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A novel constant delivery thermopneumatic micropump using surface tensions
Sensors and Actuators A 139 (2007) 210–215
A novel constant delivery thermopneumatic micropump using surface tensions Do Han Jun, Woo Young Sim, Sang Sik Yang ∗ School of Electronics Engineering, Ajou University, San 5 Wonchun-Dong, Yeongtong-Gu, Suwon 443-749, Republic of Korea Received 30 July 2006; received in revised form 11 April 2007; accepted 11 April 2007 Available online 19 April 2007
Abstract This paper presents the fabrication and test of a novel thermopneumatic micropump made of PDMS without a membrane or a valve. The driving force of the liquid discharge is the pneumatic pressure of heated air and the suction of liquid is accomplished by the surface tension. The proposed micropump has a pump chamber and a pair of air chambers. The air chambers have electric heaters on the Pyrex glass substrate. Each air chamber is connected with the pump chamber through two different air channels which behaves like check valves. The micropump is fabricated through the spin-coating process, the lithograph process, the PDMS molding process, etc. The total size of micropump is 11.7 mm × 8.8 mm × 0.7 mm. The FEM simulation was performed with air inflow and suction instead of air heating. The fabricated micropump was tested for various backward pressures ranging from 0 to 15 cm H2 O. The simulation and the test results confirmed the expected behavior of the micropump. The discharge volume is 116 nL when the input voltage of 3.5 V is applied for 4 s. This proposed micropump is feasible for disposable transdermal drug delivery system. © 2007 Elsevier B.V. All rights reserved. Keywords: Thermopneumatic; Surface tension; Micropump; Constant delivery
1. Introduction Recently microelectromechanical systems (MEMS) technology has been applied to bio-medical systems. Especially various microfluidic devices have been developed for the application to drug delivery systems (DDS). A drug delivery system releasing accurate amount of drug in time at a desired location can reduce negative side effects due to the excessive dosage that often happens in conventional drug delivery systems. The important objective of a drug delivery system is the enhancement of the treatment effectiveness. It can be achieved by the precise control of the drug delivery amount. Several groups have developed liquid dispensing actuators designed specially for drug delivery systems [1–4]. The drug delivery systems have pump-like structures that can deliver drug continuously or measure the flow rate for the accurate drug delivery. The pumps are miniaturized and have high flow rate. It, however, is difficult to fabricate since the structures are complicated. It is needed to investigate into
∗
Corresponding author. Tel.: +82 331 219 2481; fax: +82 331 212 9531. E-mail address: [email protected] (S.S. Yang).
0924-4247/$ – see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.sna.2007.04.038
a micropump so that we can reduce the cost and simplify the structure without losing the accurateness. In this paper, we propose a micropump that discharges a constant amount of liquid by using a bubble expansion by electric heating. We use the surface tension and capillary attraction in the pump chamber and channels to make one-way flow. This micropump has simple structure and is easy to fabricate. To design the shape and the size of the micropump, we perform the simulation by the finite element method (FEM) using CFD-ACE+. We prove the behavior of this micropump by experiments. 2. Micropump structure and design The schematic diagram of the micropump structure is shown in Fig. 1. The micropump consists of an inlet, an outlet, a microchannel, a pump chamber, two air chambers, two discharge air channels and two suction air channels. The air chambers have electric heaters on the Pyrex glass substrate. The pump chamber and the air chambers are connected with the four air channels through which the air is discharged and sucked. Microchannels connect the pump chamber and the inlet and the outlet port. The pump chamber shape is designed to reduce backward
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
surface tension, σ as Eqs. (1) and (2) [5,6]: sin α cos θ cos(θ + β) P = 2σ + α h w
211
(1)
where α is the fluid meniscus angle given by α=θ+β−
Fig. 1. The structure of the micropump.
