Working principle of micromixer module
Two mixing strategies were applied in the proposed micromixer: Dean force induced by the helical 3D channel structure and the mismatch of flow rates induced by twisted helical groove structures following the 3D spiral. The first strategy creates fluid velocity mismatching in the channel’s inner side and outer side by having the curved channel, leading to the formation of two opposing vortexes in the channel and thus reducing the diffusion distance of the two fluids (Chai et al. 2021; Cai et al. 2017). For the second strategy, the twisted helical groove structure contributes to fluid mixing by creating a slow fluid flow zone and therefore inducing another mismatching of fluid velocity. This fluid mismatching carries the fluid from one side towards the other side of the channel, increasing the chance of fluids contact (Vasilescu et al. 2020; Chai et al. 2021); consequently, the increased contact of different fluids enhances the molecular diffusion. As previously reported, the groove designs in the channel would not introduce strong secondary flow (Tsui et al. 2008). Additional file 1: Fig. S3 shows the simulation results of the micromixer. Increasing the inlet fluid flow ratio leads to increased pressure in the system, which is negligible for smaller flow rate ratios and shows the system can be powered by normal lab-scale pumps. The cross-section 1 (CS1) across different flow rates in Additional file 1: Fig S3 shows that the chaotic advection phenomena dominate over diffusion when the flow rate increases. However, higher flow rate ratios do not necessitate a higher mixing index since fluids take time to mix and diffuse (Additional file 1: Fig. S3). Interestingly, velocity distribution for lower flow rates shows a symmetric profile along the channel length (Additional file 1: Fig. S3C), while it becomes asymmetric for higher flow rate ratios. This phenomenon might also contribute to the reduction of the mixing index at higher flow rates.
The experimental results of the mixing index with pure water and food dye for various flow rate ratios are illustrated in Fig. 2A. The mixing efficiency of the device was higher than 95% (Additional file 1: Fig. S5) at the flow rate ratio of 1 mL/min:2 mL/min. Hence, the total flow rate of 3 mL/min was chosen as an optimised flow rate for cell harvesting. Based on the method described in Additional file 1: Section S1, the experimental mixing index is 82.7%. The discrepancy between simulation and experimental results can be attributed to the difficulties of imaging 3D printed channels with microscopy and the addition of extra noise in the picture due to the unsmooth surface of the micromixer (Rouhi et al. 2021). The micromixers have no splitting, obstacles, or sharp turning, which are appropriate for processing cells without damaging them.
Working principle of the microseparator module
The focusing position of microparticles inside a curved microfluidic channel is affected by two forces, inertial lift force (\({F}_{\mathrm{L}}\)) and Dean drag force (\({F}_{\mathrm{D}})\) (Amini et al. 2014):
$${F}_{\mathrm{L}}=\rho \left(\frac{{U}_{\mathrm{max}}}{{D}_{\mathrm{h}}}\right){C}_{\mathrm{L}}{a}^{4},$$
(1)
$${F}_{\mathrm{D}}=5.4\times {10}^{-4}\pi \mu {De}^{1.63}a.$$
(2)
\({F}_{\mathrm{L}}\) is affected by the density of fluid \(\rho\), the hydraulic diameter \({D}_{\mathrm{h}}\) (which can be calculated by \(4A/P\), \(A=\) channel cross-section and \(P=\) perimeter of the channel), the maximum fluid velocity \({U}_{\mathrm{max}}\) which is approximated as \(2\times {U}_{\mathrm{f}}\) (\({U}_{\mathrm{f}}\) is the average velocity), \({C}_{\mathrm{L}}\) which is a constant named dimensionless lift coefficient number and is dependent on the channel Reynolds number \((\mathrm{Re}=\rho {U}_{\mathrm{f}}{D}_{\mathrm{h}}/\mu , \mu\) is the viscosity of the liquid) and the diameter of particles \(a\). \({F}_{\mathrm{L}}\) consists of two forces: shear-gradient and wall-induced lift force. Shear gradient lift force pushes the particles towards the wall due to the velocity difference between the middle area and the side area of the channel. When the particles move close to the wall, the wall lift force pushes the particles away. The balancing point of inertial equilibrium position contributed to the lift force is where these two forces balance each other (Razavi Bazaz et al. 2020c).
