Reusable self driven microfluidic pump with large pumping capacity for POCT

doi:10.21203/rs.3.rs-3444991/v1

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Reusable self driven microfluidic pump with large pumping capacity for POCT

https://doi.org/10.21203/rs.3.rs-3444991/v1

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There is increasing demand for self-driven microfluidics to facilitate the implementation related technology for rapid point-of-care analysis. The driving flow of contemporary single modular micropumps can reach 0–9 µL·min^− 1. To improve pumping capacity, pump body has to be stacked; however, this increases the total volume and reduces the driving time. To address this problem, this study proposes a single modular micropump with the following parameters: driving flow of 0.25–22 µL·min^− 1, driving time up to 300 min, and volume ≤ 6000 mm³; three contact surface thicknesses of 0.1 mm, 0.3 mm, and 0.5 mm are designed. The pumping capacity varies with the thickness of polydimethylsiloxane on the effective contact surface. The pump pressure range is designed to meet different pump pressure requirements. Fluid experiments are demonstrated, such as mixing and bisection of microfluidic channels, and various performance parameters of the self-driven micropump are evaluated. As proof of concept, a self-driven colorimetric test for starch detection is demonstrated.

Self-driven micropump

Porous structure

Effective contact surface thickness

High pump pressure drive

With advances in point-of-care testing (POCT) in medicine, new research avenues are open for microfluidic technology. Microfluidic devices meet many strict requirements in POCT, such as portability and low test costs ^[1]. There have been considerable research advances in the field of microfluidic systems; however, controlling the flow continues to remain a challenge, especially in the field of POCT. In contemporary microfluidic systems, flow control is achieved using a traditional peristaltic pump, a syringe pump, or a pressurized flow controller; however, these usually need an additional external drive ^[2], making these methods not suitable for POCT. Self-driven microfluidic technology, on the other hand, is a new fluid control method that does not require external electronic components or moving parts, and it is robust, flexible, and portable. Examples of microfluidic systems are a finger driven pump ^[3], vacuum pump ^[4], capillary driven pump ^[5], paper driven pump ^[6] (including paper transverse flow and paper microfluidic components), and gravity driven pump ^{[7] [8]}. In some of these microfluidic technologies, it is challenging to accurately control the flow; also, they may not have large pumping capacities. Because of this, they are not suitable for application in fields that need a preset flow. In the past decade, degassing-driven modular plug-and-play PDMS micropumps and microfluidic chips have become an important part of self-driven microfluidic devices. Studies have demonstrated that silicone rubber has several advantages in “degassing driving fluid” technology. Silicone rubber is highly porous and has high permeability; it comprises a flexible silicone oxygen chain, provides free volume in its structure, and allows gas diffusion ^[9], because of which, it is widely used for manufacturing microfluidic devices.

The mechanism of the PDMS degassing drive flow depends on channel and equipment geometry, and it is sensitive to several parameters, as described by Zhao et al. ^[10] and Liang et al. ^[11]. Some studies have presented applications that have demonstrated the ability to control certain mechanisms based on this flow, such as point-of-care (POC) diagnostics ^{[12, 13]}, digital PCR ^[14], and DNA analysis ^{[15, 16]}. In 2012, Maeda et al. proposed a simple flow control method for a polydimethylsiloxane (PDMS)-based microfluidic device ^[15]. In 2012, Zhao et al. reported a modular PDMS plug-and-play micropump ^[11]. The pressure produced by these micropumps can be adjusted by controlling parameters such as the surface area and volume of the PDMS coating; however, these devices focus on moving low-volume samples. Moreover, the pumping flow per unit time of the micropump is only in the nanoliter range in a 0.5-mm Teflon tube ^{[11, 17, 18]}.

In 2016, Tanaka et al. ^[19] demonstrated an agarose micropattern method using degassed PDMS substrates and applied it in cell culture experiments. In 2015, Li et al. ^[20] established a degassed PDMS-based microfluidic droplet generator; it could generate monodispersed droplets. However, these pumping methods have short working durations and low efficiencies.