flow to the inlet and block the inflow from the outlet by surface tensions. The micropump operation sequence is shown in Fig. 2. The one cycle of the operation includes a discharge stroke and a suction stroke. First by the capillary attraction the liquid drug flows into the micropump from inlet as Fig. 2(a). When we apply voltage to the heater, the air in the air chamber expands and pushes the liquid in the pump chamber through air discharge channels as Fig. 2(b). During this discharge stroke the liquid in the pump chamber is discharged to the outlet by the pneumatic pressure as Fig. 2(c). If we turn off the heater, the air contracts and the air is sucked into the air chamber through the air suction channels. During this suction stroke liquid is supplied from the inlet into the pump chamber by the negative pneumatic pressure and the capillary attraction force as Fig. 2(d). At this time, liquid inflow from the outlet is blocked by the surface tension. The shapes of the channels connected to the pump chamber are to be designed appropriately to accomplish the above operation. The liquid flow in the pump chamber is controlled by the surface tension and the capillary attraction. The capillary attraction, P is determined by the channel width, w, the channel height, h, the channel expansion angle, β, the contact angle, θ,
π 2
(2)
If P is positive, the fluid keeps spreading through the expansion channel. If P is negative, the fluid is trapped at the expansion channel. The shape of the inlet-side of the pump chamber is designed to provide enough capillary attraction to achieve selfpriming. In this micropump the expansion angle of the inlet-side of the pump chamber is set to 16◦ so that P is positive. Since the liquid inflow from the pump chamber to the air chamber must be blocked during the suction stroke, we set the expansion angle over 90◦ at the connection of the air discharge channel and the air chamber to make P negative. We set the expansion angle from the outlet channel to the pump chamber to 90◦ to prevent the liquid from flowing from the outlet into the pump chamber during the suction stroke. With this angle design, the expanded air in the pump chamber is sucked into the air chamber through only the air suction channels and the liquid flows from the inlet into the pump chamber until the whole pump chamber is filled. To obtain effective liquid discharge, the air bubbles must grow from the air discharge channels with the air suction channel blocked during the discharge stroke. Since the pneumatic force is proportional to the area of the channel and the capillary attraction force is proportional to the perimeter, we set the width of the air suction channel smaller than that of the air discharge channel so that the air is trapped at the air suction channel during the discharge stroke by the capillary attraction force larger than the pneumatic force. The pump chamber size must be greater than the desired liquid discharge volume in consideration of the dead volume which is figured out from the simulation. We determine the air chamber size enough for the expanded air bubble to push out the liquid by the desired discharge volume during the discharge stroke within the restriction of the moderate temperature rise.
Fig. 2. The operation sequence of the micropump: (a) before discharge stroke, (b) beginning of discharge stroke, (c) end of discharge stroke, and (d) suction stroke.
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Fig. 3. The simulation result: (a) the initial state, (b) air expansion, (c) bubbles merging, (d) drug discharge, and (e) air suction.
The heater resistance is set to 30 . The height of micropump structure is 80 m and the total dimension of the micropump is 11.7 cm × 8.8 cm × 0.7 cm. 3. Simulations We performed the simulation using CFD-ACE+ program for the performance evaluation of the microscale devices and for the detailed size determination. In this simulation we use continuum and incompressible Navier–Strokes equations as the governing equations. Since the flow inside the channel is unsteady and has two phases in this case, we cannot obtain reliable simulation results using only these two equations. We need to use volume of fluid (VOF) method to solve the free surface problem. This method has been proposed by Hirt and Nichols [7]. Recently this method was applied to several fluid cases by Rider and Kothe [8]. In this simulation we use piecewise linear interface construction (PLIC) modeling for reconstructing fluid shape. Also we consider the surface tension in the micropump because surface tension is the important parameter in microscale. Simulation was performed under several assumptions. In the simulation the contact angles of the glass and the PDMS are 10◦ and 100◦ , respectively. The inlet is set to be a wall and the outlet
is set to be open to the atmosphere. The inject rate of air from the air chamber to the pump chamber during the discharge stroke is 9 × 10−11 kg/s and the suction rate of air from the pump chamber to the air chamber during the suction stroke is 9 × 10−11 kg/s. In simulation result, the discharge volume is 124 nL and the discharge time is 2.6 s. Fig. 3 shows the FEM simulation result of the proposed pump. Fig. 3(a) is the initial state. Fig. 3(b) shows the air bubble growing from the air discharge channel and Fig. 3(c) shows the state before the two air bubbles merge in the pump chamber during the discharge stroke. Fig. 3(d) is the end state of the discharge stroke when the fluid is totally discharged from the pump chamber. Fig. 3(e) is the state before the end of the suction stroke and illustrates that the air is sucked into the air chamber with the air discharge channel blocked. The simulation result confirms the complete behavior of the proposed micropump. 4. Fabrication process Fig. 4 shows the fabrication process of the micropump. This micropump comprises three substrates. The upper Pyrex glass substrate has microheaters for electric heating. The middle PDMS substrate has all structures of the micropump except
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
213
Fig. 4. The fabrication process of the micropump.
