects of incorporated-HMPs within the cellular aggregates. (a) Outside-in BA delivery in HMPs-free stem cell aggregates. High cell-cell contact through cadherins leads to higher adhesion forces, higher tension forces, and more compacted and spherical aggregates. Tight cell-cell junctions act as diffusion barriers and restrict the penetration of critical soluble factors to the central core of aggregates leading to the heterogeneity of differentiation. (b) Stem cell aggregates with incorporated-HMPs for homogeneous inside-out BA delivery. Incorporation of BA-laden HMPs inside the stem cell aggregates affects the differentiation process through three different ways: (i) controlled, localized, and sustained release of critical BAs throughout the aggregates; (ii) acting as spacers between the cells to enhance oxygen and nutrient penetration to the core of stem cell aggregates; (iii) regulating the mechanical properties of the cell aggregate microenvironment and triggering mechanotransductional pathways (replacing cadherins with integrins, and acting as stress concentration points to change stress distribution)

HMP/cell ratio and HMP architecture

In this approach, the HMP/cell ratio plays a key role in aggregate formation/disruption, and significantly affects the differentiation process. Therefore, adopting an optimal ratio is critical for minimizing disruption of aggregate integrity and cell interactions since, for instance, increasing the HMP to cell ratio raises the potential for HMPs to form agglomerates (Bratt-Leal et al. 2011; Sullivan 2015; Nguyen et al. 2016). Likewise, the total aggregate stiffness will be increased by enhancing this ratio leading to different differentiation pathways. Depending on the cell type, various amounts of HMPs can be incorporated within the aggregates. Table 1 summarizes the studies that investigated the effect of HMP size and HMP-to-cell ratio in aggregate formation and cellular differentiation. This ratio alongside with size and architecture of HMPs can also influence the BA release rate, in addition to the biomaterial characteristics such as porosity, cross-linking level, mesh size, and binding strength which were discussed in the previous sections (Fig. 5) (Berkland et al. 2002; Bratt-Leal et al. 2013; Li et al. 2018). For instance, Wang et al. demonstrated that in a certain time (4 weeks), smaller HMPs have a higher release of BMP2 due to the increased surface-area-to-volume ratio and shorter diffusional barrier (Wang et al. 2012). By increasing the size, fewer particles can be incorporated between the cells, and therefore, the total contact surface and GF release rate will be reduced (Berkland et al. 2003; Siepmann et al. 2004). Applying droplets with a specific core-shell architecture, composed of a liquid core surrounded by a solid shell, can significantly enhance the loading percent of bioactive agents (either hydrophilic or hydrophobic) inside the core of droplets to prolong the release process compared to the solid particles. In this case, the release profile of GFs can be regulated relying on the characteristics of the shell, as the diffusion barrier, including the thickness, the pore sizes, and the biodegradation rate (Deveza et al. 2015; Choi et al. 2016a; Li et al. 2018). For instance, by the incorporation of solid precipitates inside the shell and blocking the small pores, the release rate of agents can be significantly reduced (Zhao et al. 2011). Also, increasing the shell thickness is associated with a low release rate and prolonged duration (Yang et al. 2017b). Taken together, it can be concluded that the use of HMP-mediated delivery of BAs within cell aggregates not only provides the homogenous diffusion of BAs but also mimics cell-cell, cell-niche, and paracrine interactions via providing sustained release and natural-like spatial organization. Therefore, because of such capabilities to niche-multimodal imitation, this approach may be a more appropriate system for in vitro germ cell production and gametogenesis.

Table 1.

Summary of investigations that utilized different MP-cell ratio to perform differentiation process

Polymer	Cell type	MPs’ size	MP:cell	Growth factor	Ref.
Gelatin Heparin-gelatin	Embryonic stem cell	–	1:3	BSA BMP4 Noggin	(Bratt-Leal et al. 2013)
Agarose	Pluripotent stem cell	5.2 ±2.8 µm	2:1	–	(Bratt-Leal et al. 2011)
PLGA		4.5±1.7 µm	1:4
Gelatin		5.1±2.9 µm	1:8
Chondroitin sulfate	Human mesenchymal stem cell	10 μm	3:1	TGF-β	(Goude et al. 2014)
PLGA coated with human ESC lysate	Human embryonic stem cell	13 μm	1:50	BMP2	(Qutachi et al. 2013)
			1:50 to 1:3000	VEGF
			was investigated
pNIPMAma, (HMAM)b	Pluripotent stem cell	6.4±1 µm	1:10	BMP4	(Sullivan 2015)
			1:3
			1:1
			3:1
PLGA coated with gelatin	Embryonic stem cell	1 µm	2×106 cells (in 10 ml) + 2 mg MPs	RAc	(Carpenedo et al. 2010)
		3 µm
		11 µm
Gelatin	Human mesenchymal stem cell	3.55±1.05 μm	1:8	–	(Baraniak et al. 2012)
Gelatin-coated PLGA, PLGA	Embryonic stem cell	7.8±1.8 µm	1:2	RA	(Carpenedo et al. 2009)
			2:1
			5:1
			3 µg RA per mg of PLGA

^aPoly(N-isopropyl methacrylamdie)

^Bheparin-medethacrlamide

^cRetinoic acid

Fig. 5.

