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Disposable thermo-pneumatic micropump for bio lab-on-a-chip application
Microelectronic Engineering 86 (2009) 1337–1339
Contents lists available at ScienceDirect
Microelectronic Engineering journal homepage: www.elsevier.com/locate/mee
Disposable thermo-pneumatic micropump for bio lab-on-a-chip application Seung-Mo Ha a, Woong Cho a, Yoomin Ahn b,* a b
Dept. of Mechanical Engineering, Graduate School, Hanyang University, Seoul 133-791, Republic of Korea Dept. of Mechanical Engineering, Hanyang University, 1271 Sa 3-dong, Sangnok-gu, Ansan, Gyeonggi 426-791, Republic of Korea
a r t i c l e
i n f o
Article history: Received 27 September 2008 Received in revised form 5 December 2008 Accepted 16 December 2008 Available online 25 December 2008 Keywords: Thermo-pneumatic micropump Microscale check valve Detachable microfluidic device PDMS-glass hybrid chip
a b s t r a c t We present a disposable thermo-pneumatic micropump with microscale check valves for easy incorporation in a fully integrated biochip. The microfluidic device is comprised of biocompatible materials: glass and polydimethylsiloxane (PDMS). The chip consists of a glass substrate and three PDMS layers. In the chip, a Cr/Au heater is fabricated by micromachining on a glass substrate, while the flow channel, pump chamber, and valves are made by replica molding using layers of PDMS. The glass-based heater chip and the PDMS-based micropump chip are assembled and clamped together to enable operation of the system, and then the components are separable such that the heater chip can be used repeatedly and the PDMS chip can be used in a disposable manner. The operation of the micropump was investigated experimentally, as well as through simulations using fluid dynamics software. Superior pumping performance was achieved when the duty ratio of the voltage applied to the heater was 33%. Ó 2008 Elsevier B.V. All rights reserved.
1. Introduction Micropumps are a fundamental component of microfluidic devices [1]. By utilizing the micrpump, more advanced results can be obtained in the microsystems for drug delivery and biomedical assay. To use micropumps in promising market areas such as bio lab-on-a-chip devices, achieving lowering manufacturing costs while maintaining performance is essential [2]. Among the various types of micropumps, manufacturing costs are generally lowest for thermo-pneumatic micropumps, even though they tend to produce low flow rates relative to their size [3]. Therefore, thermo-pneumatic micropumps appear to be promising micropumps for biomedical applications. However, most thermo-pneumatic micropumps are fabricated with using silicon, which is not biocompatible and is an opaque material [4,5], as such these pumps are not suitable for biochips. For biochip application, polymerbased micropumps and microvalves are currently being studied for biochip applications [6,7]. Using well-established printed circuit board (PCB) technology, a polymer-PCB thermo-pneumatic micropump was developed with very moderate fabrication costs [6]. Typical bio lab-on-a-chip devices are hybrid PDMS-glass chips. Hence, the polymer-PCB thermo-pneumatic micropumps are not suitable for integration in bio lab-on-a-chip devices. In addition, these polymer thermo-pneumatic micropumps are still expensive for use in disposable devices, which are used only once to prevent contamination of the biological sample. As a result, micropumps
* Corresponding author. Tel.: +82 31 400 5281; fax: +82 31 406 5550. E-mail address: [email protected] (Y. Ahn). 0167-9317/$ - see front matter Ó 2008 Elsevier B.V. All rights reserved. doi:10.1016/j.mee.2008.12.046