In a curved channel, the channel’s curvature causes the inner wall (IW) fluid to flow faster than the outer wall (OW) due to the shorter distance travelled. This transverse fluid flow creates another force that affects the focusing position of the particles, which is the Dean drag force (\({F}_{\mathrm{D}}\)). \({F}_{\mathrm{D}}\) is defined in Eq. (2), where \(\mathrm{De}=\mathrm{Re}\sqrt{{D}_{\mathrm{h}}/2R}\) is the Dean number, and R is the radius of curvature; it describes the strength of \({F}_{\mathrm{D}}\). According to Eqs. (1) and (2), the forces applied to the particles are proportional to the particle size (\({F}_{\mathrm{L}}\propto {a}^{4}, {F}_{\mathrm{D}}\propto a\)). Therefore, different particle sizes have different focusing positions across the channel cross-section, and they can be collected through separate outlets (Mihandoust et al. 2020; Ozbey et al. 2019).
In a normal spiral channel, the particles inside the channel need to follow the rules of \(\mathrm{Cr}>0.07\), where \(\mathrm{Cr}=a/{D}_{\mathrm{h}}\) to be affected by the inertial forces inside the channel. In a scaled-up microfluidic channel, the increase in \({D}_{\mathrm{h}}\) results in a reduction of absolute flow velocity compared with a normal microfluidic channel. Therefore, the secondary forces applied to the microparticles were weaker, and the \(\mathrm{Cr}\) value in the scaled-up microfluidic channel was much higher than the microfluidic channels \((\mathrm{Cr}>0.17)\) (Moloudi et al. 2019; Carlo 2009).
Another factor that affects particle focusing is channel rigidness. There is no swelling or channel inflation in rigid channels compared to traditional PDMS chips; thus, the scaled-up device should have theoretically a lower \(\mathrm{Cr}\). Also, larger particles are more likely to be affected by mass and gravity since they are not neutrally buoyant (Moloudi et al. 2019), adding another variable despite flow velocity; the variable sizes of particles would also increase the difficulty in the channel design. When MCs and cells pass through the channels, focusing MCs near the IW causes the MSCs to be dispersed in the channel due to the large size difference between MCs and cells (MCs size are 150–220 µm, and MSCs are 15–20 µm). However, since large particles occupy the inner channel, the particle–particle interaction can stop some of the MSCs from going out through the inner outlet (Moloudi et al. 2018). Considering all these factors, in this study, we have designed the channel with a trapezoidal cross-section and heights of 550 µm and 620 µm, and a width of 1100 µm. This spiral chip has 6 loops and a slightly slanted enlarged inlet size to prevent clogging of MCs at the beginning of the channel (Fig. 1).
Working principle of the microconcentrator module
The zig-zag channel relies on inertial, and Dean drag forces to focus the MSCs at the centre of the channel. When Reynolds number of the channel falls in the intermediate range 1 < Re < 100, the fluid flow is laminar, between Stokes and turbulent flow regimes. Therefore, inertial forces focus the randomly dispersed particles toward certain equilibrium positions after a sufficiently long channel length. As explained above, shear-gradient and wall-induced lift force are the main forces affecting the particle focusing in straight channels, and they both contribute to the overall inertial lift force \({F}_{\mathrm{L}}\). Straight channel relies on the difference in particle sizes to focus the particles at different positions (\({F}_{\mathrm{L}}\propto {a}^{4}\)). In zig-zag channels, Dean force \({F}_{\mathrm{D}}\) is introduced differently compared to the spiral microfluidic channel. The interchanging channel direction creates a mismatch of fluid flow velocity in an alternating pattern and introduces Dean force, accelerating the focusing of particles inside the channel.
A zig-zag channel has three focusing modes across different flow rates. When \({F}_{\mathrm{L}}{<F}_{\mathrm{D}}\), the particles focus at the side of the channels. When \({F}_{\mathrm{L}}{>F}_{\mathrm{D}}\), the particles were focused in the middle of the channel due to due to the strong \({F}_{\mathrm{L}}\). When \({F}_{\mathrm{L}}{\sim F}_{\mathrm{D}}\), particles are in the transition mode. For the aim of this study, MSCs need to satisfy the condition of \({F}_{\mathrm{L}}{>F}_{\mathrm{D}}.\) One primary advantage of the zig-zag channel is its operating ranges of flow rates, i.e., it can focus particles at the centre over a wide range of flow rates. After careful evaluations, the zig-zag channel with a cross-section of 360 µm × 60 µm, 60° angle has been proposed to concentrate cells after the spiral microfluidic device. To avoid clogging of zig-zag channels caused by the remaining MCs in the target outlet, some obstacles were planted at the target outlet of the spiral to ensure no MCs could enter the zig-zag channel.
