- Open Access
Efficient generation of hepatic cells from mesenchymal stromal cells by an innovative bio-microfluidic cell culture device
© The Author(s). 2016
- Received: 20 April 2016
- Accepted: 20 July 2016
- Published: 19 August 2016
Mesenchymal stromal cells (MSCs) are multipotent and have great potential in cell therapy. Previously we reported the differentiation potential of human MSCs into hepatocytes in vitro and that these cells can rescue fulminant hepatic failure. However, the conventional static culture method neither maintains growth factors at an optimal level constantly nor removes cellular waste efficiently. In addition, not only is the duration of differentiating hepatocyte lineage cells from MSCs required to improve, but also the need for a large number of hepatocytes for cell therapy has not to date been addressed fully. The purpose of this study is to design and develop an innovative microfluidic device to overcome these shortcomings.
We designed and fabricated a microfluidic device and a culture system for hepatic differentiation of MSCs using our protocol reported previously. The microfluidic device contains a large culture chamber with a stable uniform flow to allow homogeneous distribution and expansion as well as efficient induction of hepatic differentiation for MSCs.
The device enables real-time observation under light microscopy and exhibits a better differentiation efficiency for MSCs compared with conventional static culture. MSCs grown in the microfluidic device showed a higher level of hepatocyte marker gene expression under hepatic induction. Functional analysis of hepatic differentiation demonstrated significantly higher urea production in the microfluidic device after 21 days of hepatic differentiation.
The microfluidic device allows the generation of a large number of MSCs and induces hepatic differentiation of MSCs efficiently. The device can be adapted for scale-up production of hepatic cells from MSCs for cellular therapy.
- Mesenchymal stromal cells
- Hepatic differentiation
- Microfluidic device
Liver transplantation is the only definitive treatment for end-stage liver diseases . However, the lack of donor organs has been a major obstacle to the treatment modality. Besides, the long-term use of immuno-suppressants after liver transplantation also brings undesired adverse effects . For these reasons, alternative treatment options are desired. Cell therapy, which offers a potential solution to the problem of donor shortage, has become highly attractive in recent years along with the rapid progress in stem cell technology. However, the unsolved ethical concerns and controversies over the use of human embryos to derive embryonic stem cells (ESCs) have substantially hampered the progress of their clinical applications, not to mention that ESCs are allogeneic in nature and may require further immue-suppression treatment after implantation. Mesenchymal stromal cells (MSCs), which can be obtained from a number of somatic tissues including adipose tissue [3, 4], amniotic fluid [5, 6], placenta , and umbilical cord blood , are easily available. We have previously reported the differentiation potential of human MSCs into hepatocytes in vitro based on a novel two-step protocol [8, 9]; these cells can rescue fulminant hepatic failure induced by CCl4 . Recently, a number of studies have used similar protocols for hepatic differentiation of MSCs based on our previous reports [11–14]. About 4–6 weeks are required to complete the hepatic differentiation process with these protocols. In addition, the need for a large number of hepatocytes for cell therapy has not been addressed fully in these studies.
Shear stress and fluid friction force generated from continuous fluid flow may significantly affect hepatic differentiation of MSCs because liver regeneration in vivo is related to portal pressure, reflecting fluid shear stress . However, very few studies have examined the effects of fluid shear stress on hepatic differentiation. Therefore, it is valuable to investigate hepatic differentiation of MSCs under continuous fluid flow using a microfluidic device which mimics the shear flow in the microenvironment. Microfluidic devices have been utilized in cell culture [16–18], cell differentiation [19–22], dynamic gene expression [23, 24], and test of cellular response to chemical gradients [25–27]. Studies of hepatic differentiation from MSCs using microfluidic devices  and bioreactors [28–40], together with 3D scaffolds , have been reported. However, there are shortcomings when using conventional microfluidic devices to promote hepatic differentiation. First, an uneven flow in a small culture chamber may affect the quality and yield of differentiated cells . Second, cell injection through seeding microchannels causes an uneven cell distribution which could affect cellular interaction and therefore influence differentiation [17, 19, 20, 26, 42–48]. In addition, when cell injection is completed during the injection processes, some cells still remain in the injection microchannel and are not delivered into the culture chamber. Thus, it is imperative to develop an innovative microfluidic device that minimizes these limitations.
