Syringe-assisted point-of-care micropumping utilizing the gas permeability of polydimethylsiloxane
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- Xu, L., Lee, H. & Oh, K.W. Microfluid Nanofluid (2014) 17: 745. doi:10.1007/s10404-014-1356-4
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By utilizing the high gas permeability of polydimethylsiloxane (PDMS), a simple syringe-assisted pumping method was introduced. A dead-end microfluidic channel was partially surrounded by an embedded microchamber, with a thin PDMS wall isolating the dead-end channel and the embedded microchamber. A syringe was connected with the microchamber port by a short tube, and the syringe plunger was manually pulled out to generate low pressure inside the microchamber. When sample liquid was loaded in the inlet port, air trapped in the dead-end channel would diffuse into the surrounding microchamber through the PDMS wall, creating an instantaneous pumping of the liquid inside the dead-end channel. By only pulling the syringe manually, a constant low flow with a rate ranging from 0.089 to 4 nl/s was realized as functions of two key parameters: the PDMS wall thickness and the overlap area between the dead-end channel and the surrounded microchamber. This method enabled point-of-care pumping without pre-evacuating the PDMS devices in a bulky vacuum chamber.
KeywordsPoint-of-care Polydimethylsiloxane (PDMS) Pump
One of the challenging unit operations in point-of-care testing (POCT) diagnostics is autonomous, controllable, and on-demand pumping. Obviously, manual injection of sample liquids with a syringe or a pipette is one of the simplest methods. However, such direct injection method is not suitable for use where constant and controllable pumping is required over the injection time. Currently, most of the pumping methods in POCT are employed passively, mainly based on capillary force (Gervais and Delamarche 2009). Since no external energy is required, the capillary-driven pumping scheme is attractive for many POCT systems. Also based on the capillary effect, alternative ways such as paper-based and textile-based pumping are studied (Nilghaz et al. 2012; Yetisen et al. 2013). However, this pumping method mainly relies on the wetting properties of different testing liquids, which are unstable. Thus, it is difficult to generate stable and steady pumping (Ziegler et al. 2008). Another interesting point-of-care micropumping method is realized by utilizing the elastic deformation of PDMS (Weibel et al. 2007; Li et al. 2012). First, the microfluidic device made of PDMS is deformed by using either thumb or screw. Then, once the external force is removed, the PDMS device will go back to its initial form, which will withdraw liquids flowing inside the channel. In general, the flow rate generated by this pumping method is not linear or constant. In order to keep the liquids flowing inside the channel, the external force has to be repeatedly applied and removed.
Utilizing the high gas permeability of PDMS to achieve pumping has been considered as a promising way. High gas permeability is one of unique properties of PDMS (Merkel et al. 2000). This has allowed easy removal of trapped air bubbles out of microchannels (Ong et al. 2007; Sung et al. 2010). In addition, vacuum-assisted self-powered pumping has been introduced by Hosokawa et al. (2004). Thus, by loading liquid at the inlet port of a dead-end channel in the degassed device, the fluid can be drawn into the channel. However, PDMS devices need to be placed in a vacuum chamber to be pre-vacuumed. This is due to the reabsorption of air inside the vacuumed PDMS devices after air exposure to reach the new equilibrium under the atmosphere. This method was used to perform a simple immunoassay by adding two different types of liquids (Hosokawa et al. 2006). Based on the same principle, a self-powered microfluidic blood analysis system (Dimov et al. 2011; Liang et al. 2011) and a viscometer (Tang and Zheng 2011) have been successfully realized.
Although the vacuum-assisted degas-driven flow is a convenient way to adopt without external pumps, there are still several practical limitations: (1) Before a device can work as a pump, it needs to be stored in a vacuum chamber for more than 30 min to degas the PDMS bulk or in a sealed shrink-wrap vacuum packaging; (2) after the device is exposed to air, its pumping ability decays immediately with time, so the device should be used immediately after air exposure; (3) since pumping ability decays with time, the flow rate generated from the vacuum-assisted flow is nonlinear and uncontrollable; alternatively, instead of pre-evacuating the entire device, a low-pressure source was connected to the chamber and separated with a dead-end channel by a thin PDMS membrane (Mark and Bruce 2006). By controlling the thickness of the membrane, different flow rates were realized. However, the process to make this kind of multilayer sandwich structures is laborious due to precise alignment between the top and bottom layers and careful control of the PDMS membrane thickness and uniformity. Instead of using sandwich structures, a dead-end microchannel surrounded by a vacuumed microchamber in the same layer was employed for PCR test (Trung et al. 2010). In this approach, a constant vacuum was applied by external source and the flow rate is controlled only by varying the PDMS wall thickness between the dead-end channel and surrounded microchamber, which gives a limited range of loading flow rate (0.3–0.7 nl/s).