used in bio lab-on-a-chip devices must be improved to reduce costs. In our work, an improved polymer-based thermo-pneumatic micropump was designed with the goal of reduced fabrication costs for use in disposable devices. A low-cost polydimethylsiloxane (PDMS)-glass hybrid chip integrating a thermo-pneumatic micropump and microscale check valves was developed for applications in bio lab-on-a-chip devices. The micropump and microvalves were made of biocompatible and transparent materials (PDMS and glass), so as to be applicable to biochips. For biocompatibility and disposability reasons, the costly microheater portion of the pumping system was designed to be reused by detaching it from the disposable micropump portion of the system that is in contact with the biological sample. In addition, by using passive check valves, backward flow and fluid leakage were blocked and flow control was stable and precise. 2. Design and fabrication A schematic of the PDMS-glass micropump system is shown in Fig. 1. Voltage applied to a microheater unit fabricated on a glass substrate generates heat via electrical resistance. Air in the microheater chamber expands as the temperature increases and deforms a PDMS membrane located over the heater chamber. Consequently, fluid over the membrane is discharged to the outside of the pump chamber. Next, the applied voltage is cut off, restoring the membrane to the former state, which causes fluid to flow into the chamber. To ensure pumping in one direction only, PDMS check valves were fabricated and located at the inlet and outlet of the pump chamber. If fluid flows in the forward direction, the PDMS valve
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S.-M. Ha et al. / Microelectronic Engineering 86 (2009) 1337–1339
1.13 W m 1 K 1, respectively. The assumed convection heat transfer coefficient was 10 W m 2 °C 1. In the pumping experiments, the signal for a DC pulse voltage was generated with a desired frequency and duty ratio using a commercial control program (LabVIEW 7.0, National Instruments, US) and was sent to a signal conditioner (SC2345, National Instruments) through a data acquisition board (PCI-6024E, National Instruments). The filtered signal from the signal conditioner was then transferred to a power supply (E3631A, Agilent). Finally, the amplified voltage was applied to the microheater. To measure the temperature of the microheater chamber, a K-type thermocouple was inserted in the chip. The micropump was filled with deionized water by syringe and then operated to measure the flow rate without backpressure. The micropump was tested using various square pulse voltages applied to the microheater. Flow rates were measured after a sufficient period of operation and under zero pressure difference between the inlet and outlet. Flow rates were calculated by gauging the amount of fluid that passed through the outlet tube during a pumping time of 30 min. Fig. 1. Configuration of the micropump with microvalve: (a) cross-sectional view and (b) schematic of detachable microfluidic device.
flap is displaced by the fluid pressure and the gap between the flap and the valve seat widens, such that the valve opens. If fluid flows in the backward direction, the flap adheres to the valve seat and the valve closes. In order to use the micropump in a disposable manner, the glass-based heater chip was designed to be detachable from the PDMS-based pump chip. The PDMS chip and the glass chip are mechanically assembled using a jig and a clamp, as shown in Fig. 1b. The inexpensive PDMS chip can be discarded after use, but the relatively expensive microheater chip can be reused. The microheater was made using a Cr/Au (20/100 nm) thin film, which was deposited on a 500 lm thick glass wafer (Pyrex 7740, Dow Corning, US) by thermal evaporation and patterned by photolithography. Chambers, valves, and channels were micromachined in PDMS by replica molding. The PDMS chip was composed of three layers that were bonded together by O2 plasma treatment. During plasma treatment, valve areas were locally masked to prevent bonding between the flap and the valve seat. Inlet and outlet access holes (1 mm in diameter) were mechanically punched in the top PDMS layer. The diameter of the chamber was 7 mm; the width and depth of the channels were 200 and 180 lm, respectively; the thickness of the PDMS membrane was 300 lm. One of promising applications of the micropump is drug delivery. Maximum flow rates of reported micropumps for the drug delivery are 0.1– 3500 ll/min [8]. In this study, the pumping system was designed for the maximum flow rate of 1000 ll/min. To produce this flow rate, deformation of the membrane was calculated as about 1.3 mm at 1.0 Hz frequency. This is within the height of the chamber, so the system is expected to operate properly.