the microheater. The bottom Pyrex glass substrate is a supporting structure to hold the PDMS structure. The fabrication process consisted of three steps. First for the ˚ layer is deposited and patterned microheater Ti/Au (500/1500 A) on the Pyrex glass substrate as upper layer. The heater resistance is about 30 . Then the inlet and the outlet are made by electrochemical discharge machining (ECDM) process on the glass wafer. The second step is to make SU-8 mold and PDMS layer. For SU-8 mold fabrication, negative thick photoresister (SU-8-
2100) is spun at 500 rpm for 40 s and at 2500 rpm for 60 s on the silicon wafer. The spin-coated photoresister is cured at 65 ◦ C for 25 min and at 90 ◦ C for 50 min. Then the micropump structure is patterned. A 10:1 mixture of PDMS prepolymer and the curing agent is mixed. This prepared PDMS mixture is poured onto the SU-8 mold and cured at 65 ◦ C for 3 h in a vacuum chamber. After curing the PDMS layer is released. Finally the released PDMS layer and Pyrex glass substrates are bonded after the treatment with O2 plasma for 10 s. The photograph of the fabricated micropump is shown in Fig. 5. Fig. 5(a) is a fabricated micropump image and Fig. 5(b) is a SEM image of PDMS layer. 5. Measurements and results In this paper, the high speed video camera is used for the observation of the micropump operations. The discharge time and the self-filling time of the micropump are approximately measured by the frame analysis of the recorded video data. The measurement setup of micropump is shown in Fig. 6. In this test the liquid is water. The discharge time is closely related with the input voltage. As the input voltage increases, the discharge time decreases since the micropump needs a proper input energy to accomplish one discharge cycle. The interval time is independent
Fig. 5. The image of the micropump: (a) photograph of the fabricated micropump and (b) SEM image of PDMS layer.
Fig. 6. The measurement setup.
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Fig. 7. The photograph during the operation: (a) t = 0 s, (b) t = 1.5 s, (c) t = 2.1 s, and (d) t = 4 s.
of the input voltage as far as the input energy is fixed. In this paper, I designed the heater resistance so that the micropump can be operated at the level of TTL output voltage. If the input voltage is low as much as 2 V, it takes over 10 s to discharge. The input voltage selected arbitrarily is 3.5 V which leads to a moderate discharge time less than 5 s. Fig. 7 shows the sequential capture images of the micropump during the operation for the input voltage of 3.5 V. Fig. 7(a)–(d) shows the meniscus of water and the air bubble when t = 0, 1.5, 2.1 and 4 s, respectively. The water and the air bubble movements observed in the experiment are the same as simulation result. The discharge volume of the micropump is calculated from the meniscus movement in the outlet silicone tube observed with
Fig. 8. The accumulated discharged volume when backward pressure is 0 cm H2 O.
a high speed video camera. When applied voltage is 3.5 V and no backward pressure exists, the discharge time is 4 s and selffilling time is 3 min and 52 s. The measured discharge volume is 116 nL. We measured the discharge volume ten times for the repeatability test under zero backward pressure every 5 and 10 min, respectively. Fig. 8 shows the accumulated discharge volume for the two discharge intervals. The two graphs coincide in the discharge volume. When no backward pressure exists, the discharge volume does not depend on the discharge interval over 5 min. The maximum deviation of the discharge volume is 2.5 and 2.3 nL for 5 and 10 min of the discharge interval, respectively. We observed the discharge volume of the fabricated micropump for various backward pressures ranging from 0 to 15 cm
Fig. 9. The accumulated discharge volume.