Various effective parameters of HMPs in BA release rate and stem cell differentiation. (a) The effect of biomaterial type: the polymer type and its associated modifications can physically and chemically affect the release rate of BAs as well as paracrine interactions. (b) The effect of HMPs size: By increasing the HMP size, fewer particles are incorporated between the cells, and the total contact surface is also reduced leading to a lower BA release rate. Also, larger particles have a longer diffusion barrier leading to a more sustained release of BAs. (c) The effect of HMP/cell ratio: increasing the number of HMPs, when the total cell number is constant, enhances the total release rate of BAs. (d) HMP architecture: application of HMPs with more complex architecture, such as core-shell HMPs, enables higher BAs loading, and thus, higher release rate

Conventional methods for HMP production

The slightest variations in size, composition, and structure of HMPs impact the delivery performance and BA release profile which in turn affect the stem cell niche and the cell differentiation process (Thiele 2016; Li et al. 2018). Therefore, a preferable method with high controllability in the particle-generation process is required to produce HMPs with precise characteristics relevant to their applications in biological studies. There are a variety of strategies utilized for HMP fabrication including emulsion polymerization, electrohydrodynamic spraying, mechanical fragmentation, and DBM devices. Herein, we briefly explain the limitations and strengths of these techniques in producing advanced HMPs suitable for stem cell differentiation, although there are some excellent review papers for Enthusiasts that have comprehensively investigated different aspects of current HMP fabrication methods (Choi et al. 2017; Daly et al. 2020).

Emulsion polymerization

Conventional emulsion polymerization, also known as batch emulsion, is the most common technique for HMP fabrication. The technique enables the simple and high-speed production of HMPs in a short time. In this method, the hydrogel prepolymer, cross-linking agent, and surfactant are added to an oil phase. Then, mechanical agitation such as high-speed stirring is utilized to generate the necessary shear stresses for breaking up the droplets and forming a microparticulate water-in-oil emulsion (Liu et al. 2014; Li et al. 2016). The droplet size can be determined by tuning the time and the intensity of the mixing process. This method suffers from several major drawbacks including large polydispersity, weak reproducibility, and inability in tuning internal and external architecture to produce advanced HMPs as a result of low controllability over the individual droplet generation (Xu et al. 2009; Gupta et al. 2017; Li et al. 2018). Polydisperse droplets propose different diffusion rates resulting in a non-uniform release rate of BAs, and thus, non-uniform stem cell differentiation. Also, in the case of cell-laden HMPs, due to the size variation of the HMPs, the precise control of cell numbers within HMPs is impossible. Further, since the advanced HMPs with the multi-compartmental structure are great candidates for more efficient differentiation through delivering multiple critical BAs with individual release profiles, the inability of the emulsion polymerization method to produce these advanced HMPs makes them less attractive for efficient stem cell differentiation applications. In addition, random loading of BAs in HMPs leads to huge waste and makes it uneconomical for expensive BAs such as BMP4 (Esfandiari et al. 2017).

Electro-hydrodynamic spraying

Electro-hydrodynamic spraying is another conventional method employed to produce HMPs using the electrical field (Bock et al. 2011; Qayyum et al. 2017; Gansau et al. 2018). In this method, the hydrogel precursor is pumped through a syringe while an electric field is externally exerted to the system. As a result of applied voltage, the surface tension of the droplet at the needle tip is overcame by the electrostatic force leading to the breakup of the droplet into the small droplets. Then, the charged small droplets are collected and cross-linked in a collector bath (Naqvi et al. 2016). The droplet size can be controlled by determining various parameters including polymer concentration, solution viscosity, flow rate, applied voltage, and needle diameter (Choi et al. 2017; Gansau et al. 2018). The main strength of this approach is its ability to produce small MPs with size as little as 0.1 µm (Pancholi et al. 2009; Choi et al. 2017). However, similar to the batch emulsion method, this technique generates polydisperse droplets with dispersity indexes higher than 5%, which restricts its usage for efficient HMP-based stem cell differentiation (Young et al. 2012; Qayyum et al. 2017).