4. Results and discussion Deformation at the center of the PDMS membrane covering the heater chamber was simulated as a function of the temperature of the chamber. The deformation was negligible when the temperature was less than about 40 °C but increased exponentially above 40 °C. The variation of the temperature of the heater chamber as a function of the power applied to the heater was simulated and measured, as shown in Fig. 2. An electrical power of 7.2 W was applied to the heater for 30 s, and then the chip was allowed to cool by convection for 90 s to a temperature of 20 °C. To investigate the performance of the check valves, deionized water was injected using a syringe into the fabricated micropump device. Injection was possible in the forward direction but was impossible in the backward direction. Fig. 3 shows the variation of the pumping flow rate according to the applied electrical power (duty ratio of 50%) for three different frequencies. In general, the flow rate increased with increasing electrical power. This is because the temperature and pressure of the heater chamber increases with an increase in electrical power, resulting in the membrane deforming greatly and an increase in the flow rate. In Fig. 3, flow rate was also affected by the applied electrical frequency. The flow rate increased as the frequency was increased for applied power levels less than about 4.0 W. When the power was greater than about 4.0 W, a maximum flow rate was reached using a frequency of 3.0 Hz, and the flow rates were smaller for frequencies less than or greater than 3.0 Hz. This figure reveals that
3. Simulations and experiments Commercial fluid dynamic analysis programs (CFD ACE+, ESI Group, France) were used to simulate the temperature of the heater chamber and the deformation of the PDMS membrane as a function of electrical power applied to the heater. In the simulations, fluid in the chip was assumed to be deionized water. The microheater was assumed to be a wall heat source. In the heat transfer analysis, heat conduction within the chip and free convection from the chip surface to ambient air were considered, but radiation from the chip was not considered. The ambient air was assumed to be 20 °C. The thermal conductivities of deionized water, PDMS, and glass were assumed to be 0.613, 0.17,
Fig. 2. Comparison between simulated results and experimental measurements of the air temperature in the chamber.
S.-M. Ha et al. / Microelectronic Engineering 86 (2009) 1337–1339
Fig. 3. Flow rate as a function of applied power (duty ratio: 50%) for different frequencies.
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ber deforms to a greater extent when the voltage-on period is greater than the voltage-off period for the pulsed voltage. However, the membrane cannot return fully to the original state for high duty ratios, so the flow rate is not stable. Meanwhile, when the duty ratio is low, the membrane deformation is small due to the short duration of applied power. Hence, the flow rate is not greater than about 54 ll/min, but the flow rate is constant because the membrane deformation is completely restored between pulses. If the duty ratio is less than 33%, the flow rate is constant of about 7 ll/min but too slow. Therefore, when the duty ratio is optimal, the flow rate should be controlled constantly and be as fast as possible. In silicon-glass thermo-pneumatic micropumps with integrated microscale check valves in which the pump membrane is made of silicon, a maximum flow rate of 34–55 ll/min was obtained and the operating frequency was 1–5 Hz [4,5]. A polymer-based thermo-pneumatic micropump with microscale check valves and a polyimide membrane was reported by Wego and Pagel [6]. The polymer thermo-pneumatic micropump produced a maxium flow rate of 530 ll/min at an operating frequency of 2 Hz. Meanwhile, the maximum flow rate of the newly developed micropump was measured as 50–80 ll/min at an operating frequency of less than 1.0 Hz. Therefore, the frequency of the new micropump is not much lower; the flow rate is somewhat better than for the silicon membrane micropump but is still slower than that for the polyimide membrane micropump. Compared with other micropumps for the drug delivery, the new micropump is needed lower power but provides smaller flow rate [8].
5. Conclusions
Fig. 4. Flow rate as a function of applied power (frequency: 0.1 Hz) for different duty ratios.