Microsystems are typically designed to analyze small sample sizes; however, some biomedical and environmental applications—that require detection of low concentrations of analytes—require processing of large volumes of fluids. For example, at least 1 mL of blood must be analyzed to detect one circulating tumor cell (CTC) in the early stage of the disease ^[21]. The same is true for the detection of nucleic acids, proteins, bacteria, or viruses in blood ^[22]. In such cases, a preconcentration step is usually required ^[23]. Contemporary self-driven microsystems can typically process a maximum of few microliters at a time, limiting their applications. In 2020, Jaione et al. proposed a modular polymer micropump that enables high-volume sample drive by using a combination of multiple modules ^[24]. However, when high pumping capacity is required in a short pumping time, several modules need to be overlapped, which makes the pump body too large.

To solve the abovementioned problems, we developed a self-driven micropump with a large driving capacity and stable pumping performance, as shown in Fig. 1. Compared with contemporary methods, the injection molding process using PDMS is highly versatile, simple, and enables rapid prototyping. Studies have shown that the pressure generated by PDMS is proportional to the effective surface area of the micropump ^[10]. Therefore, the micropump should have a higher effective area, that is, its porosity in a certain volume should be higher, to enable greater driving ability. Accordingly, in this study, we used a “sugar particle mixing method,” that is, a mixture of a PDMS crosslinking agent and brown sugar particles is cured. Then, the brown sugar is melted in a bath to achieve elasticity, porosity, and a large surface area of PDMS, achieving a high driving force in the microfluidic device. Moreover, by controlling the thickness of the silicone package layer, the pumping capacity of the pump body can be controlled, providing flexibility for use in self-driven micropumps with different specifications and pumping needs.

This study describes the performance of self-driven micropumps and demonstrates their versatility with microfluidic devices designed for some common microfluidic operations using microfluidic networks such as mixing and bisection. Finally, we evaluate the performance of these devices and perform an independent colorimetric test for starch detection.

Soft lithography has often been used to manufacture microfluidic devices, and PDMS is the most commonly used resin in soft lithography ^{[25, 26]}. To better understand the performance of the self-driven pump, the dynamics of gas transport through the PDMS thin layer were evaluated, and PDMS-related parameters were obtained, such as solubility, diffusivity, and permeability ^{[27, 28]}. Based on the PDMS dissolution rate and permeability for the design and production of self-driven micropump.

2.1 Preparation of the self-driven micropump

A self-driven micropump has two parts, namely, the internal “gluten” body and the external sealing layer, as shown in Fig. 2. The pumping power of micropumps can be changed by changing the effective surface area to volume ratio (S/V). Accordingly, we design a “gluten” porous structure with interconnected pores and various pore sizes. The volume of the micropump is fixed, and its porosity is improved, maximizing the effective area of the micropump and achieving high pump pressure.

2.1.1 Production of “gluten body”

The production of “gluten body” is mainly carried out via the “sugar particle mixing method,” as shown in Fig. 3. Dow Corning 184 silicone rubber (Dow Corning DC184PDMS, Mandarin Technology) was mixed with brown sugar granules (commercial common brown sugar) for casting. The base component was mixed thoroughly with the hardener in a ratio of 10:1, and then, brown sugar particles were mixed in (ratio of 1:2). The mixture of PDMS and brown sugar was poured into cylindrical molds manufactured using stereolithography 3D printing (Stratasys 3D Printing SLA printers). These molds were placed in a 90°C vacuum drying oven (DZF, Shanghai Lichen Bunxi Instrument Technology Co., Ltd.) for 1 h to fully crosslink the PDMS and separate the molded structure. Then, these direct cured elastomers were subjected to water bath heating, because of which, the brown sugar particles precipitated entirely. These were dried, which sealed the PDMS mixture perfusion into a “gluten”-like structure, considerably decreasing the porosity of the self-driven micropump as well as the pump pressure.

To solve this problem, the surface of the elastomer is coated with a thin layer of the PDMS crosslinking agent before heating in the water bath and then cured again at a high temperature. This ensures that the surface of the elastomer is perforated evenly before being heated in the water bath. All brown sugar particles inside the elastomer are separated and spread out to obtain “wet gluten.” The “wet gluten” is placed in a vacuum drying oven at 150°C for 3 h; this aids draining of all the water in the gluten, and the final dried “gluten” is obtained.

The as-obtained “gluten” has high hardness, good elasticity, high porosity, and a film on the outer surface, which can effectively block the infiltration of the PDMS crosslinking agent used to seal the layer.