Mechanical fragmentation

Contrary to the two previous techniques in which the cross-linking process is fulfilled after the droplet formation, in the mechanical fragmentation method, a preformed bulk hydrogel is mechanically broken up into the microdroplets with the minimum size of 15 µm (Sinclair et al. 2018). The breaking up step can be implemented by either applying a rotational blender or deriving the hydrogel through the steel meshes with predefined pore sizes. Due to the low controllability of this approach over the droplet generation process, specifying the characteristics such as size, shape, and internal architecture is impossible. Although this technique benefits from simplicity and high production rate, producing cell-laden HMPs seems impractical due to the high shear stresses generated in the breaking up step. Taken together, all these disadvantages make this technique inappropriate for being used in HMP-based stem cell differentiation. To circumvent the limitations of conventional methods, DBM systems have recently been introduced as the most efficient and sophisticated platform for high-throughput production of advanced HMPs with tunable compositions and geometrical characteristics in a highly controllable manner (Kim et al. 2014; Lee et al. 2016; Wang et al. 2017a; Li et al. 2018). The comparison of various 3D strategies applied for directing stem cell differentiation is summarized in Table 2.

Table 2.

Summary of different 3D cultures with growth factor delivery strategies using diverse materials for stem cell differentiations to target lineage

Strategies	Source cell type	Target lineage	Applied signaling mediators	Applied hydrogel	Delivery techniques	Reference
Soluble delivery	Embryonic stem cells	Mesoderm, ectoderm	BMP4, noggin	–	–	(Bratt-Leal et al. 2013)
	Embryonic stem cells	PGC	BMP4 BMP8b	–	–	(Hayashi et al. 2011)
Emulsion polymerization	Embryonic stem cells	Germ cell	BMP4	Alginate sulfate	MPs incorporation within cellular aggregates	(Esfandiari et al. 2017)
	Adipose stem cells	Nucleus pulposus	RhGDF-6, human serum albumin	PLGA polyethylene glycol (PEG)-PLGA	Bioactive microgel in a 3D cell-laden collagen	(Hodgkinson et al. 2019)
	Embryonic stem cells	Visceral endoderm	RA	PLGA	MP incorporation within cellular aggregates	(Carpenedo et al. 2009)
	Human embryonic stem cells	Vascular	VEGF, BFGF, PIGF	PLGA	MP incorporation within cellular aggregates	(Ferreira et al. 2008)
	Embryonic stem cells	Osteoblast, endothelial	BMP2, VEGF	PLGA	MP incorporation within cellular aggregates	(Qutachi et al. 2013)
	Embryonic stem cells	Endoderm, mesoderm	ECM proteins	Agarose, PLGA, gelatin	MP incorporation within cellular aggregates	(Bratt-Leal et al. 2011)
	Human embryonic stem cells	Chondrocyte	TGF-β1, BMP2	Gelatin, mineral-coated hydroxyapatite	MP incorporation within cellular aggregates	(Dang et al. 2016)
Electrospraying	Pre-osteoblast cells	Osteoblast	VEGF, BMP7	PLGA	Bioactive microgel	(Carpenedo et al. 2009)
	Bone marrow stromal stem cells	Osteoblast	BMP2, SDF-1	Alginate-chitosan	Bioactive microgel	(Xu et al. 2018)
Droplet-based microfluidics	Mouse embryonic stem cells	Endoderm	Nodal FGF-2	PEG+ Heparin	BA-stem cell laden microgels	(Siltanen et al. 2016)
	Human dental mesenchymal stem cells	Chondrocyte	TGF-β1	RGD-modified alginate	BA-stem cell laden microgels	(Moshaverinia et al. 2013)
	Human neuronal stem cells	Neuron	ECM proteins	Alginate + ECM	Stem cell encapsulation within matrigel-coated microcapsules	(Alessandri et al. 2016)
	Bone marrow-derived mesenchymal stem cells	Osteoblast	BMP2	GelMA	BA-stem cell laden microgels	(Zhao et al. 2016)
	Human mesenchymal stem cells	Osteoblast	BMP2	Poly(vinyl alcohol)	BA-stem cell laden microgels	(Hou et al. 2018)

