Wireless powered thermo-pneumatic micropump using frequency-controlled heater

Sensors and Actuators A 233 (2015) 1–8

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Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna

Wireless powered thermo-pneumatic micropump using frequency-controlled heater Pei Song Chee a,b , Marwan Nafea Minjal a , Pei Ling Leow a , Mohamed Sultan Mohamed Ali a,c,∗ a

Faculty of Electrical Engineering, Universiti Teknologi Malaysia, 81310 Skudai, Johor, Malaysia Faculty of Engineering and Science, Universiti Tunku Abdul Rahman, 43000 Bandar Sungai Long, Selangor, Malaysia c Flextronics, Pelabuhan Tanjung Pelepas (PTP), 81560 Gelang Patah, Johor, Malaysia b

a r t i c l e

i n f o

Article history: Received 28 January 2015 Received in revised form 9 June 2015 Accepted 17 June 2015 Available online 22 June 2015 Keywords: Wireless actuation Micropump Thermo-pneumatic

a b s t r a c t This paper reports a novel, wirelessly powered micropump based on thermo-pneumatic actuation using a frequency-controlled heater. The micropump operates wirelessly through the energy transfer to a frequency-dependent heater, which was placed underneath the heating chamber of the pump. Heat is generated at the wireless heater when the external magnetic field is tuned to the resonant frequency of the heater. The enclosed air in the chamber expands and forces the liquid to flow out from the reservoir. The developed device is able to pump a total volume of 4 ml in a single stroke when the external field frequency is tuned to the resonant frequency of the heater at the output power of 0.22 W. Multiple strokes pumping are feasible to be performed with the volume variation of ∼2.8% between each stroke. Flow rate performance of the micropump ranges from 1.01 ␮L/min to 5.24 ␮L/min by manipulating the heating power from 0.07 W to 0.89 W. In addition, numerical simulation was performed to study the influence of the heat transfer to the sample liquid. The presented micropump exclusively offers a promising solution in biomedical implantation devices due to its remotely powered functionality, free from bubble trapping and biocompatible feature. © 2015 Elsevier B.V. All rights reserved.

1. Introduction Micropumps have been widely applied in self-contained biomedical and micro-total-analysis system (␮TAS) areas, including transdermal drug delivery and therapeutic implants devices due to their precise dosage controllability. Numerous actuation mechanisms have been developed towards the compatibility of the micropump devices for clinical applications including refractory epilepsy treatment [1], in-vitro diabetics injection [2], fibroblast growth factor (bFGF) controlled for tissue regeneration [3] and pain management [4]. Of other actuation mechanisms such as pneumatic [5–7], piezoelectric [8,9] and electromagnetic principles [10–12], thermo-pneumatic actuation benefits from low operating voltage and no involvement of external peripheral driving equipment setup [13], making it a promising option for clinical biomedical usage. A thermo-pneumatic micropump can be operated through the motion of the membrane’s deflection induced by the heating and cooling of the air due to the temperature gradient of the heating

∗ Corresponding author. http://dx.doi.org/10.1016/j.sna.2015.06.017 0924-4247/© 2015 Elsevier B.V. All rights reserved.

region [14,15]. The pump, in order to control its flow direction, is often coupled with a pair of passive check valve [16] or active valve. Alternatively, a series of multiple actuators arrangement generates a peristaltic motion, which can also convey the flow in a desired direction [17,18]. In these approaches, however, the devices tend to have limited longevity due to the aging and easy tearing of the moving elements such as the thin elastic membrane and flap valves. These constraints deteriorate when multiple actuators are involved. Moreover, they are sensitive to the trapping of bubbles within the moving boundary, and eventually weaken the pumping performance [19]. A straight forward single stroke micropump can potentially address the above-mentioned issues [20]. In this case, the liquid is pushed from a pre-filled reservoir to the outlet microchannel based on the volume expansion of a trapped air upon heating operation. This actuation mechanism eliminates the risk from the mechanical failure of the moving elements and was successfully deployed in portable point-of-care diagnostics [21] and discrete droplet manipulation [22]. Nonetheless, the mobility of the devices was severely restrained by the utilization of a wired and battery interface, which further narrows their application ranges. The ability of wireless power and control of the thermal operation appears to be a key to promote the practicality of the

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Fig. 1. Schematic diagram of a RF powered thermo-pneumatic micropump with (a) device setup with an external coil (including a close up view of the fluid directing channel) (b) a cross section view of the pump, showing frequency-selective working principle using external RF field.

micropump device. This idea can be effectively realized through the wireless radiofrequency (RF) actuation method. The micropump can be wirelessly powered either by converting the RF signal to stress wave propagation [23] or through thermal heating when the field frequency is tuned to the resonant frequency of the heater. The latter principle was demonstrated from our previous work [24] by using field frequency modulation to control the motion of shape memory alloy (SMA) actuators for a microsyringe device. This paper describes the development of a thermo-pneumatic micropump that is powered through RF wireless heating, and controlled by frequency manipulation of external magnetic field. The device was designed with no moving element, i.e., oscillation of passive check valve and membrane, and thus, ensuring a more reliable and durable pumping operation. The wireless flow rate control using the fabricated micropump was carried out as the experimental demonstration of the developed RF thermo-pneumatic actuation principle.