2.1.2 Preparing the external sealing layer

The external sealing layer was cast with Dow Corning 184 silicone rubber (Dow Corning DC184PDMS, Huawan Technology), as shown in Fig. 4 (basic components and curing agent in the ratio 10:1). After degassing, the PDMS crosslinking agent was poured into the mold (SLA printer of Stratasys 3D Printing Company). The mold was placed in a vacuum drying oven at 90°C (Italian FALC) for 1 h to fully crosslink PDMS. The as-obtained structure was separated from the mold, and it has a cylindrical side sealing layer. At the bottom, PDMS films (Hefei Science material KYQN series PDMS film, thicknesses of 0.1 mm, 0.3 mm, or 0.5 mm) were used. The film was cut to an appropriate size and glued with the previously obtained PDMS cylindrical side.

2.1.3 Assembly

The cylindrical side of the “gluten body” was coated with a layer of PDMS crosslinking agent, placed in the outer sealing layer, and then, a layer of PDMS crosslinking agent was cast on the top of the “gluten body,” achieving the sealing effect. Then, this assembly was placed in a vacuum drying oven at 90°C for 1 h, obtaining the final self-driven micropump body.

2.2 Preparation of multilayer PMMA microfluidic chip

The microfluidic chip was made using “Sanming treatment,” as shown in Fig. 5. The acrylic microfluidic chip comprised a layer of PMMA board (hongyishun flagship store), a layer of silica gel film (thickness = 0.1 mm, Shentong rubber and plastic), and a layer of PMMA (hongyishun flagship store). It was clamped and sealed by a bolt structure. PMMA (thickness = 4 mm) was cut to the required size (length = 120 mm, width = 60 mm) using a CO₂ laser system (vls2.30 desktop universal laser system) and perforated for use as the equipment base. A Graphtec cutting plotter ce6000-40 (CPS cutter printer system) was used to cut microfluidic channels (width = 0.8 mm, length = 430 mm, and height = 0.8 mm) on one of the PMMA boards. A silica gel film (thickness = 0.1 mm, Shentong rubber and plastic) was used as the transition sealing layer between two PMMA plates. The final assembly comprised two PMMA boards with a silica gel film in between, and this assembly was laminated and sealed using 10 bolts. Epoxy resin glue (abida 6005 epoxy resin AB glue) was used to seal the sides.

This assembly was used to fabricate a starch detection device, with a 4-mm thick PMMA substrate. The PMMA layer had a groove for depositing samples, and gel starch in the channel. The Graphtec cutting plotter ce6000 − 40 (CPS cutter printer system) was used to cut microfluidic channels on the transparent PMMA (thickness = 4 mm) substrates. As mentioned previously, the top layer used to close the device was also a PMMA substrate, with a silica gel film (thickness = 0.1 mm) sandwiched in the middle. Then, 10 µL starch (Tianjin Zhiyuan Chemical Reagent Co., Ltd.) was placed in 30 mL water and boiled at 100°C to promote gelation. Finally, the previously described lamination method was used to seal the assembly after the sample was added. To test the colorimetric reaction, after the self-driven micropump was assembled, a pipette was used to move the iodine standard solution (C (1 / 2I2) = 0.01002 mol/L, Yida Technology Co., Ltd.) to the inlet of the microfluidic device.

2.3 Driving and characterization of the micropump

The as-obtained self-driven was degassed in a vacuum drying oven (DZF type, Shanghai lichen Bangxi Instrument Technology Co., Ltd.) at 0.81 MPa for 8 h and vacuum packaged (sv-204 vacuum sealing machine, sammic, Spain) for storage in an airless environment. This vacuum packaging process does not use a plastic box and thus allows long-term storage of the pump, because of which, it is also easily available.

The micropump was pasted on the multilayer PMMA microfluidic device via a PSA patch; the flow curve caused by the micropump over time was measured using a flow characterization device. The flow characterization device was placed on the stage of an inverted microscope (ix-51, Olympus, Japan) equipped with a CCD camera (DP70, Olympus, Japan), the food dye solution (1%, w / V, Shanghai Jiahui Fine Chemical Co., Ltd.) was dripped at the entrance, and then, the CCD camera was used to capture the image at an interval of 1 min until the movement in front of the liquid stopped. The pumping pressure was evaluated by monitoring the length of the liquid column in the microchannel.