Droplet-based microfluidic systems

Nowadays, DBM technology is a well-known technology with widespread applications in directing stem cell differentiation and developmental biology (Moreno et al. 2015; Samal et al. 2019). Their great abilities in precise manipulation of multiple fluids at the millimeter-to-micrometer scale and balancing the acting forces make them extremely attractive for high-throughput, automatic, and standard generation of HMPs. Producing advanced HMPs with well-defined size and shape, tailored structures and features, tunable physical and chemical properties, minimum material waste, high-level monodispersity, and integrated functions for programmable release of BA are some of the abilities of designable DBM devices as a result of their high controllability over the droplet generation process (Champion et al. 2007; Campos et al. 2013; Zhang et al. 2017; Li et al. 2018). The uniform HMP production is important because it enables (i) the predictability of BA encapsulation amount and releasing profile in the case of cell-free BA-laden HMPs; and (ii) controlling the cell number in each HMPs in cell-laden HMP strategy (De La Vega et al. 2013; He et al. 2019). DBM devices are able to produce monodisperse droplets with the minimum size range of 5–10 µm, the maximum size of 500 µm, and the dispersity indexes of 1–2% (Bardin et al. 2013; Campàs et al. 2014; Pittermannová et al. 2016). This is an acceptable range with appropriate accuracy for HMP-base stem cell differentiation. This technology enables producing HMPs with the multi-compartmental structure for separate co-encapsulation and synergistic delivery of multiple BAs (both hydrophobic and hydrophilic agents) and independent release rate in a spatiotemporal fashion, which is required for stronger and more specific differentiation (Lesher-Perez et al. 2013; Teo et al. 2016). The fabrication process of HMPs in DBM devices occurs due to the difference in flow rate between the oil and aqueous phases (Shembekar et al. 2016). Simultaneous injection of dispersed and continuous phases into the separate microchannels and balancing between the parameters of surface tension and viscous shear stress result in the droplet formation of the dispersed phase. More precisely, at proper flow rates, once the drag force surpasses the interfacial tension, the droplets of the dispersed phase are broken into specific and uniform sizes in a “one droplet at a time” fashion (Yu et al. 2010; Deng et al. 2011a; Li et al. 2018). The shape, size, and production rate of the HMPs can be optimized by tuning the flow rates, the viscosity and density of dispersed/continuous phases, surfactant properties, and channel dimensions and geometry (Leng et al. 2010; Li et al. 2018). To prevent the coalescence of microdroplets, the cross-linking process must be implemented as soon as possible after droplet formation through either on-chip or off-chip strategies. There are different cross-linking mechanisms reported in the literature; the most prominent methods for gelation in microfluidic devices are chemical agents (alginate), photo-cross-linking (GelMA, PEGDA, and Sil-MA), and thermal assembly (gelatin and agarose) (Xia et al. 2016; Mohamed et al. 2020). The photo-cross-linking strategy is a simple method for gelling droplets in transparent microfluidic devices. However, this method is not appropriate for cross-linking cell-laden droplets since it may injure the cells as a result of ultra-violet (UV) exposure (Pfeifer et al. 2005; Huang et al. 2017). Alternatively, delivering chemical agents is a more appropriate strategy to cross-link/polymerize the aqueous precursor and gelation (Wu et al. 2016). Gelation of sodium alginate droplets via Ca²⁺ is the most notable example of this strategy. The thermal assembly, also, can be utilized to polymerize hydrogels such as collagen, and agarose. Since the low volume of microfluidic devices, the temperature can simply and quickly set to the desired point required for the gelation process (Mazaheri et al. 2016; Wang et al. 2016).

On-chip manipulation strategies

Despite the substantial progressions in advanced droplet generation with new upgraded versions of DBM systems, they are not yet applied in the stem cell differentiation field. The DBM devices currently utilized for stem cell differentiation have several important limitations. By considering the recent advancements in DBM devices and integrating several on-chip manipulation mechanisms, these limitations can be resolved. In this part of the review, three on-chip manipulation mechanisms, including pump-less derivation, on-chip extraction, and on-chip cell scattering components, are highlighted.

Pump-less derivation

An important limitation of most DBM platforms is their integration with pumps for driving the flow inside the device. Although the device operation can be optimized in a higher controllable fashion when applying syringe pumps, its major limitation is revealed when an extremely low volume of expensive material is needed. For instance, BMP4 is an essential and very expensive GF required for germ cell derivation. However, due to the pumping system required for launching the DBM device, a significant volume of the total hydrogel-BMP4 mixture will be wasted inside the tube and syringe. Also, the delayed response time of these systems that are caused by large tubing resistance and device material compressibility increases the required time for stabilizing particle generation time in the order of seconds to minutes leading to additional BMP4 wasting (Churski et al. 2013; Chong et al. 2016). To circumvent this problem, various pump-less DBM devices acting based on suction, centrifuge, pressure, and gravity-driven flow have been recently introduced (Chong et al. 2016; Damiati et al. 2018). For instance, Ahmed et al. could generate monodisperse micro-droplets by using a simple centrifugation-based microfluidic device. They mounted their device and reagents to a scaffold in a compact form small enough to be inserted into a 50-mL falcon tube, and then placed the tube into a common lab centrifuge for the spinning process. They could control the flow rates and, thus, droplet sizes through regulating centrifuge speed, reservoir heights, and inlet channel resistance (Fig. 6a) (Byun et al. 2014; Ahmed et al. 2016). In another study, Langer et al. (2018) developed a novel strategy for droplet formation in microfluidic devices that is so appropriate for cases with very small liquid volumes or expensive loaded agents such as BMP4. They applied a standard laboratory adjustable micropipette as a negative pressure generator for pumping the fluids from inlet pipette tips and producing homogeneous droplets from whole the sample (Fig. 6b). Another group presented a simple and efficient method to generate monodisperse microparticles by applying a hand-operated syringe to create vacuum at the outlet of the microfluidic device to suck the reagents through the channels and generate the microparticles (Abate and Weitz 2011). For controlling the particle size and production frequency, their device was equipped with hydrodynamic resistors and single-layer membrane valves to specify the channels’ hydrodynamic resistances and flow rates in real time. This method is highly suitable for droplet generation of expensive and limited materials given the significant reduction of waste percentage (Fig. 6c). Several similar studies utilized syringe vacuum microfluidics, finger-powered devices, and 3D printed pumping lid for the generation of droplets with a minimized loss of expensive samples (Begolo et al. 2014; Iwai et al. 2014; Dutka et al. 2016).