To solve the aforementioned problems, we developed a novel bio-microfluidic device system to enhance the hepatic differentiation capacity of MSCs. The microfluidic device can also improve cell distribution and scale-up cell productions. We further compared the performance of the new system with that of the classic system.
Design and fabrication of the microfluidic device and culture system
The microfluidic device was designed to have a culture chamber dimension of 10 mm × 40 mm × 350 μm (width × length × height), with a culture area of 400 mm2. The device was assembled in five layers (Fig. 1) consisting of a lower layer of a culture substrate, on top of an intermediate layer formed by two patterned glass and two patterned polydimethylsiloxane (PDMS) membranes (Sylgard 184; DowCorning, Midland, MI, USA), with a top layer of polymethyl methacrylate (PMMA), including three adaptors for producing the vacuum, medium inlet, and outlet. The PDMS membranes were prepared and fabricated according to the manufacturer’s instructions. These PDMS membranes were patterned by a CO2 laser machine and the glass was patterned by an ultrasonic drilling machine (LUD-1200; Lapidary & Sonic Enterprises, Taipei, Taiwan). The substrate was made from a polystyrene plate (PS) (25 mm × 75 mm) cut from a culture dish using a CO2 laser. Finally, the patterned glass and PDMS were bonded together by a plasma treatment system (PX-250; Nordson, Westlake, OH, USA) and stuck to the PMMA adaptor with double-sided tape to completely assemble the microfluidic device. The microfluidic device, which included a cell culture chamber, a vacuum, and air bubble trap regions, was placed on top of the PS culture substrate. The function of the vacuum region was to seal the culture substrates within the microfluidic device by negative pressure. The pressure applied for sealing is about 85 mmHg. For future large-scale studies, the culture chamber can be further scaled up (up to now, its maximal culture area is 32,400 mm2, as shown in Additional file 1: Figure S1). In addition, the device was sterilized by γ-ray radiation before the experiments.
Cultivation of MSCs
MSCs were harvested from the bone marrow of postnatal 7-week-old C57BL/6 J mice (National Laboratory Animal Center, Taipei, Taiwan). Approval for the experiment was obtained from the Taipei Veterans General Hospital Institutional Animal Care and Use Committee (IACUC) regarding the use of animals prior to commencement of the experiments. For maintenance and culture expansion, MSCs were maintained in Dulbecco’s modified Eagle’s medium with 1000 mg/L glucose (LG-DMEM; Sigma-Aldrich, St. Louis, MO, USA) supplemented with 10 % fetal bovine serum (FBS; Gibco Invitrogen, Carlsbad, CA, USA), 100 units/ml penicillin, 100 μg/ml streptomycin, 2 mM l-glutamine (Gibco Invitrogen), 10 ng/ml basic fibroblast growth factor (bFGF; Sigma-Aldrich), and 10 ng/ml epidermal growth factor (EGF; R&D Systems, Minneapolis, MN, USA). Cells were seeded at a density of 3 × 103 cells/cm2 (30–40 % confluence). They were subcultured and expanded when reaching 80–90 % confluence. Confluent cells were detached with 0.1 % trypsin-EDTA (Gibco Invitrogen), rinsed twice with PBS, and centrifuged at 200 × g for 5 minutes. Cell pellets were rinsed twice with PBS and resuspended in culture medium. The cells were re-seeded at a density of 8 × 103 cells/cm2 prior to hepatic differentiation under the same culture conditions. The culture medium was replaced three times a week. All cultures were maintained at 37 °C in a humidified atmosphere containing 5 % CO2.