Pressure balance of microfluidic system
Combining multiple microfluidic devices in one system requires careful arrangement to balance the fluid flow and pressure change. An electronic circuit was used as an analogy for our system to understand better the fluid behaviour in the system (Additional file 1: Fig. S5). These microfluidic devices resemble the resistors that reduce the pressure input from the pumps, similar to the voltage drop in an electronic circuit (Oh et al. 2012). Keeping the flow rate and pressure stable according to the following equation is the key point of the successful operation of this system:
$$Q=\frac{\pi {{R}_{\mathrm{H}}}^{4}}{8\mu }\frac{\Delta p}{L},$$
(3)
where \(Q\) is the volumetric flow rate, \({R}_{\mathrm{H}}\) is the hydraulic resistance of the channel, \(\mu\) is the viscosity, \(\Delta p\) is the pressure drop, and \(L\) is the channel length. In a serial circuit, \(Q\) (which is current I in the electronic circuit) remains constant in each device, thus \({Q}_{\mathrm{spiral}}={Q}_{\mathrm{mixer}1}={Q}_{\mathrm{mixer}2}\). \({Q}_{\mathrm{mixer}1}\) has two inputs, one from the peristaltic pump, and one from the syringe pump. In a parallel circuit, the current of the circuit \({Q}_{\mathrm{mixer}}={Q}_{\mathrm{inlet}1}+{Q}_{\mathrm{inlet}2}\). The working flow rates of micromixers and zig-zag channels are more flexible, while the spiral microfluidic device only works under a specific flow rate. To achieve this flow rate, we change the flow rate of the two pumps according to \({Q}_{\mathrm{spiral}}={Q}_{\mathrm{inlet}1}+{Q}_{\mathrm{inlet}2}\). The outlet’s resistance of the spirals affects the focusing of the MCs in the inner outlet. Therefore, the fluid pressure of the zig-zag channel must be balanced with the pressure-damping channel connecting to the inner outlet of the spiral device. This pressure-damping channel needs to have the same hydraulic resistance \({R}_{\mathrm{H}}\) to the zig-zag channel, which can be calculated by Eq. (4) (Oh et al. 2012):
$${R}_{\mathrm{H}}=\frac{8\eta L}{\pi {D}_{\mathrm{h}}},$$
(4)
where \(\eta\) is the viscosity and \(L\) is the finite length of the channel. Since \({D}_{\mathrm{h}}\) of the channel is fixed and \({R}_{\mathrm{H}}\propto 8L\), changing the length of the pressure-damping channel to reach R3 = R4 balances the pressure of the system and would not affect the particle focusing positions in the spiral channel (Additional file 1: Fig. S6). This system potentially eliminates the debris larger than cells through spiral channel, and removes debris smaller than the cells through the zig-zag channel.
Evaluation of different modules with fluorescent microbeads and microcarriers
The maximum capacity and optimal flow rate of the spiral microfluidic device was determined by passing a different concentration of MCs through the device across a range of flow rate. As shown in Fig. 2B and Additional file 1: Fig. S6, from 2.0 to 4.0 mL/min, the focusing position of the MCs gradually shifts to the outer outlet. Noticeably, 3.0 mL/min is the critical flow rate that runs under high throughput while still focusing the MCs at the inner outlet. MCs with a concentration higher than 1% escape from the outer outlet even at a lower flow rate. However, MCs with a concentration of 0.75% can be sufficiently removed from the inner outlet at a flow rate of 3 mL/min. At the flow rate of 3 mL/min (2 mL/min from the bioreactor, 1 mL/min from the enzyme reservoir), the fluid mixing efficiency reached 95% after the first micromixer (Additional file 1: Fig. S4). The addition of the enzyme from the syringe pump inlet of the micromixer dilutes the sample.