Fig. 6.

On-chip manipulation strategies. Pump-less DBM devices acting based on (a) centrifugation, (b) micropipette-suction, and (c) negative pressure with control valves for regulating particle size. (d) A DBM device integrated with an on-chip extraction module capable of separating the droplets from the oil phase into an aqueous phase. The red and blue colors are the oil and the aqueous solutions, respectively. (e) Integration of a cell-scattering unit to the DBM device can safely disperse the cell clusters into single cells which is appropriate for accurate specification of cell numbers within each droplet. Reprinted, with permission, from Ahmed et al. (2016) (a), Langer et al. (2018) (b), Abate and Weitz (2011) (c), Deng et al. (2011b) (d), Kim (2015) (e)

On-chip extraction

The conventional method for off-chip extraction of microparticles from carrier oil into the aqueous phase is a time-consuming and labor-intensive process including several steps of off-chip centrifugation and washing. However, due to the centrifugal\compressive forces, and prolonged exposure to cell-incompatible nutrient-oxygen depleted oil, several important disadvantages appear including inefficient retrieval, microparticle aggregation, remaining oil on microparticle surface, and compromised cell viability in the case of cell-laden HMPs (Zhang et al. 2007; Deng et al. 2011b; Hong et al. 2012; Huang and He 2014). These drawbacks impact the efficiency of high-throughput germ cell derivation in the laboratory. The on-chip immediate extraction unit is considered another functional component that can be integrated with DBM systems to resolve these limitations (Deng et al. 2011b).

In the case of cell-laden HMPs, applying this module can minimize cellular injuries during the production and collection process (Sun et al. 2018). This component operates through immediate on-chip sorting of cell-laden microcapsules from cell-incompatible, nutrient, and oxygen-depleted oil phase into the aqueous phase. Up to now, several on-chip strategies have been introduced to extract microcapsules including oil phase depletion, mechanical filter extraction, interfacial tension–based extraction, and dielectrophoresis (DEP)-based extraction (Hong et al. 2012; Huang et al. 2015b; Sun et al. 2018). For instance, Deng et al. (2011b) could successfully transfer HMPs from the oil phase into the aqueous phase through continuous infusion of aqueous solution into the multiple side channels. This infusion could remove the oil carrier toward the specified outlet assigned on the opposite side. They showed that compared to the conventional methods, the cell viability has been significantly improved when the on-chip extraction module is incorporated in DBM devices. An attractive idea that they benefited from it in their device was on-chip polymerization of alginate before the extraction process by designing a separate inlet for calcium chloride (CaCl₂) injection. More precisely, their device could simultaneously generate alginate and CaCl₂ droplets and merge them before the extraction part to solidify the droplets (Fig. 6d). Another strategy is dispensing magnetic nanoparticles into the cell suspension and applying a magnetic field outside the chip to extract microgels from oil to the aqueous phase (Yoon et al. 2013). However, utilizing the magnetic nanoparticles may affect the pathway of stem cell differentiation, or even lead to strong cell apoptosis (Dai et al. 2019). As another strategy for large particles, establishing a stable interface between oil and aqueous phases leads to the transfer of the suspended microparticles toward the interface boundary which happens as a result of the slip-boundary conditions. Then, upon contacting the interface, they will be dragged into the aqueous phase due to the interfacial tension force (Huang and He 2014; Salehi et al. 2020). In another research, the same group integrated an efficient on-chip DEP-based extraction module with their DBM device to separate cell-laden hydrogel microcapsules (Hong et al. 2012; Huang et al. 2015b). Also, incorporating the mechanical filter into the device can prevent the microparticles from passing through the filter gates and sink them into the aqueous solution (Hong et al. 2012).

On-chip scattering

A further and important advantage of applying DBM devices for generating cell-contained HMPs is their better ability to control the cell numbers in each droplet, and thus, the accurate specification of required oxygen level for aggregate formation (Wu et al. 2008; Kashaninejad et al. 2016). For instance, Agarwal et al. utilized a microfluidic platform to generate core-shell microcapsules with almost 20 embryonic cells in the aqueous liquid core. Their results showed that 20 embryonic stem cells are sufficient to produce a single ESC aggregate in each droplet after a week (Agarwal et al. 2013). Even for further uniformity of cell numbers encapsulated in each droplet, and thus forming uniformly sized cellular aggregates, a cell-scattering component can be integrated with DBM devices to disperse cell clusters into the single cells. To do so, Kim designed a scattering unit that sequentially performs splitting, stretching, and bumping processes on the initial cell clusters. He demonstrated that utilizing a cell scattering component improves the uniformity of encapsulated cells in HMPs (Fig. 6e) (Kim 2015).