To overcome these limitations of using the vacuum-assisted pumping, we have designed a simple point-of-care pumping method using a single-layer structure that does not require pre-evacuation of PDMS devices in a vacuum chamber. A dead-end microfluidic channel is partly surrounded by an embedded microchamber, with a thin PDMS wall separating the dead-end channel and the embedded microchamber. A syringe connected with the microchamber port is employed to provide a negative pressure source. Low pressure generated inside the microchamber by pulling out the syringe plunger will draw air trapped in the dead-end channel through the thin PDMS wall, creating an instantaneous pumping of the liquid inside the dead-end channel. The flow rate can be regulated by controlling the wall thickness and the overlap area between the dead-end channel and the surrounded microchamber. Employing only a hand-held syringe, this method allows for on-demand pumping without pre-evacuating the PDMS devices in a bulky vacuum chamber. In this paper, we have systematically investigated the major parameters (e.g., wall thickness, overlap area) that can generate constant and controllable flow rates.
A rough estimate of the characteristic time to allow for the constant and steady-state air flux across the thin PDMW wall can be obtained by examining the diffusion time (tD) across the PDMS wall: tD ≈ w2D−1, where w is the thickness of the PDMS wall and D is the diffusion coefficient of air in PDMS. For example, if w = 50 μm and D = 3.4 × 10−9 m2 s−1 (diffusion coefficient of air in PDMS at atomsphere), the characteristic time to initialize the steady-state air flux is tD1 ≈ 0.74 s. Another characteristic time to diminish the steady-state air flux would depend on the air diffusion from the surface into the microchamber across the thick PDMS layer (e.g., ~5 mm), which is tD2 ≈ 2 h. Therefore, the air flux will be kept steady and constant if devices are operated within tD1 ≪ t ≪ tD2.
3 Design and fabrication
3.2 Microfluidic device fabrication
All devices were fabricated by a standard soft lithography process (Xia and Whitesides 1998). A 3-inch silicon wafer with one side polished (University wafers, South Boston, MA, USA) was submerged into buffered hydrofluoric acid (BHF) at room temperature for 5 min to remove the thin native silicon dioxide layer. After that, the wafer was cleaned by using acetone and methanol, respectively, and then rinsed in deionized water before blown dry by filtered nitrogen gas. After cleaning, the cleaned wafer was placed on a hot plate at 120 °C for 5 min in order to make it completely dehydrated. SU-8 (SU-8 2050, Micro-Chem Corp, Newton, MA, USA) was then spin coated on top of the wafer by using the spin coater (WS-650Mz NPP from Laurell Technologies, North Wales, PA, USA) to the target thickness. After spin coating, soft bake was performed on a leveled hot plate for 3 and 9 min at 65 and 95 °C, respectively. After soft bake, UV photolithography was carried out by using a contact mask aligner. After UV exposure, the post-exposure bake (PEB) was conducted on the leveled hot plate for 2 and 7 min at 65 and 95 °C, respectively, followed by development in SU-8 developer for 5 min. After developing, the wafer was cleaned with isopropyl alcohol. Finally, the wafer was blown dry with filtered nitrogen gas and then placed on the hot plate at 100 °C for 5 min to evaporate any residual of isopropyl alcohol.