The microcarrier concentration used for cell culture was 1.29% v/v% (1 g in 80 mL media). Therefore, MCs’ volume and concentration for cell harvesting before entering the microfluidic gadget were set to 70 mL to reach 0.75% when the sample arrived at the spiral microfluidic chip. The volume was calculated by the following equations: target concentration (0.75%)/dilution factor in micromixer (2/3)/concentration in culture (1.29%) × volume in culture (80 mL). As such, 40 mL of TrypLE was added since there was 30 mL of media inside the bioreactor after 50 mL of supernatant was taken away. The flow rate was set at 2 mL/min from the bioreactor and 1 mL/min TrypLE from the syringe pump, so the total flow rate of 3 mL/min fluid proceeded into the spiral. To demonstrate the inertial forces in the system do not damage the MCs, we passed MCs through the two micromixers and one spiral chip setup under a 20 mL/min flow rate. The results showed that the gentle forces applied by the micromixer do not change the shape and size of the MCs (Fig. 2C). Various inertial microfluidic channel designs can be used in this application as evidenced in our previous publications (Moloudi et al. 2018). In this study, we have showcased a rigid channel in the processing of large particle through the power of 3D printed inertial microfluidics.
The zig-zag channel was responsible for further concentrating the harvested cells. Since it was connected to the outer outlet of the spiral, the operation flow rate of the zig-zag channel needed to match the flow rate of the outer outlet of the spiral. The zig-zag concentrator was tested with 15 and 20 µm beads across different flow rates. The results showed that from 1.6 to 1.9 mL/min, the beads were concentrated 100% in the middle outlet (Fig. 2D). The beads were concentrated ~ 3.5 times, with ~ 70% of the volume removed, indicating good dewatering efficiency of the device.
Application showcase
Harvesting MSCs from bioreactor using the microfluidic system
To investigate the efficiency of the microfluidic gadgets on cell detachment, the cells were stained with Hoechst before passing through the mixer. To ensure the complete detachment of cells in the micromixers, a one-inlet micromixer was added at the end to increase the interaction of cells and enzyme under the same mixing efficiency (Vasilescu et al. 2020). Figure 3A shows microcarrier-cell suspension before cell harvesting in which cells covered the whole surface of MCs. The growth of healthy MSCs on MCs commonly leads to cell–MCs aggregation (Ferrari et al. 2012) (Additional file 1: Fig S7). Therefore, to prevent the blockage of microfluidic devices, the cells–MCs suspension was incubated with enzyme for 5 min in the incubator to detach these aggregates. Figure 3B shows the MSCs were detached from MCs’ surface by enzymatic treatment and gentle mechanical force after passing through the micromixers.
The media containing detached cells and MCs from the micromixers were then passed through the spiral. Later, they were collected separately from two outlets (Fig. 3C). 94.11% of MCs were successfully removed in the first round of separation. 76.62 ± 2.1% and 17.21 ± 0.6% cells were recovered from the OW outlet in the first and second pass, respectively, and 6.16 ± 1.80% cell loss through the IW outlet at the end of the process (Fig. 3D, Additional file 1: Fig. S8). The sum of yield (sum of cells harvested from the OW outlet over the total cell harvest from all outlets) can reach ~ 94%. Adding some obstacles at the outlet leads to 100% of the microcarrier removal rate, making it ready for clinical applications. Additional file 1: Fig. S9 shows the tight focusing band of MSCs in the middle outlet and the removal of small debris in the outer outlets. The cell solutions were collected from the outer outlets, and no cell was found in the waste outlet. Cells were concentrated 4.5 times compared to the pre-filtered samples. Although the counting results showed that the recovery rate was higher than 100%, a small number of cell loss could potentially happen due to the heterogeneity, clumping of cells, or attachment to the tubing or channel walls.
MSCs viability and proliferation after microfluidic cell harvesting
Cell viability was assessed immediately after harvesting. The live and dead staining results indicate that the microfluidic device did not compromise the viability of cells (Fig. 4A). MTS (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium) assay illustrating the metabolic activity of cells harvested by the device is also similar to the control. In the microfluidic group, the absorbance of media at 490 nm wavelength increased over time which indicates that cells have slightly higher metabolic activity than the control group, although the difference is not significant (Fig. 4B). Cell attachment, morphology, and proliferation were evaluated by staining the post-harvesting cells using DAPI and phalloidin. The fluorescent microscopy images in Fig. 4C and D indicate cells harvested with the microfluidic device have comparatively better cell attachment (Additional file 1: Fig. S10) than the control group on the first day of culture. After 3–5 days of culture, both groups of cells were confluent in the wells, and no significant difference in the growth rate was observed. Additionally, cells maintained their spindle morphology after harvesting with the device, and the size of cells was around 13–17 µm in both groups. The number of harvested cells after 1, 3, 5 days of cell seeding was counted by ImageJ to verify the MTS results. The results confirm that the microfluidic system does not affect cell attachment and growth after harvesting (Additional file 1: Fig. S10).