High-throughput culture of cell-BA-laden HMPs

DBM technology is intrinsically high-throughput which means numerous HMPs are generated each time the system is launched. In this strategy, each HMP operates as a micro-bioreactor and a suitable platform for cellular aggregate formation. Long-term culturing of cell-laden HMPs can be accomplished either off-chip or on-chip. In the case of the off-chip culturing method, after droplet generation, extraction, and collection, the cell-laden HMPs can be cultured within the common containers such as media-filled Petri dishes or multi-well plates for an extended period of time (Wang and Wang 2014; Wang et al. 2015). However, the coalescence of HMPs seems inevitable when the cell-laden HMPs are cultured in these containers (Moshksayan et al. 2018). Integrating an appropriate culture chamber downstream of the DBM devices to house the generated cell-laden HMPs is a great strategy to automatize all the processes on a single device, and to improve the performance and efficiency by reducing consumption of culture medium and providing better control and manipulation of the microenvironment. This strategy allows for rapid and high-throughput aggregation, reduces the user-device interaction, and adopts automatic procedures instead of inefficient manual handling (Tabeling 2005; Zhang et al. 2017; Moshksayan et al. 2018). Many studies have demonstrated the competence of such integrated DBM platforms for the long-term culturing of cellular aggregates while maintaining their viability at a high level (Kwapiszewska et al. 2014; Kim et al. 2015; Ruppen et al. 2015). Although this kind of integrated platforms have not been utilized for in vitro germ cell derivation, their correlated benefits make them very suitable and attractive for automatic and high-throughput stem cell aggregate formation and germ cell induction.

In the case of on-chip aggregate formation, it is highly critical to immediately remove surrounding sheath fluid after droplet formation and before entering the droplets into the anchorage sites. As a result, after droplet generation, the cell-laden HMPs are immediately extracted from the oil phase and then enter the docking chambers. The polymerization process can be accomplished either before the extraction process or after the placement inside the chambers depending on the hydrogel type and solidification mechanism (conceptual design) (Deng et al. 2011a). Up to now, various microfluidic devices equipped with droplet anchorage sites have been developed. For instance, Sabhachandani et al. developed an integrated high-throughput DBM system containing a droplet generation compartment and a docking chamber (Sabhachandani et al. 2016). The role of the first part is to produce cell-BA-laden HMPs and transfer them directly to the second part in which the cell-laden HMPs are entrapped for further generation, maintenance, stimulation, and analysis of the cellular aggregates (Fig. 7a). In another study, Tomasi and coworkers could generate and homogeneously entrap 5000 cell-laden microgels in anchoring microwells in less than 10 min (Tomasi et al. 2013) (Fig. 7b). Furthermore, McMillan and colleagues developed an integrated microfluidic platform with the ability to contain 2000 monodisperse cell-laden droplets. They could maintain the cell viability at a high level by continuous perfusion of culture medium (Fig. 7c) (McMillan et al. 2016).

Fig. 7.

Integrated DBM device with permanent\temporary anchorage sites. (a-c) Various reported designs of permanent anchorage sites integrated with DBM devices for high-throughput on-chip cell aggregate formation and storage. (d-e) The temporary anchorage sites developed relying on pneumatic structures to address the challenges of permanent anchorage sites such as channel blockage and to improve retrieval efficiency and safety. Reprinted, with permission, from Sabhachandani et al. (2016) (a), Tomasi et al. (2013) and Chemical and Biological Microsystems Society (CBMS) (b), McMillan et al. (2016) (c), Anada et al. (2010) (d), Liu et al. (2015a) (e)

Utilizing permanent anchorage sites is a common problem of most on-chip culturing devices since in this case the flow could be blocked by droplets. As reported in several studies, designing a bypass channel may be a solution to overcome this challenge (Das et al. 2013; Ruppen et al. 2014). However, since the droplet generation with desired properties is stabled after a while, the culture-specific part of the device must be out of the circuit at the beginning of the process when utilizing the bypass strategy. Otherwise, the anchorage sites are filled with initial non-homogeneous droplets in terms of size and cell numbers, and thus, no empty sites are left for desired droplets. Safe, easy, and efficient harvesting of the cell aggregates from the device after formation and differentiation into the desired lineage is another important challenge of the on-chip aggregation approach with permanent anchorage sites. Several mechanisms have been developed to address this deficiency. The main solution is replacing permanent anchorage sites with temporary ones through integrating well-established pneumatic microstructures into the DBM devices (Liu et al. 2015b, a). To do so, Anada et al. developed a novel two-layered microfluidic device equipped with a thin polydimethylsiloxane membrane that can rapidly generate thousands of hemispherical cavities through decompression of the bottom channel to deform the membrane. As a result, they could noninvasively retrieve the generated spheroids upon restoring the pressure to atmospheric pressure (Fig. 7d) (Anada et al. 2010). Similarly, Liu et al. developed a microfluidic device with a U-shaped pneumatic microstructure for dynamic micromanipulation of anchorages to efficiently retrieve the aggregates from the device (Fig. 7e) (Liu et al. 2015a). In conclusion, integrated DBM devices with temporary anchorage sites and on-chip extraction module can be considered a promising strategy for high-throughput in vitro germ cell derivation.