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
H2 O. In this test the discharge time, the discharge interval and the input voltage are 4 s, 10 min and 3.5 V, respectively. Fig. 9 shows the accumulated discharge volume for various backward pressures. As the backward pressure increases from 0 to 15 cm H2 O, the average discharge volume decreases gradually from 116 to 87 nL. The standard deviation of the discharge volume is about 2 nL for each backward pressure. Within the backward pressure test range, the discharge volume deviation is less than 15%. 6. Conclusions In this paper, a thermopneumatic constant delivery valveless micropump was proposed and designed with the aid of FEM simulation. The simulation result confirms the perfect operation of the proposed micropump. The micropump was fabricated by micromachining including SU-8 molding and irreversible PDMS bonding. The discharge volume of the micropump was measured for various backward pressure from 0 to 15 cm H2 O when the input voltage is 3.5 V. The micropump discharges 116 nL in 4 s at zero backward pressure when the input voltage is 3.5 V. As the backward pressure increases, the discharge volume decreases gradually. The repeatability of the pump discharge volume is about 2%. This micropump has advantages such as simplicity of the structure and easiness of the fabrication. We conclude that the proposed micropump is feasible for disposable transdermal drug delivery devices. Acknowledgements This research has been supported by the Intelligent Microsystem Center (IMC), which carries out one of the 21st century’s Frontier R&D Projects sponsored by the Korea Ministry of Commerce, Industry and Energy. References [1] D.A.L. Van, T. McGuire, R. Langer, Small-scale systems for in vivo drug delivery, Nat. Biotechnol. 21 (10) (2003) 1184–1191.
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[2] K.P. Kamper, J. Dopper, W. Ehrfeld, S. Oberbeck, A self-filling low-cost membrane micropump, in: Proceedings of the 11th Annual International Workshop on Micro Electro Mechanical Systems, January 1998, pp. 432–437. [3] C. Grosjean, Y.-C. Tai, A thermopneumatic peristaltic micropump, in: Proceedings of the International Conference on Solid-State Sensors and Actuators (Transducers ’99), June 1999, pp. 1776–1779. [4] D. Maillefer, S. Gamper, B. Frehner, P. Balmer, H. Van Lintel, P. Renaud, A high-performance silicon micropump for disposable drug delivery systems, in: Proceedings of the 14th IEEE international conference on MEMS, January 2001, pp. 413–417. [5] S.-J. Kim, Y.B. Shin, D.-S. Lee, H. Yang, K. Kim, S.H. Park, Y.T. Kim, Capillary-driven passive confluence microvalve based on an aspect ratio concept, in: Proceedings of the 6th Korean MEMS Conference, April 2004, pp. 97–100. [6] K.H. Chung, S.-J. Kim, H.B. Pyo, H.-S. Yang, Polymer microfluidic device accomplishing self-wash and fluid replacement via capillary force, in: Proceedings of the 6th Korean MEMS Conference, April 2004, pp. 331– 336. [7] C.W. Hirt, B.D. Nichols, Volume of fluid (VOF) method for the dynamics of free boundaries, J. Comput. Phys. 39 (1) (1981) 201– 225. [8] W.J. Rider, D.B. Kothe, Reconstructing volume tracking, J. Comput. Phys. 141 (2) (1998) 112–152.