the optimum frequency for producing a maximum flow rate varies with the applied electrical power. Similarly, previous studies on thermo-pneumatic micropumps using thin films of silicon or polyimide reported that the flow rate increased with the electrical frequency, but beyond a threshold frequency, the rate decreased as the frequency increased [4,6]. This may be related to the results of Fig. 2, where the heater temperature increased and decreased nonlinearly with the duration of applied electrical power. The reason is thought to be that at high frequency, the response of the heater chamber to cooling was not fast enough for the deformed membrane to revert to its original state. For effective pumping, it is important to find the optimum duty ratio, which is the ratio of cooling time to heating time of the chamber. At the optimum duty ratio, the membrane of the pump chamber is able to completely return to the initial state before being deformed again with the next pulse. The effect of the duty ratio on the flow rate is shown in Fig. 4. At a frequency of 0.1 Hz, the flow rate increased as the duty ratio increased. The increase in flow rate was not constant with respect to the applied electrical power at a duty ratio of 50%. At a duty ratio of 33%, the flow rate was not greater than that at a duty ratio of 55%, but the flow rate increased linearly with applied power. At a relatively low duty ratio of 10%, the increase in flow rate was linear with respect to applied power, but the flow rate was not fast. If the duty ratio is high, the flow rate is fast because the membrane over the pump cham-
In this paper, a PDMS-based thermo-pneumatic micropump with micro check valves was designed and fabricated. A two-part detachable device was developed for application in disposable biochips, in which the comparatively high-cost microheater portion can be isolated after use and is reusable. The pumping system was simulated by a finite element analysis program, and the effect of the applied electrical power, voltage pulse frequency, and duty ratio on the flow rate was investigated. The pumping rate generally increased as the applied power and the duty ratio increased. However, as the frequency increased, the flow rate increased and then decreased. In the pumping experiments, the most stable and optimum performance of the fabricated micropump was obtained at a duty ratio of about 33% for a frequency of 0.1 Hz, which agreed with simulation results. The developed micropump may be economically and practically applicable to bio lab-on-a-chip devices due to the low fabrication costs, simplicity, and moderate flow rates that can be achieved with the micropump. References [1] [2] [3] [4] [5] [6] [7] [8]
S. Shoji, M. Esashi, J. Micromech. Microeng. 4 (1994) 157–171. P. Woias, Sens. Actuator B – Chem. 105 (2005) 28–38. D.J. Laser, J.G. Santiago, J. Micromech. Microeng. 14 (2004) R35–R64. F.C.M. Van De Pol, H.T.G. Van Lintel, M. Elwenspoek, J.H.J. Fluitman, Sens. Actuator A – Phys. 21–23 (1990) 198–202. M. Elwenspoek, T.S.J. Lammerink, R. Miyake, J.H.J. Fluitman, J. Micromech. Miroeng. 4 (1994) 227–245. A. Wego, L. Pagel, Sens. Actuator A – Phys. 88 (2001) 220–226. S.-W. Lee, D.-J. Kim, Y. Ahn, Y.G. Chai, Proc. Inst. Mech. Eng. Part C – J. Eng. Mech. Eng. Sci. 220 (2006) 1283–1288. A. Nisar, N. Afzulpurkar, B. Mahaisavariya, A. Tuantranont, Sens. Actuator B – Chem. 130 (2008) 917–942.
Contents lists available at ScienceDirect
Microelectronic Engineering journal homepage: www.elsevier.com/locate/mee
Disposable thermo-pneumatic micropump for bio lab-on-a-chip application Seung-Mo Ha a, Woong Cho a, Yoomin Ahn b,* a b
Dept. of Mechanical Engineering, Graduate School, Hanyang University, Seoul 133-791, Republic of Korea Dept. of Mechanical Engineering, Hanyang University, 1271 Sa 3-dong, Sangnok-gu, Ansan, Gyeonggi 426-791, Republic of Korea
a r t i c l e
i n f o
Article history: Received 27 September 2008 Received in revised form 5 December 2008 Accepted 16 December 2008 Available online 25 December 2008 Keywords: Thermo-pneumatic micropump Microscale check valve Detachable microfluidic device PDMS-glass hybrid chip
a b s t r a c t We present a disposable thermo-pneumatic micropump with microscale check valves for easy incorporation in a fully integrated biochip. The microfluidic device is comprised of biocompatible materials: glass and polydimethylsiloxane (PDMS). The chip consists of a glass substrate and three PDMS layers. In the chip, a Cr/Au heater is fabricated by micromachining on a glass substrate, while the flow channel, pump chamber, and valves are made by replica molding using layers of PDMS. The glass-based heater chip and the PDMS-based micropump chip are assembled and clamped together to enable operation of the system, and then the components are separable such that the heater chip can be used repeatedly and the PDMS chip can be used in a disposable manner. The operation of the micropump was investigated experimentally, as well as through simulations using fluid dynamics software. Superior pumping performance was achieved when the duty ratio of the voltage applied to the heater was 33%. Ó 2008 Elsevier B.V. All rights reserved.