In addition to the pumping flow, we also studied the sensing time of the flow driven by the PDMS pump. A drop of the food dye solution (1%, w / V) was added to the inlet of the microchannel through a micropipette. When the inlet was blocked, the CCD camera was used to capture images at an interval of 1 s.

3.1 Performance of the self-driven micropump

The degassed PDMS micropump could absorb air from the closed system into the PDMS. This absorption process produced negative pressure and triggered movement of the fluid in the microchannel. Zhao et al. reported a PDMS micropump ^[17] in which the pressure generated by its PDMS was directly proportional to the effective surface area of the micropump. However, the flow rate per unit time generated by these micropumps in a Teflon tube (inner diameter = 0.5 mm) was only in the nanoliter range. Jaione et al. proposed a PDMS modular micropump (volume = 4000 mm³) with a driving flow of up to 8 µL·min^− 1.

We designed a self-driven PDMS micropump with a high driving flow, small volume, and large surface area. The volume of the designed micropump is < 6000 mm³, and it can produce a driving flow of 0.25–22 µL·min^− 1, with a driving time of up to 300 min. The PDMS thickness of the effective contact surface (the surface connected with the microfluidic chip) affects the driving flow of the self-driven micropump. Three effective contact surface thicknesses were designed (0.1 mm, 0.3 mm, and 0.5 mm), and the corresponding driving flows were observed, as shown in Fig. 6, to satisfy needs of different applications.

When we measured the drive flow generated by each micropump, a rapid increase in its maximum was observed within the first few minutes of operation. The time required to reach the maximum drive varied from 30 s to 2 min; thinner effective contact surface led to greater drive flow and longer drive time. Thereafter, the flow rate decayed in a double exponential waveform, and then entered a slow decay state, gradually reaching equilibrium. The flow velocity distribution at each time point was observed and recorded, as shown in Fig. 2. In the first 10 min, the driving capacity of each micropump decreased rapidly. After 20 min, the driving capacity approached a linear trend and then declined slowly and steadily.

The pumping flow ranges of the self-driven micropump with effective contact surface thicknesses of 0.5 mm, 0.3 mm, and 0.1 mm are 6.3–0.75 µL·min^− 1, 9.3–0.5 µL·min^− 1, and 22–0.5 µL·min^− 1, and the corresponding pumping times were 150 min, 200 min, and 320 min, respectively. The pumping performance was considerably better than that of the PDMS modular micropump proposed by Jaione et al. Therefore, we determined that thinner effective contact surface of the self-driven micropump results in larger pumping capacity and longer driving time. Based on these findings, custom self-driven micropumps with different effective contact surface thicknesses can be designed for different needs to achieve effective flow control.

The permeability of the micropump material to air is one of the key parameters for determining the pressure and fluid flow rate generated by the micropump in the microchannel ^[29]. The average flow velocity of micropumps with different effective contact surface thicknesses was observed. The active times of these self-driven micropumps were inconsistent because of different effective contact surface thicknesses; therefore, we observed the average flow velocity in the first 120 min, as shown in Fig. 6D.

Next, we evaluated the volume of liquid displaced by the micropump. Our self-driven micropumps with effective thicknesses of 0.5 mm, 0.3 mm, and 0.1 mm replaced approximately 275 µL, 550.4 µL, and 1100.8 µL, respectively. The amount of liquid that can be displaced by the micropump is determined by its flow rate and activity time. Our findings indicate that these micropumps can be used to pump a wide range of fluids, down to the milliliter range, by varying the effective contact surface thickness, which is a major breakthrough in self-driven microfluidics.

To further explore the stable pumping capacity of the self-driven micropump, two and three pigments were transported into the microchannel, and the formation and maintenance time of laminar flow were observed using a microscope, as shown in Fig. 7a and b. The laminar flow formation process occurred in < 70 s, demonstrating the considerable driving effect of the self-driven micropump, which further proves that the self-driven micropump has application potential in complex microfluidic devices. The laminar flow state was maintained for 160 min. Compared with the PDMS micropump with no porous structure, the PDMS sponge vacuum pump can maintain the laminar flow for a long time, which is mainly attributed to the porous structure of the self-driven micropump “gluten.” It provides more space for vacuum storage and for absorbing liquid samples.