Biographies Do Han Jun was born in Korea in 1981. He received his BS and MS degrees in electronics engineering from Ajou University in 2004 and 2006, respectively. He is now a PhD candidate in Ajou University. His current research is concentrated on microelectromechanical devices such as microfluidic devices. Woo Young Sim was born in Korea in 1973. He received his BS and MS degrees in electrical and electronics engineering from Ajou University in 1998 and 2000, respectively. He is pursuing the PhD degree with the Electronics & Computer Engineering Department, Ajou University. His main interests are in the field of Bio-MEMS and Cell-MEMS using microfluidic devices. Sang Sik Yang was born in Korea in 1958. He received his BS and MS degrees in mechanical engineering from Seoul National University in 1980 and 1983, respectively. In 1988, he received his PhD degree in mechanical engineering from University of California, Berkeley. He was then a research assistant professor at New Jersey Institute of Technology. Since 1989, he has been a professor in the School of Electrical Engineering at Ajou University. His research interests include microfluidic devices, microsensors, microactuators and their applications to bio-chips and/or biomedical devices.
A novel constant delivery thermopneumatic micropump using surface tensions Do Han Jun, Woo Young Sim, Sang Sik Yang ∗ School of Electronics Engineering, Ajou University, San 5 Wonchun-Dong, Yeongtong-Gu, Suwon 443-749, Republic of Korea Received 30 July 2006; received in revised form 11 April 2007; accepted 11 April 2007 Available online 19 April 2007
Abstract This paper presents the fabrication and test of a novel thermopneumatic micropump made of PDMS without a membrane or a valve. The driving force of the liquid discharge is the pneumatic pressure of heated air and the suction of liquid is accomplished by the surface tension. The proposed micropump has a pump chamber and a pair of air chambers. The air chambers have electric heaters on the Pyrex glass substrate. Each air chamber is connected with the pump chamber through two different air channels which behaves like check valves. The micropump is fabricated through the spin-coating process, the lithograph process, the PDMS molding process, etc. The total size of micropump is 11.7 mm × 8.8 mm × 0.7 mm. The FEM simulation was performed with air inflow and suction instead of air heating. The fabricated micropump was tested for various backward pressures ranging from 0 to 15 cm H2 O. The simulation and the test results confirmed the expected behavior of the micropump. The discharge volume is 116 nL when the input voltage of 3.5 V is applied for 4 s. This proposed micropump is feasible for disposable transdermal drug delivery system. © 2007 Elsevier B.V. All rights reserved. Keywords: Thermopneumatic; Surface tension; Micropump; Constant delivery
1. Introduction Recently microelectromechanical systems (MEMS) technology has been applied to bio-medical systems. Especially various microfluidic devices have been developed for the application to drug delivery systems (DDS). A drug delivery system releasing accurate amount of drug in time at a desired location can reduce negative side effects due to the excessive dosage that often happens in conventional drug delivery systems. The important objective of a drug delivery system is the enhancement of the treatment effectiveness. It can be achieved by the precise control of the drug delivery amount. Several groups have developed liquid dispensing actuators designed specially for drug delivery systems [1–4]. The drug delivery systems have pump-like structures that can deliver drug continuously or measure the flow rate for the accurate drug delivery. The pumps are miniaturized and have high flow rate. It, however, is difficult to fabricate since the structures are complicated. It is needed to investigate into
∗
Corresponding author. Tel.: +82 331 219 2481; fax: +82 331 212 9531. E-mail address: [email protected] (S.S. Yang).
0924-4247/$ – see front matter © 2007 Elsevier B.V. All rights reserved. doi:10.1016/j.sna.2007.04.038
a micropump so that we can reduce the cost and simplify the structure without losing the accurateness. In this paper, we propose a micropump that discharges a constant amount of liquid by using a bubble expansion by electric heating. We use the surface tension and capillary attraction in the pump chamber and channels to make one-way flow. This micropump has simple structure and is easy to fabricate. To design the shape and the size of the micropump, we perform the simulation by the finite element method (FEM) using CFD-ACE+. We prove the behavior of this micropump by experiments. 2. Micropump structure and design The schematic diagram of the micropump structure is shown in Fig. 1. The micropump consists of an inlet, an outlet, a microchannel, a pump chamber, two air chambers, two discharge air channels and two suction air channels. The air chambers have electric heaters on the Pyrex glass substrate. The pump chamber and the air chambers are connected with the four air channels through which the air is discharged and sucked. Microchannels connect the pump chamber and the inlet and the outlet port. The pump chamber shape is designed to reduce backward
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
surface tension, σ as Eqs. (1) and (2) [5,6]: sin α cos θ cos(θ + β) P = 2σ + α h w
211
(1)
where α is the fluid meniscus angle given by α=θ+β−
Fig. 1. The structure of the micropump.