Proliferation and hepatic differentiation of MSCs on the microfluidic device
The procedures for proliferation and hepatic differentiation of MSCs on the culture dish and the microfluidic device are described in the supplementary material (Additional file 1: Figure S2). Hepatic differentiation was initiated using the two-step protocol we reported previously . Mouse MSCs were used for hepatic differentiation and therefore the differentiation time is about 3–4 weeks . Step-1 induction medium, consisting of Iscove’s modified Dulbecco’s medium (IMDM; Gibco BRL, Grand Island, NY, USA) supplemented with 20 ng/ml hepatocyte growth factor (HGF; R&D Systems), 10 ng/ml bFGF, 0.61 g/L nicotinamide (Sigma-Aldrich), and 100 units/ml penicillin, 100 μg/ml streptomycin, 2 mM l-glutamine, was used for induction in the first 7 days. Step-2 maturation medium, consisting of IMDM supplemented with 20 ng/ml oncostatin M (ProSpec, East Brunswick, NJ, USA), 1 μmol/L dexamethasone (Sigma-Aldrich), and 50 mg/ml insulin–transferrin–selenium (6.25 mg/ml insulin, 6.25 mg/ml transferrin, 6.25 ng/ml selenious acid, ITS+ premix; Becton Dickinson, Franklin Lakes, NJ, USA), was used for induction for 2 weeks. During the hepatic differentiation, induction medium was supplied from the syringe and injected into the chamber of the microfluidic device through the pipeline, and the outlet was connected to the waste tube. Cellular waste products were removed continuously inside the chamber. The flow rate was 100 μl/hour. For the control group, MSCs were cultured on the PS without continuous flow and were induced by the same protocol.
Functional analysis, flow field simulation, and statistical analysis
Details of the materials and methods used for RNA extraction, quantitative real-time PCR, immunofluorescent staining, flow cytometry analysis, uptake of low-density lipoprotein (LDL), urea production assay, flow field simulation and observation, and statistical analysis are described in Additional file 1.
Design, assembly, and performance test of the bio-microfluidic culture system
A five-layered microfluidic device assembled from one layer of PS, patterned PDMS, patterned glass, and PMMA adaptors was designed and fabricated (Fig. 1a). Together with the other components of the bio-microfluidic system, the device was effectively utilized for the maintenance and hepatic differentiation of MSCs (Figs. 1b and 2). Since the force generated by the flow rate of the culture medium can influence the proliferation and differentiation of MSCs, the perfusion performance in our system was initially simulated and tested (Fig. 1c, d). We aimed at maintaining a constant flow rate and homogeneous medium flow of media in the cell culture chamber. The simulated field of velocity distribution is illustrated in Fig. 1c. The simulation result showed a uniform flow profile in the culture region of the microfluidic device. In addition, the distribution of rhodamine showed that the flow field near the center line and the margin was uniform (Fig. 1d) and matched the simulation result. The Reynolds number (Re) of our microfluidic device for a 100 μl/hour flow rate was 3.6 × 10–5. The flow is considered laminar flow when Re is smaller than 2100, and turbulent flow when larger than 4000 . The 100 μl/hour flow rate in our device was therefore a laminar flow.
Efficiency of air bubble removal was further validated. The air bubble-removal structure based on the principle of negative pressure prevented experimental failures and inconsistencies of experimental data caused by air bubble interferences inside the culture chamber. Figure 1e is the schematic representation of the air bubble-removal structure. The material of the barrier wall between the microfluidic channel and the negative-pressure channel was PDMS. When air bubbles enter the microfluidic channel along with culture medium, they can be removed by the air bubble-removal structures by applying negative pressure through the gas-permeable PDMS. A movie of this process demonstrates the slow disappearance of a large air bubble injected into the microchannel of the device 12 hours previously (see Additional file 2: Video 1). The result showed that the “bubble traps” within the device were effective for removing the air bubbles. After validation of the quality of the device and performance of the perfusion system, the microfluidic device was used for the subsequent experiments of proliferation and differentiation of MSCs.