1. Introduction Micropumps are a fundamental component of microfluidic devices [1]. By utilizing the micrpump, more advanced results can be obtained in the microsystems for drug delivery and biomedical assay. To use micropumps in promising market areas such as bio lab-on-a-chip devices, achieving lowering manufacturing costs while maintaining performance is essential [2]. Among the various types of micropumps, manufacturing costs are generally lowest for thermo-pneumatic micropumps, even though they tend to produce low flow rates relative to their size [3]. Therefore, thermo-pneumatic micropumps appear to be promising micropumps for biomedical applications. However, most thermo-pneumatic micropumps are fabricated with using silicon, which is not biocompatible and is an opaque material [4,5], as such these pumps are not suitable for biochips. For biochip application, polymerbased micropumps and microvalves are currently being studied for biochip applications [6,7]. Using well-established printed circuit board (PCB) technology, a polymer-PCB thermo-pneumatic micropump was developed with very moderate fabrication costs [6]. Typical bio lab-on-a-chip devices are hybrid PDMS-glass chips. Hence, the polymer-PCB thermo-pneumatic micropumps are not suitable for integration in bio lab-on-a-chip devices. In addition, these polymer thermo-pneumatic micropumps are still expensive for use in disposable devices, which are used only once to prevent contamination of the biological sample. As a result, micropumps
* Corresponding author. Tel.: +82 31 400 5281; fax: +82 31 406 5550. E-mail address: [email protected] (Y. Ahn). 0167-9317/$ - see front matter Ó 2008 Elsevier B.V. All rights reserved. doi:10.1016/j.mee.2008.12.046
used in bio lab-on-a-chip devices must be improved to reduce costs. In our work, an improved polymer-based thermo-pneumatic micropump was designed with the goal of reduced fabrication costs for use in disposable devices. A low-cost polydimethylsiloxane (PDMS)-glass hybrid chip integrating a thermo-pneumatic micropump and microscale check valves was developed for applications in bio lab-on-a-chip devices. The micropump and microvalves were made of biocompatible and transparent materials (PDMS and glass), so as to be applicable to biochips. For biocompatibility and disposability reasons, the costly microheater portion of the pumping system was designed to be reused by detaching it from the disposable micropump portion of the system that is in contact with the biological sample. In addition, by using passive check valves, backward flow and fluid leakage were blocked and flow control was stable and precise. 2. Design and fabrication A schematic of the PDMS-glass micropump system is shown in Fig. 1. Voltage applied to a microheater unit fabricated on a glass substrate generates heat via electrical resistance. Air in the microheater chamber expands as the temperature increases and deforms a PDMS membrane located over the heater chamber. Consequently, fluid over the membrane is discharged to the outside of the pump chamber. Next, the applied voltage is cut off, restoring the membrane to the former state, which causes fluid to flow into the chamber. To ensure pumping in one direction only, PDMS check valves were fabricated and located at the inlet and outlet of the pump chamber. If fluid flows in the forward direction, the PDMS valve
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S.-M. Ha et al. / Microelectronic Engineering 86 (2009) 1337–1339
1.13 W m 1 K 1, respectively. The assumed convection heat transfer coefficient was 10 W m 2 °C 1. In the pumping experiments, the signal for a DC pulse voltage was generated with a desired frequency and duty ratio using a commercial control program (LabVIEW 7.0, National Instruments, US) and was sent to a signal conditioner (SC2345, National Instruments) through a data acquisition board (PCI-6024E, National Instruments). The filtered signal from the signal conditioner was then transferred to a power supply (E3631A, Agilent). Finally, the amplified voltage was applied to the microheater. To measure the temperature of the microheater chamber, a K-type thermocouple was inserted in the chip. The micropump was filled with deionized water by syringe and then operated to measure the flow rate without backpressure. The micropump was tested using various square pulse voltages applied to the microheater. Flow rates were measured after a sufficient period of operation and under zero pressure difference between the inlet and outlet. Flow rates were calculated by gauging the amount of fluid that passed through the outlet tube during a pumping time of 30 min. Fig. 1. Configuration of the micropump with microvalve: (a) cross-sectional view and (b) schematic of detachable microfluidic device.