flow to the inlet and block the inflow from the outlet by surface tensions. The micropump operation sequence is shown in Fig. 2. The one cycle of the operation includes a discharge stroke and a suction stroke. First by the capillary attraction the liquid drug flows into the micropump from inlet as Fig. 2(a). When we apply voltage to the heater, the air in the air chamber expands and pushes the liquid in the pump chamber through air discharge channels as Fig. 2(b). During this discharge stroke the liquid in the pump chamber is discharged to the outlet by the pneumatic pressure as Fig. 2(c). If we turn off the heater, the air contracts and the air is sucked into the air chamber through the air suction channels. During this suction stroke liquid is supplied from the inlet into the pump chamber by the negative pneumatic pressure and the capillary attraction force as Fig. 2(d). At this time, liquid inflow from the outlet is blocked by the surface tension. The shapes of the channels connected to the pump chamber are to be designed appropriately to accomplish the above operation. The liquid flow in the pump chamber is controlled by the surface tension and the capillary attraction. The capillary attraction, P is determined by the channel width, w, the channel height, h, the channel expansion angle, β, the contact angle, θ,
π 2
(2)
If P is positive, the fluid keeps spreading through the expansion channel. If P is negative, the fluid is trapped at the expansion channel. The shape of the inlet-side of the pump chamber is designed to provide enough capillary attraction to achieve selfpriming. In this micropump the expansion angle of the inlet-side of the pump chamber is set to 16◦ so that P is positive. Since the liquid inflow from the pump chamber to the air chamber must be blocked during the suction stroke, we set the expansion angle over 90◦ at the connection of the air discharge channel and the air chamber to make P negative. We set the expansion angle from the outlet channel to the pump chamber to 90◦ to prevent the liquid from flowing from the outlet into the pump chamber during the suction stroke. With this angle design, the expanded air in the pump chamber is sucked into the air chamber through only the air suction channels and the liquid flows from the inlet into the pump chamber until the whole pump chamber is filled. To obtain effective liquid discharge, the air bubbles must grow from the air discharge channels with the air suction channel blocked during the discharge stroke. Since the pneumatic force is proportional to the area of the channel and the capillary attraction force is proportional to the perimeter, we set the width of the air suction channel smaller than that of the air discharge channel so that the air is trapped at the air suction channel during the discharge stroke by the capillary attraction force larger than the pneumatic force. The pump chamber size must be greater than the desired liquid discharge volume in consideration of the dead volume which is figured out from the simulation. We determine the air chamber size enough for the expanded air bubble to push out the liquid by the desired discharge volume during the discharge stroke within the restriction of the moderate temperature rise.
Fig. 2. The operation sequence of the micropump: (a) before discharge stroke, (b) beginning of discharge stroke, (c) end of discharge stroke, and (d) suction stroke.
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Fig. 3. The simulation result: (a) the initial state, (b) air expansion, (c) bubbles merging, (d) drug discharge, and (e) air suction.