A prepolymer of PDMS (Sylgard 184, Dow Corning) and corresponding curing agent was thoroughly mixed at a ratio of 10: 1 (wt/wt). Then, the mixed PDMS was degassed in a vacuum chamber for 20 min to remove all the air bubbles. In order to peel off the PDMS from the wafer mold easily, hexamethyldisilazane (Sigma Aldrich, Saint Louis, MO, USA) was silanized on the surface of the wafer mold in vacuum chamber at room temperature for 30 min. After that, the PDMS mixture was carefully poured onto the wafer mold and then cured at 80 °C for 3 h. In order to make sure the thickness of the PDMS bulk does not affect the experiment, the thickness of all the PDMS devices was made to be more than 1 cm. After PDMS was cured and peeled off from the wafer mold, holes were punched on the PDMS replicas for the connection with the syringe, followed by oxygen plasma treatment for irreversible bonding between PDMS and glass slide at the top surface. Lastly, the device was baked over a hot plate for 48 h at 70 °C to improve the bonding strength and stabilize the surface property of the devices.
3.3 Test setup and procedure
The fabricated microfluidic device was connected through a silicone tube to a glass syringe with maximum volume of 200 μl. The glass syringe was used to generate the low pressure inside the microchamber of the fabricated microfluidic devices. In order to verify that the volume of the syringe does not affect the flow rate, a glass syringe with maximum volume of 100 μl was also used in the same devices. And the results were similar to the glass syringe with maximum volume of 100 μl (the variation in flow rate was within 10 %). Therefore, we adopted the glass syringe with maximum volume of 200 μl in all the following tests. All of the devices were tested three times. First, the syringe was connected and pulled to have volume expansion to 200 μl (as shown in Step I and II in Fig. 1d). After pulling the plunger, DI water was loaded at the inlets after around 5 s (as shown in Step III in Fig. 1d). All the flow processes were recorded by a Nikon stereo-type microscope and camera set. By analyzing the recorded video clips, volume–time curves were plotted (KDS100 W, Fisher Scientific, IL, USA).
4 Results and discussion
As shown in Fig. 2b, before liquid reaches the part of the channel surrounded by the microchamber, the liquid volume V is linear to the pumping time t, which indicates that the flow rate Q remains constant (Phase I). As the flow enters the part of the channel surrounded by the microchamber after a certain time tC, however, the flow rate exponentially decreases (Phase II).
Similarly, we investigated the syringe-assisted pumping related to the overlap area, while the thickness of the PDMS wall was kept unchanged (w = 50 μm), as shown in Fig. 3a. The pumped total volume was plotted in Fig. 3b. In Phase II, exponential decay was shown with the same time constant τ for different overlap areas. This is due to the fixed PDMS wall thickness, resulting in the same air flux F for all cases. The flow rate of the device with almost zero overlap area (S0 = d × h = 0.0035 mm2) was extremely low (0.089 nl s−1), taking ~9 min to draw in water completely. As only the very end of the channel overlapped with the microchamber, the pumped total volume was linear with time during the whole pumping process (Fig. 3c). In Phase I, as expected from Eq. 3, the flow rate QI was linearly proportional to the overlap surface area (Fig. 3d).
When t ≈ tD2, the syringe-assisted pumping will stop because the pressure difference completely vanishes inside the microchamber due to the air diffusion from outside to the microchamber. However, we can restore the low pressure or vacuum inside the microchamber by simply reconnecting the syringe and pulling the plunger again. Therefore, unlike the vacuum-assisted degas-driven flow, the proposed syringe-assisted method permits instantaneous, recurring, point-of-care pumping. Another advantage of the syringe-assisted pumping is that the air diffusion is bidirectional between the channel and microchamber depending on the polarity of pressure difference. By pushing the plunger, the drawn-in liquid solution can be retrieved from the dead-end channel. Thus, the syringe can supply not only a vacuum source (e.g., low pressure) but also a pressure source (e.g., high pressure) to drive the fluid flow into and from the dead-end channel. This will enable a simple point-of-care pumping system that requires multiple incubations and washing processes.
By adjusting the thickness of the PDMS wall and overlap area between the channel and microchamber, we have controlled the flow rate and generated a constant flow before the fluid reaches the part of channel surrounded by the microchamber. By the proposed method, a syringe-assisted, instantaneous, recurring, bidirectional, point-of-care pumping system without external pumps or vacuum chambers has been realized with controllable flow rate ranging from 0.089 to 4 nl s−1.
This work was partially supported by grants from NSF (ECCS-1002255 and ECCS-0736501).
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