High-throughput generation of HMP-incorporated cellular aggregates

Off-chip aggregate formation

The current methods for cell aggregate formation with incorporated-HMPs are based on off-chip strategy so that after BA-laden HMP generation, collection, and extraction, they are mixed with stem cells in a specified ratio and then cultured in specific containers such as hanging drop platforms or round bottom microwell plates to enforce the cell-HMP accumulation. However, these methods are non-automated, inflexible, time-consuming, labor-intensive, low-throughput, and susceptible to variability and error making them inappropriate for in vitro germ cell induction (Aijian and Garrell 2015). On the other hand, being high-throughput is the primary characteristic of DBM systems leading to the fast production of millions of HMPs in a short time, while conventional aggregate generators are low-throughput. Therefore, for applying all the generated HMPs and less waste of expensive BAs, common methods for 3D cell aggregate generation must be improved. In recent decades, microfluidic technology has been drastically improved in providing sophisticated platforms for controllable and high-throughput generation of 3D cell aggregates as well as their culture and harvesting. Since the ability to design microfluidic devices, various strategies with different stimulations such as gravity, magnetic force, acoustic, and rotation have been adopted to form and culture cellular aggregates (Alhasan et al. 2016; Cottet et al. 2019; Sudeepthi et al. 2019). For instance, to eliminate high-risk and tedious manual handling of the hanging drop method, digital microfluidic systems have been recently utilized to flexibly and automatically handle liquids for high-throughput droplet generation (Aijian and Garrell 2015). In fact, this technology acts as an automatic hanging drop method by automating the droplet generation, 3D aggregate formation and maintenance, culture medium exchange, and addition/removal of any solution (Fig. 8a) (Bender et al. 2016). Therefore, unlike the conventional hanging-drop method, long-term 3D aggregate culture will be possible in an automatic and high-throughput manner when using digital microfluidic systems. Digital microfluidics utilizes electrical fields to spatiotemporally manipulate the liquids for discrete droplet generation, translation, splitting, mixing, and merging relying on electro-wetting properties of the liquids (Fair 2007; Jones 2008; Nelson and Kim 2012). For instance, Aijian et al. (Aijian and Garrell 2015) developed a digital microfluidic to produce cell-laden droplets and generate cellular aggregates. Also, they could automatically exchange the culture medium around the spheroids without extracting them. Another important capacity of digital microfluidic systems is their ability to simply retrieve aggregates from the device after formation and differentiation for further processing. The same strategy can also be applied for the mixture of stem cells and HMPs instead of only-cell suspension. According to this, cell-HMP suspension is placed on the device surface for further automatic splitting, and cell-HMP-contained droplet generation. By applying parallel electrodes in a digital microfluidic device, many droplets can be simultaneously and automatically generated. The valuable capabilities mentioned above have made digital microfluidics more appropriate than conventional hanging-drop for high-throughput germ cell induction in the laboratory. Another off-chip high-throughput strategy is the injection of cell/HMP suspension into the specific microfluidic devices with lots of round-bottom microwells. As an example, Chen et al. (2015) developed a microfluidic aggregate formation platform fabricated by a novel poly-2-hydroxyethyl methacrylate (poly HEMA) (non-adherent polymer) to quickly generate 1024 aggregates within the microchambers (Fig. 8b). In another study, Chen et al. introduced a high-throughput acoustic aggregate fabrication method that can dynamically produce 12000 multicellular aggregates within several minutes. Despite its major advantages such as high efficiency, contactless and contamination-free condition, size controllable, and excellent biocompatibility, this system suffers from two major limitations in producing stem cell aggregates with incorporated HMPs: (i) acoustic stimuli may affect cell differentiation and (ii) the ratio of cell-HMP could not be accurately determined (Fig. 8c) (Chen et al. 2019). A minor limitation of off-chip strategies is their inability to accurately determine cell-HMP ratio while this ratio is a critical parameter for the stem cell aggregate formation and germ cell induction. Therefore, although many different off-chip strategies have been developed for cell-only aggregate formations, they are not the best choices for the generation of HMP-incorporated aggregates due to their related limitations.