Between the time the PDMS pump is installed on the PDMS microfluidic unit and the time when the flow starts, there is a small delay, which we define as flow induction time; it is a key parameter that influences fluid manipulation in a microfluidic system. In a microfluidic device, channel geometry, surface characteristics, and flow sensing time depend on the performance of the installed PDMS pump. ^[17] Therefore, we also examined the induction time of the flow driven by the PDMS pump (Fig. 8). The induction time of pumping with effective contact surface thicknesses of 0.1 mm, 0.3 mm, and 0.5 mm was 1.5 s, 2.3 s, and 3 s, respectively.

3.2 Independent colorimetric test for starch detection

To demonstrate a self-driven colorimetric test, we fabricated an automatic device for starch detection via the reaction of an iodine solution with starch by combining the self-driven micropump with the microfluidic chip. First, the self-driven micropump was connected to the multilayer microfluidic device. The microchannel configuration was a single channel (microchannel cross sectional size: 0.8 mm × 0.8 mm) with a circular groove (diameter = 5 mm, depth = 6 mm), and in this, a transparent starch gel (Tianjin Zhiyuan Chemical Reagents Co., Ltd.) was placed. During starch testing, when the iodine standard solution (C (1/2I2) = 0.01002 mol/L, Eida Technology Co., Ltd.) reached the groove and came in contact with the starch gel, the color changed to “blue-black,” as shown in Fig. 9. This reaction is widely known as an iodine test, and it is commonly used for starch detection in food samples ^[30]. A self-driven micropump with an effective contact surface thickness of 0.1 mm of PDMS was installed on the interface of the microfluidic chip (diameter = 20 mm, depth = 6 mm), and 100 µL of the iodine standard solution was loaded into the inlet. After 3 min, the iodine solution had completely filled the microchannels, and the starch gel had completely turned blue-black, as shown in Fig. 9a. Finally, the iodine solution flowed through the microchannel to the reservoir. The whole experiment lasted for 5 min. Despite the high flow resistance and large size of the device, the ability of the microfluidic device to operate autonomously demonstrates the possibility of using self-driven micropumps and microfluidic chips to fabricate functionally integrated self-driven analysis microdevices.

Rapid clinical analysis is one of the most popular applications of microfluidic technology, and it is mainly implemented using self-driven microfluidic technology. This study presents a highly versatile self-driven microfluidic system that can pump large flows of liquid. The “brown sugar crystal mixture method” is adopted, which renders the PDMS self-driven micropump high porosity, increases the effective surface area to volume ratio (S/V), and considerably improves its driving ability. The driving force of the self-driven micropump could be varied by changing the PDMS thickness of the effective contact surface, enabling the design of the desired pumping capacity and flow control without changing the volume of the pump. We aim to investigate this in future research, and we are sure it will provide significant insights for applications. A thinner effective contact surface results in greater pumping capacity and longer driving time; the driving time of the self-driven micropump reached 180–300 min, which is a considerable increase.

Combining self-driven micropumps with multilayer PMMA microfluidic chips is a popular strategy for the development of self-driven microfluidic devices because of the following advantages: 1. It allows rapid change of filter element design, and 2. it enables reuse. Finally, we demonstrated an autonomous colorimetric test for starch detection. The possible use of self-actuated microdevices in diagnosing large biological samples was demonstrated, as well as the transport of fluids with different characteristics. The presented system is easily fabricated and enables configuration of different microfluidic channels for various functions, because of which, it is suitable for a wide range of applications, such as sample loading, mixing, separation, and cell culture. The simple combination of self-driven micropumps and multilayer PMMA microfluidic chips can overcome several limitations of contemporary self-driven devices, and it has the advantages of miniaturization and integration, which demonstrates its potential in applications that require precise flow control.

Conflict of Interest

The authors declare no conflict of interest.

Funding

This research received no external funding.

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No competing interests reported.

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Reusable self driven microfluidic pump with large pumping capacity for POCT

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Abstract

Figures

1 Introduction

2. Materials and Methods

2.1 Preparation of the self-driven micropump

2.1.1 Production of “gluten body”

2.1.2 Preparing the external sealing layer

2.1.3 Assembly

2.2 Preparation of multilayer PMMA microfluidic chip

2.3 Driving and characterization of the micropump

3. Result and Discussion

3.1 Performance of the self-driven micropump

3.2 Independent colorimetric test for starch detection

4. Conclusion

Declarations

Conflict of Interest

Funding

References

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