flap is displaced by the fluid pressure and the gap between the flap and the valve seat widens, such that the valve opens. If fluid flows in the backward direction, the flap adheres to the valve seat and the valve closes. In order to use the micropump in a disposable manner, the glass-based heater chip was designed to be detachable from the PDMS-based pump chip. The PDMS chip and the glass chip are mechanically assembled using a jig and a clamp, as shown in Fig. 1b. The inexpensive PDMS chip can be discarded after use, but the relatively expensive microheater chip can be reused. The microheater was made using a Cr/Au (20/100 nm) thin film, which was deposited on a 500 lm thick glass wafer (Pyrex 7740, Dow Corning, US) by thermal evaporation and patterned by photolithography. Chambers, valves, and channels were micromachined in PDMS by replica molding. The PDMS chip was composed of three layers that were bonded together by O2 plasma treatment. During plasma treatment, valve areas were locally masked to prevent bonding between the flap and the valve seat. Inlet and outlet access holes (1 mm in diameter) were mechanically punched in the top PDMS layer. The diameter of the chamber was 7 mm; the width and depth of the channels were 200 and 180 lm, respectively; the thickness of the PDMS membrane was 300 lm. One of promising applications of the micropump is drug delivery. Maximum flow rates of reported micropumps for the drug delivery are 0.1– 3500 ll/min [8]. In this study, the pumping system was designed for the maximum flow rate of 1000 ll/min. To produce this flow rate, deformation of the membrane was calculated as about 1.3 mm at 1.0 Hz frequency. This is within the height of the chamber, so the system is expected to operate properly.
4. Results and discussion Deformation at the center of the PDMS membrane covering the heater chamber was simulated as a function of the temperature of the chamber. The deformation was negligible when the temperature was less than about 40 °C but increased exponentially above 40 °C. The variation of the temperature of the heater chamber as a function of the power applied to the heater was simulated and measured, as shown in Fig. 2. An electrical power of 7.2 W was applied to the heater for 30 s, and then the chip was allowed to cool by convection for 90 s to a temperature of 20 °C. To investigate the performance of the check valves, deionized water was injected using a syringe into the fabricated micropump device. Injection was possible in the forward direction but was impossible in the backward direction. Fig. 3 shows the variation of the pumping flow rate according to the applied electrical power (duty ratio of 50%) for three different frequencies. In general, the flow rate increased with increasing electrical power. This is because the temperature and pressure of the heater chamber increases with an increase in electrical power, resulting in the membrane deforming greatly and an increase in the flow rate. In Fig. 3, flow rate was also affected by the applied electrical frequency. The flow rate increased as the frequency was increased for applied power levels less than about 4.0 W. When the power was greater than about 4.0 W, a maximum flow rate was reached using a frequency of 3.0 Hz, and the flow rates were smaller for frequencies less than or greater than 3.0 Hz. This figure reveals that
3. Simulations and experiments Commercial fluid dynamic analysis programs (CFD ACE+, ESI Group, France) were used to simulate the temperature of the heater chamber and the deformation of the PDMS membrane as a function of electrical power applied to the heater. In the simulations, fluid in the chip was assumed to be deionized water. The microheater was assumed to be a wall heat source. In the heat transfer analysis, heat conduction within the chip and free convection from the chip surface to ambient air were considered, but radiation from the chip was not considered. The ambient air was assumed to be 20 °C. The thermal conductivities of deionized water, PDMS, and glass were assumed to be 0.613, 0.17,
Fig. 2. Comparison between simulated results and experimental measurements of the air temperature in the chamber.
S.-M. Ha et al. / Microelectronic Engineering 86 (2009) 1337–1339
Fig. 3. Flow rate as a function of applied power (duty ratio: 50%) for different frequencies.
1339
ber deforms to a greater extent when the voltage-on period is greater than the voltage-off period for the pulsed voltage. However, the membrane cannot return fully to the original state for high duty ratios, so the flow rate is not stable. Meanwhile, when the duty ratio is low, the membrane deformation is small due to the short duration of applied power. Hence, the flow rate is not greater than about 54 ll/min, but the flow rate is constant because the membrane deformation is completely restored between pulses. If the duty ratio is less than 33%, the flow rate is constant of about 7 ll/min but too slow. Therefore, when the duty ratio is optimal, the flow rate should be controlled constantly and be as fast as possible. In silicon-glass thermo-pneumatic micropumps with integrated microscale check valves in which the pump membrane is made of silicon, a maximum flow rate of 34–55 ll/min was obtained and the operating frequency was 1–5 Hz [4,5]. A polymer-based thermo-pneumatic micropump with microscale check valves and a polyimide membrane was reported by Wego and Pagel [6]. The polymer thermo-pneumatic micropump produced a maxium flow rate of 530 ll/min at an operating frequency of 2 Hz. Meanwhile, the maximum flow rate of the newly developed micropump was measured as 50–80 ll/min at an operating frequency of less than 1.0 Hz. Therefore, the frequency of the new micropump is not much lower; the flow rate is somewhat better than for the silicon membrane micropump but is still slower than that for the polyimide membrane micropump. Compared with other micropumps for the drug delivery, the new micropump is needed lower power but provides smaller flow rate [8].