The heater resistance is set to 30 . The height of micropump structure is 80 m and the total dimension of the micropump is 11.7 cm × 8.8 cm × 0.7 cm. 3. Simulations We performed the simulation using CFD-ACE+ program for the performance evaluation of the microscale devices and for the detailed size determination. In this simulation we use continuum and incompressible Navier–Strokes equations as the governing equations. Since the flow inside the channel is unsteady and has two phases in this case, we cannot obtain reliable simulation results using only these two equations. We need to use volume of fluid (VOF) method to solve the free surface problem. This method has been proposed by Hirt and Nichols [7]. Recently this method was applied to several fluid cases by Rider and Kothe [8]. In this simulation we use piecewise linear interface construction (PLIC) modeling for reconstructing fluid shape. Also we consider the surface tension in the micropump because surface tension is the important parameter in microscale. Simulation was performed under several assumptions. In the simulation the contact angles of the glass and the PDMS are 10◦ and 100◦ , respectively. The inlet is set to be a wall and the outlet
is set to be open to the atmosphere. The inject rate of air from the air chamber to the pump chamber during the discharge stroke is 9 × 10−11 kg/s and the suction rate of air from the pump chamber to the air chamber during the suction stroke is 9 × 10−11 kg/s. In simulation result, the discharge volume is 124 nL and the discharge time is 2.6 s. Fig. 3 shows the FEM simulation result of the proposed pump. Fig. 3(a) is the initial state. Fig. 3(b) shows the air bubble growing from the air discharge channel and Fig. 3(c) shows the state before the two air bubbles merge in the pump chamber during the discharge stroke. Fig. 3(d) is the end state of the discharge stroke when the fluid is totally discharged from the pump chamber. Fig. 3(e) is the state before the end of the suction stroke and illustrates that the air is sucked into the air chamber with the air discharge channel blocked. The simulation result confirms the complete behavior of the proposed micropump. 4. Fabrication process Fig. 4 shows the fabrication process of the micropump. This micropump comprises three substrates. The upper Pyrex glass substrate has microheaters for electric heating. The middle PDMS substrate has all structures of the micropump except
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
213
Fig. 4. The fabrication process of the micropump.
the microheater. The bottom Pyrex glass substrate is a supporting structure to hold the PDMS structure. The fabrication process consisted of three steps. First for the ˚ layer is deposited and patterned microheater Ti/Au (500/1500 A) on the Pyrex glass substrate as upper layer. The heater resistance is about 30 . Then the inlet and the outlet are made by electrochemical discharge machining (ECDM) process on the glass wafer. The second step is to make SU-8 mold and PDMS layer. For SU-8 mold fabrication, negative thick photoresister (SU-8-
2100) is spun at 500 rpm for 40 s and at 2500 rpm for 60 s on the silicon wafer. The spin-coated photoresister is cured at 65 ◦ C for 25 min and at 90 ◦ C for 50 min. Then the micropump structure is patterned. A 10:1 mixture of PDMS prepolymer and the curing agent is mixed. This prepared PDMS mixture is poured onto the SU-8 mold and cured at 65 ◦ C for 3 h in a vacuum chamber. After curing the PDMS layer is released. Finally the released PDMS layer and Pyrex glass substrates are bonded after the treatment with O2 plasma for 10 s. The photograph of the fabricated micropump is shown in Fig. 5. Fig. 5(a) is a fabricated micropump image and Fig. 5(b) is a SEM image of PDMS layer. 5. Measurements and results In this paper, the high speed video camera is used for the observation of the micropump operations. The discharge time and the self-filling time of the micropump are approximately measured by the frame analysis of the recorded video data. The measurement setup of micropump is shown in Fig. 6. In this test the liquid is water. The discharge time is closely related with the input voltage. As the input voltage increases, the discharge time decreases since the micropump needs a proper input energy to accomplish one discharge cycle. The interval time is independent
Fig. 5. The image of the micropump: (a) photograph of the fabricated micropump and (b) SEM image of PDMS layer.
Fig. 6. The measurement setup.