Stem cell properties and therapeutic properties of the harvested MSCs
To confirm the stemness and multipotency of the harvested cells, the MSC surface markers were evaluated and trilineage differentiation was performed. CD90, CD73, and CD105 were stained with fluorescent antibodies (ThermoFisher, Australia) staining and counted by a flow cytometer (CytoFLEX LX, Beckman Coulter, USA). Figure 5A shows 98%, 100%, and 100% of the cells express CD90, CD73, and CD105, respectively, confirming the well-preserved MSCs identity. To assess the multipotency of cells after harvesting, cells were stained with Oil Red, Alizarin Red, and Alcian Blue staining after treating with adipogenic and osteogenic/chondrogenic induction media, respectively (Fig. 5B). Formation of bright red stain calcium deposits stained by Alizarin Red S confirmed osteoblastic phenotype of cells. Additionally, presence of red lipid droplets stained by Oil Red O verified the adipocyte phenotype, and the blue glycosaminoglycan complex staining showed the presents of chondrogenic cells. These results indicate that cells retained their differentiation potential.
The therapeutic effect of harvested MSCs is verified by staining the surface therapeutic proteins and analysis of the cytokines in the cultured supernatant. Figure 5C shows the changes in the expression level of the surface therapeutic proteins after priming with TNF-α and IFN-Υ for 24 h. HLA-G is a protein that prohibits the growth of lymphocytes, which expression level does not change with priming (Nasef et al. 2007; Selmani et al. 2009; Najar et al. 2012). The expression level of HLA-G in both microfluidic and control groups remained constant after priming. CD54 (iCAM) is a T-cell activation-related protein that is sensitive to inflammation, and the expression level of this protein increased significantly after priming (Rubtsov et al. 2017; Tang et al. 2018). Figure 5C shows that the expression level of CD54 increased 100% after priming in both groups. These results prove that there is no significant difference in the therapeutic properties of the MSCs after proceeding through the microfluidic device. Next, using the Custom ProcartaPlex Multiplex immunoassay panel, we analysed the secretion profile of the harvested cells compared to the secretion profile of cells passaged stably in multilayer cell factories. The results showed that the harvested cells expressed a similar or lower level of HGF, IL-6, CCL2, VEGF-A, and TNF-RI compared to the passage 4, passage 8 multilayer cell factory grown controls, the expression level of SDF-1 alpha and TIMP-1 are much higher than the control group (Fig. 5D).
Multiplexing the microfluidic harvesting system for large-scale application
A multiplexed system was built with the same printing protocols to demonstrate the capability of scaling up the microfluidic system for large-scale applications. The system consists of five layers (Fig. 6); the first layer is the fluid splitting layer; it has one inlet for cell and microcarrier solutions to enter the system and another inlet for the digestive enzyme with a flow rate of 8 mL/min for the cell and microcarrier solution and 4 mL/min for the enzyme. These two inlets split the total flow into four even sets and enter the 4 micromixers evenly in the second layer. The micromixers have inserted holes for the pins to anchor the positions and prevent leakage. The flow rate in each micromixer is 3 mL/min for detaching cells from MCs. The third layer is the spiral layer, with a pin inserted into the outlet of the micromixers. The solutions collected from each of the two micromixers were evenly split into two spiral microfluidic devices, and each spiral received 3 mL/min liquid flow to separate cells from MCs. Then, the fourth layer, a splitting layer was used as the bottom layer of the spiral. Two holes were opened at the outlets of the spirals, and this layer was bonded with the fifth layer spiral layer with double adhesive tape. Lastly, a whole 3D printed layer with 4 zig-zag channels and pressure-damping channels was attached to the splitting layer with double adhesive tape. The inner outlets of each spiral are connected to one pressure-damping channel, and the outer outlets of each spiral are connected to one zig-zag channel. The cross-sectional area ratio of the inner and outer outlet is 2:3; the flow rate of the outer outlet is, therefore 1.8 mL/min for each spiral. As shown in Fig. 2, the zig-zag channel can focus the cells from 1.4–1.9 mL/min. This flexible working range of the zig-zag channel ensures the cells focus on the middle outlet of the device and reduce the requirement of precision of the pressure-damping channel. The total flow rate of the cell outlet was 7.2 mL/min, while the MCs outlet was 4.8 mL/min.