Fig. 8.

Various high-throughput microfluidic systems for off-chip aggregate formation. (a) Digital microfluidics acting as an automatic hanging drop method to flexibly and automatically handle cell suspension for high-throughput droplet generation. (b) Microwell array microfluidic system. This chip can generate 1024 aggregates within the microchambers by using a novel polyHEMA. The fabricated device contains 1024 microwells within a core area of 2 cm by 2 cm. The cells are loaded in the inlet and then flow through the microwell array to the outlet. (c) Acoustofluidic cell aggregate fabrication chip. The embedded acoustic wave generator on this device aids the cell aggregates generation in a high-throughput manner. Reprinted, with permission, from Bender et al. (2016) (a), Chen et al. (2015) (b), Chen et al. (2019) (c)

On-chip aggregate formation

Inspired by cell-laden HMPs in which each HMP provides an appropriate platform for culture, growth, and differentiation of cells, applying liquid-core/solid-shell HMPs seems a promising strategy for germ cell derivation by producing high-throughput and automatic stem cell aggregates with incorporated BMP4-laden HMPs. The ability of this strategy to modulate the permeability of shell structure for the suitable exchange of nutrients and oxygen enables a long-term culture of generated stem cell aggregates (Chen et al. 2016). In a two-step strategy, after droplet generation, separation, and collection from the first DBM device and then mixing with stem cells in a favorable ratio, the cell-HMP suspension is injected in the second DBM device to be encapsulated in liquid-core/solid-shell HMPs and cultured in the downstream culture chambers. However, such a two-step approach is often laborious. Microfluidic systems are highly potential to be specifically designed for integrating all mentioned steps on a single device. Figure 9 shows the conceptual design of an integrated DBM device for both on-chip generation and culturing of stem cell aggregates with incorporated BA-laden HMPs in a high-throughput and automatic manner. As is clear in the figure, this device is composed of two series of droplet generator components, a cell scattering component, two extraction units, one pneumatic valve, and the culture chambers. The pneumatic valve has been integrated into the system to control the size and the frequency of droplet production. Integrating the cell-scattering module helps to provide a completely uniform cell suspension without clusters. Both of these modules improve the efficiency of germ cell derivation through enhancing the accuracy of the cell-HMP ratio in each droplet. At the end of the process, a large number of core-shell droplets containing stem cells and BA-laden HMPs are generated and placed in anchorage sites in a very short term. After the sedimentation of stem cells and microparticles, cell-cell interactions are increased and the compact aggregates are spontaneously formed while containing BA-laden microgels. We believe that through integrating all the crucial components on a single device, this platform meets all the requirements for high-throughput, automatic, and efficient in vitro stem cell differentiation into the germ cells.

Fig. 9.

Conceptual design of an integrated DBM device for both on-chip generation and culturing of stem cell aggregates with incorporated BA-laden HMPs. This platform is composed of five inlets; I₁: the inlet for BA-laden hydrogel; I₂ and I₅: the inlet for oil phase; I₃: the inlet for cell suspension; and I₄: the inlet for shell hydrogel. The object of the pneumatic valve is controlling the droplet size and production frequency. Two extraction modules are considered for on-chip extraction of HMPs from the oil phase into the aqueous phase. A cell scattering module is designed after cell suspension inlet to disperse cluster cells into single cells. The culture chambers are also integrated into the device to contain the core-shell HMPs for on-chip culturing, aggregate formation, and differentiation up to the germ cells

Future perspectives

The marriage of stem cell biology and engineering technologies seems highly valuable to make a bright future for germ cell research. In recent years, promising advances of DBM systems have revealed their great potential in producing smart and advanced HMPs which suggest them as the sophisticated platforms for efficient in vitro germ cell induction. Advanced HMPs with tunable morphology, structure, and characteristics have enabled controlled and sustained delivery of critical BAs. These platforms have proved their competence in faithful replication of 3D natural microenvironment and cellular niches. Although to the best of our knowledge these platforms have not yet been applied for in vitro germ cell derivation, they have shown promising outcomes in stem cell differentiation into other lineages such as endoderm, chondrocyte, osteoblast, and neuron. According to this, comprehensive investigation and deep understanding of current advances, limitations, and potentials of engineering technologies utilized in other fields would enable the scientists to employ them for the development of more efficient procedures for in vitro germ cell induction from stem cells.

Supplementary Information

Below is the link to the electronic supplementary material.

Author contribution

R. S. H. and M. A. H. contributed equally to conceiving and writing the paper; Z. G., M. A. H., and R. S. H. designed and produced the figures; and F. E. supervised the work.

Funding

This work was supported by a grant provided by the Iran National Science Foundation (INSF; grant number: 97001442) and Royan Institute (grant number: 96000166).

Declarations

Conflict of interest

The authors declare no competing interests.