Morphology, proliferation, pH, and immunophenotype of MSCs cultured under the microfluidic device
Previous studies have shown that an optimum pH value is critical for cells to achieve high confluence and efficient enzyme activity . The pH level of the culture medium in the microfluidic device and culture dish was therefore measured during the first 3 days of culturing. The result showed that the pH value of the medium in the microfluidic device was maintained at about 7.3, but decreased from 7.4 to 6.8 in the culture dish and continued to decrease over time (Fig. 3c).
Hepatic differentiation of MSCs in the microfluidic device
In this study, a perfusion-based microfluidic system was designed to enhance the production of differentiated hepatocytes from MSCs. The design showed substantial advantage in promoting hepatogenesis of MSCs. Our design has two major advantages over other similar devices which use negative pressure to seal the open cover, such as the hybrid microfluidic–vacuum system [54, 55]. First, the use of LDW techniques reduced the fabrication time of our microfluidic system from days to hours compared with the conventional micro-electromechanical techniques . Also, the fabrication of a large culture chamber is made easy with the help of LDW techniques. Second, unlike other previous designs [54, 55], cells are not pressed against the microfluidic device at the edges of culture chamber during assembly due to a close match between the cell culture region and the culture chamber in our system; this design prevents cell damage during assembly. In addition, the microfluidic device can be reversibly bonded and unbonded to the culture substrate, which is useful for further immunocytochemistry or manipulation of cell colonies. For example, unwanted colonies can be removed easily without affecting the healthy ones during maintenance. This device may also be suitable for the differentiation of ESCs or induced pluripotent stem cells (iPSCs).
Besides a larger cell culture surface area, a geometry that allows uniform flow distribution within the culture chamber is also a featured characteristic which confers several advantages for hepatic differentiation of MSCs. The culture chamber of our microfluidic device (culture area ≥ 400 mm2) was significantly larger than that previously reported (culture area = 7 mm2) . Incorporating this device into an automated system, scale-up production for cell therapy is possible. Moreover, the flow field distribution of the culture medium inside the chamber is more uniform because of the quadrilateral geometry of the culture chamber, because the flow field of a rectangular chamber creates a more uniform flow (Fig. 1c, d). On the contrary, uneven flow occurs in a circular chamber (Additional file 1: Figure S5). Hence, a more homogeneous flow distribution within the culture chamber can be achieved in our device, which offers a favorable environment for hepatic differentiation of MSCs.
The microfluidic system, which mimics the shear flow microenvironment, theoretically confers advantages over the static culture system. It supplies a continuous flow of fresh medium and facilitates constant removal of metabolic wastes, which provides a constant microenvironment to maintain cell viability and function over a longer culture period. On the contrary, there are inherent limitations in the traditional in-vitro static culture systems, such as nutrient supply and metabolic end-product accumulation, including release of toxins that negatively influences the cell growth . The utilization of our bio-micofluidic device with a favorable environment efficiently facilitates and accelerates hepatic differentiation of MSCs. In addition, a previous study has demonstrated that the microenvironment is essential for the stemness properties of stem cells . The cellular microenvironments of MSCs in the microfluidic device and the culture dish were similar, and so was the cell viability of MSCs.
The microenvironment in the culture chamber is the key for efficient hepatic differentiation of MSCs in this study. The continuous flow not only constantly supplies nutrients for hepatic differentiation but also removes the unhealthy cells that loosely attached onto the culture chamber. As a consequence, hepatocyte-like cells with homogeneous and favorable functional activities are obtained. A previous study indicates that removal of cell-secreted factors suppresses cell growth and differentiation . However, this action did not seem to suppress cell growth and differentiation in our study. This might be due to the use of a two-step hepatic differentiation protocol which offers necessary growth factors to support hepatogenesis . These factors can be maintained at a relatively stable and higher concentration in the microfluidic system compared with the static culture. A higher expression of hepatic marker genes in the microfluidic device group may be attributable to the mechanical microenvironment as well as more homogeneous cell distribution in the device.