5. Conclusions
Fig. 4. Flow rate as a function of applied power (frequency: 0.1 Hz) for different duty ratios.
the optimum frequency for producing a maximum flow rate varies with the applied electrical power. Similarly, previous studies on thermo-pneumatic micropumps using thin films of silicon or polyimide reported that the flow rate increased with the electrical frequency, but beyond a threshold frequency, the rate decreased as the frequency increased [4,6]. This may be related to the results of Fig. 2, where the heater temperature increased and decreased nonlinearly with the duration of applied electrical power. The reason is thought to be that at high frequency, the response of the heater chamber to cooling was not fast enough for the deformed membrane to revert to its original state. For effective pumping, it is important to find the optimum duty ratio, which is the ratio of cooling time to heating time of the chamber. At the optimum duty ratio, the membrane of the pump chamber is able to completely return to the initial state before being deformed again with the next pulse. The effect of the duty ratio on the flow rate is shown in Fig. 4. At a frequency of 0.1 Hz, the flow rate increased as the duty ratio increased. The increase in flow rate was not constant with respect to the applied electrical power at a duty ratio of 50%. At a duty ratio of 33%, the flow rate was not greater than that at a duty ratio of 55%, but the flow rate increased linearly with applied power. At a relatively low duty ratio of 10%, the increase in flow rate was linear with respect to applied power, but the flow rate was not fast. If the duty ratio is high, the flow rate is fast because the membrane over the pump cham-
In this paper, a PDMS-based thermo-pneumatic micropump with micro check valves was designed and fabricated. A two-part detachable device was developed for application in disposable biochips, in which the comparatively high-cost microheater portion can be isolated after use and is reusable. The pumping system was simulated by a finite element analysis program, and the effect of the applied electrical power, voltage pulse frequency, and duty ratio on the flow rate was investigated. The pumping rate generally increased as the applied power and the duty ratio increased. However, as the frequency increased, the flow rate increased and then decreased. In the pumping experiments, the most stable and optimum performance of the fabricated micropump was obtained at a duty ratio of about 33% for a frequency of 0.1 Hz, which agreed with simulation results. The developed micropump may be economically and practically applicable to bio lab-on-a-chip devices due to the low fabrication costs, simplicity, and moderate flow rates that can be achieved with the micropump. References [1] [2] [3] [4] [5] [6] [7] [8]
S. Shoji, M. Esashi, J. Micromech. Microeng. 4 (1994) 157–171. P. Woias, Sens. Actuator B – Chem. 105 (2005) 28–38. D.J. Laser, J.G. Santiago, J. Micromech. Microeng. 14 (2004) R35–R64. F.C.M. Van De Pol, H.T.G. Van Lintel, M. Elwenspoek, J.H.J. Fluitman, Sens. Actuator A – Phys. 21–23 (1990) 198–202. M. Elwenspoek, T.S.J. Lammerink, R. Miyake, J.H.J. Fluitman, J. Micromech. Miroeng. 4 (1994) 227–245. A. Wego, L. Pagel, Sens. Actuator A – Phys. 88 (2001) 220–226. S.-W. Lee, D.-J. Kim, Y. Ahn, Y.G. Chai, Proc. Inst. Mech. Eng. Part C – J. Eng. Mech. Eng. Sci. 220 (2006) 1283–1288. A. Nisar, N. Afzulpurkar, B. Mahaisavariya, A. Tuantranont, Sens. Actuator B – Chem. 130 (2008) 917–942.