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Fig. 7. The photograph during the operation: (a) t = 0 s, (b) t = 1.5 s, (c) t = 2.1 s, and (d) t = 4 s.
of the input voltage as far as the input energy is fixed. In this paper, I designed the heater resistance so that the micropump can be operated at the level of TTL output voltage. If the input voltage is low as much as 2 V, it takes over 10 s to discharge. The input voltage selected arbitrarily is 3.5 V which leads to a moderate discharge time less than 5 s. Fig. 7 shows the sequential capture images of the micropump during the operation for the input voltage of 3.5 V. Fig. 7(a)–(d) shows the meniscus of water and the air bubble when t = 0, 1.5, 2.1 and 4 s, respectively. The water and the air bubble movements observed in the experiment are the same as simulation result. The discharge volume of the micropump is calculated from the meniscus movement in the outlet silicone tube observed with
Fig. 8. The accumulated discharged volume when backward pressure is 0 cm H2 O.
a high speed video camera. When applied voltage is 3.5 V and no backward pressure exists, the discharge time is 4 s and selffilling time is 3 min and 52 s. The measured discharge volume is 116 nL. We measured the discharge volume ten times for the repeatability test under zero backward pressure every 5 and 10 min, respectively. Fig. 8 shows the accumulated discharge volume for the two discharge intervals. The two graphs coincide in the discharge volume. When no backward pressure exists, the discharge volume does not depend on the discharge interval over 5 min. The maximum deviation of the discharge volume is 2.5 and 2.3 nL for 5 and 10 min of the discharge interval, respectively. We observed the discharge volume of the fabricated micropump for various backward pressures ranging from 0 to 15 cm
Fig. 9. The accumulated discharge volume.
D.H. Jun et al. / Sensors and Actuators A 139 (2007) 210–215
H2 O. In this test the discharge time, the discharge interval and the input voltage are 4 s, 10 min and 3.5 V, respectively. Fig. 9 shows the accumulated discharge volume for various backward pressures. As the backward pressure increases from 0 to 15 cm H2 O, the average discharge volume decreases gradually from 116 to 87 nL. The standard deviation of the discharge volume is about 2 nL for each backward pressure. Within the backward pressure test range, the discharge volume deviation is less than 15%. 6. Conclusions In this paper, a thermopneumatic constant delivery valveless micropump was proposed and designed with the aid of FEM simulation. The simulation result confirms the perfect operation of the proposed micropump. The micropump was fabricated by micromachining including SU-8 molding and irreversible PDMS bonding. The discharge volume of the micropump was measured for various backward pressure from 0 to 15 cm H2 O when the input voltage is 3.5 V. The micropump discharges 116 nL in 4 s at zero backward pressure when the input voltage is 3.5 V. As the backward pressure increases, the discharge volume decreases gradually. The repeatability of the pump discharge volume is about 2%. This micropump has advantages such as simplicity of the structure and easiness of the fabrication. We conclude that the proposed micropump is feasible for disposable transdermal drug delivery devices. Acknowledgements This research has been supported by the Intelligent Microsystem Center (IMC), which carries out one of the 21st century’s Frontier R&D Projects sponsored by the Korea Ministry of Commerce, Industry and Energy. References [1] D.A.L. Van, T. McGuire, R. Langer, Small-scale systems for in vivo drug delivery, Nat. Biotechnol. 21 (10) (2003) 1184–1191.
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Biographies Do Han Jun was born in Korea in 1981. He received his BS and MS degrees in electronics engineering from Ajou University in 2004 and 2006, respectively. He is now a PhD candidate in Ajou University. His current research is concentrated on microelectromechanical devices such as microfluidic devices. Woo Young Sim was born in Korea in 1973. He received his BS and MS degrees in electrical and electronics engineering from Ajou University in 1998 and 2000, respectively. He is pursuing the PhD degree with the Electronics & Computer Engineering Department, Ajou University. His main interests are in the field of Bio-MEMS and Cell-MEMS using microfluidic devices. Sang Sik Yang was born in Korea in 1958. He received his BS and MS degrees in mechanical engineering from Seoul National University in 1980 and 1983, respectively. In 1988, he received his PhD degree in mechanical engineering from University of California, Berkeley. He was then a research assistant professor at New Jersey Institute of Technology. Since 1989, he has been a professor in the School of Electrical Engineering at Ajou University. His research interests include microfluidic devices, microsensors, microactuators and their applications to bio-chips and/or biomedical devices.