Shear stress plays a crucial role in regulating proliferation, differentiation, and cellular morphology of MSCs [59–61]. The shear stress generated in our microfluidic device was much lower (0.00142 Dyne/cm2) than those previously reported to regulate the proliferation and osteogenic differentiation of MSCs (0.3–2.7 Dyne/cm2) [59–61]. It did not affect the lineage commitment of MSCs because there were no significant difference between the microfluidic system and static culture in surface immunophenotype after 3 days of maintenance (Fig. 4). This means that shear stress was not a key factor for stemness properties of MSCs in our study. Moreover, a previous study indicates that low shear stress could further stimulate maturation signals through sensory systems on polarized hepatic cells . Indeed, higher expression of hepatic marker genes was observed in the microfluidic device and the mechanical signals might have contributed to functional maturation of hepatocyte-like cells.
Based on our results, MSCs in the microfluidic device after 3 weeks of hepatic induction expressed a high level of AFP and HNF4. Because AFP is specifically expressed in hepatocyte progenitors, it is therefore one of the indicators for hepatic differentiation [63, 64]. The high level of AFP after 3 weeks of hepatic differentiation means that use of the microfluidic device can accelerate the hepatic differentiation. The results indicated that both hepatocyte progenitors (AFP high expression) and matured hepatocytes (HNF4 high expression) in the microfluidic device are greater than those in the culture dish, due to the optimal differentiation conditions. Since previous studies have indicated that hepatic progenitor cells (or fetal liver cells) possess greater regeneration potential than fully matured hepatocytes in experimental liver failure models [49, 65, 66], the cells differentiated in the microfluidic device should be suitable for cell therapy compared with those generated from 2D static culture. The hepatic differentiation from MSCs in the microfluidic device is therefore a more efficient method than that in the culture dish.
An innovative microfluidic system that allows homogeneous cell seeding and real-time observation of cell morphology has been developed. The system efficiently generates hepatocyte-like cells from MSCs with more rapid functional maturation than a conventional static culture system. This microfluidic system may be further developed for large-scale production of hepatocyte-like cells to meet the high demand in the cell therapy industry. More efforts will be devoted to further scale-up and automation of this novel device.
bFGF, basic fibroblast growth factor; EGF, epidermal growth factor; ESC, embryonic stem cell; FBS, fetal bovine serum; HGF, hepatocyte growth factor; IMDM, Iscove’s modified Dulbecco’s medium; iPSC, induced pluripotent stem cell; LDL low-density lipoprotein. Re, Reynolds number; LDW, laser direct writing; LG-DMEM, Dulbecco’s modified Eagle’s medium with 1000 mg/L glucose; MSC, mesenchymal stromal cell; PDMS, polydimethylsiloxane; PMMA, polymethyl methacrylate; PS, polystyrene plate
This work was supported in part by the Novel Bioengineering and Technological Approaches to Solve Two Major Health Problems in Taiwan sponsored by the Taiwan Ministry of Science and Technology Academic Excellence Program under Grant Number MOST 105-2633-B-009-003. The authors acknowledge financial support from the Ministry of Science and Technology, Taiwan (MOST 103-2314-B-010-053-MY3, MOST 104-2321-B-010-008, and MOST 105-2911-I-010-506). This study was also supported by Aiming for the Top University Plan, a grant from the Ministry of Education.
Availability of data and materials
Additional files contain supporting data which include supplementary materials and methods, one table, five figures, and two videos.
M-HY is the first author. M-HY conceived and designed the study, performed the experiments, assembled and interpreted the data, and wrote the manuscript. Y-YW performed the experiments, assembled and interpreted the data, and wrote the manuscript. Y-SL analyzed and interpreted the data and wrote the manuscript. MR wrote the manuscript. JH-CH and OK-SL conceived and designed the study, interpreted the data, and wrote the manuscript. All authors read and approved the final manuscript.
The authors declare that they have no competing interests.
Consent for publication
The authors consent to publication of all details, images, or videos for this manuscript.
Ethics approval and consent to participate
Approval was obtained for the experiment from the Taipei Veterans General Hospital Institutional Animal Care and Use Committee (IACUC) regarding the use of animals prior to commencement of the experiments.
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