2. Design and working principle The design of the developed micropump (with a total size of 22 × 7 × 4 mm3 ) consists of a heating chamber, a loading reservoir and an outlet hole, which made of PDMS biocompatible material. A microchannel (500 ␮m thickness, 4 mm length) is used to dispense the liquid sample from the loading reservoir (with loading capacity of 9.8 ␮L) to the outlet port. Fig. 1a illustrates the construction design of the developed micropump in this effort. A diffuser, with the opening angle of 12◦ , is formed in the microchannel to create a fluidic resistance between the loading reservoir and the heating chamber to avoid the overflow of the liquid into the chamber. In this work, the pump displaces the liquid in a single direction with no involvement of any mechanical moving membrane and no continuous inlet liquid uptake. Hence, the leakage flow of the diffuser in a pumping operation is eliminated. The detail of the diffuser design parameter was reported in [25]. The layer of fluid directing channel

was enclosed with thick cover lid. The cover lid has the wall thickness of 2 mm depth, aiming to reduce the possible air leakage from the heating chamber to the surrounding. Pumping in the device is achieved by the heat transfer from a planar LC wireless heater (Fig. 1b). The heater serves as a frequency-dependent actuator and is integrated beneath the heating chamber. The LC wireless heater effectively produces Joule heat when exposed to an external RF magnetic field (generated from the external coil) that matches the resonant frequency of the heater, fresonant = √1 [26]. The rate of temperature change is related to 2 LC

the conduction of energy transfer, Q through a PDMS layer. Its relationship is expressed using Fourier’s equation: Q = kA

T x

(1)

where k, A, T and x are respectively denoted as the thermal conductivity of the PDMS, cross sectional surface area between the LC wireless heater and the heating chamber, temperature difference between the LC wireless heater and the heating chamber, and its gap separation distance. From the equation, the rate of conduction heat transfer can be maximized by utilized a smaller gap separation. However, a thinner separation layer might susceptible to deflection and may reduce the pressure to regulate the fluid flow. In our effort, the attached LC wireless heater is made up from a thin polyimide film (∼47 ␮m) which has about 3 orders of magnitude stiffer than PDMS [27], making it rigid to reduce the flexing of the bottom surface of heating chamber. Upon the heat transfer to the chamber, the volume of the air in the heating chamber increases and pushes the liquid from the loading reservoir to the outlet port based on Charles’ law. Charles’s law describes the volume of an ideal air, V, is directly proportional to its absolute temperature, T, given that the pressure, P, and number of moles are held constant: V˛T

(2)

the air is enclosed within the heating chamber through the filling of the liquid sample by using a precision pipette. Hence, the number

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Fig. 2. Simulation of thermo-pneumatic pump (a) temperature profile of the device (b) cross section temperature plot from point A1 to A2.

of air molecules in the chamber is a constant. The volume of the air in the chamber at room temperature, V1 is increased to volume V2 upon the increase of temperature, T, according to the formulation of [13]: V =

V1 T V2

(3)

where, V = V2 − V1 and T = T2 − T1 . To have the change of volume as a driving force to displace the liquid, the pressure at the reservoir must be equal with the pressure at the heating chamber. Therefore, the liquid loading hole was then sealed with silicone adhesive before heating operation. Another alternative method to draw the liquid into the empty reservoir is by heating the chamber to 50 ◦ C for ∼5 min with the loading hole sealed. Then, a sample liquid of 9.8 ␮L (this amount was selected to fill up the entire volume of the reservoir, which is 9.8 ␮L) can be injected at the outlet port. The air contraction upon cooling operation will drag the liquid in, filling the empty space of the reservoir. Excessive volume of sample injection might cause the liquid overflow into the heating chamber, which might result in slow heating process (low pumping rate) due to the volume reduction of the air in the heating chamber. 3. Numerical simulation Numerical simulation of the temperature distribution profile across the thermo-pneumatic actuated micropump is performed using COMSOL Multiphysics® . Excessive heat transfer to the sample solution when dealing with temperature-sensitive material such as protein is one of the main issues in thermo-pneumatic pumps [15]. In addition, for drug delivery application, the increase of the temperature in the reservoir might have adverse chemical effect on the drug. This study, therefore, investigates the heat transfer from the heating chamber to the reservoir, where the sample or solution will be stored. In the simulation model, a steady-state heat transfer module was utilized with the assumption of no heat loss to the surroundings and uniform heating in the LC wireless heater. As depicted in Fig. 2, a heat flux boundary condition was set at the bottom surface of the heating chamber. The heat source boundary domain was set at 51.8 ◦ C (experimentally determined from the LC wireless heater with the heating power of 0.22 W). Zero flux condition was assigned along the side surface. Non-uniform meshes were generated over the model and the initial temperature of the device was fixed at 22 ◦ C (room temperature). A mesh sensitivity test was conducted to ensure the analysis result was independent of the meshing densities. Fig. 2a illustrates the temperature distribution profile of the device while Fig. 2b plots the temperature profile across the micropump along A1 − A2 line. According to the simulation result,

the temperature achieved a maximum value of ∼47 ◦ C at the heating chamber through thermal conduction from the LC wireless heater. The simulated model was then verified with the experimental measurement of the actual device using IR (infrared) thermal camera. As can be seen in Fig. 2b, the experimental measurement showed a narrow temperature profile with lower temperature elevation (∼46 ◦ C) compared to the simulated value (typically varied by ∼1 ◦ C). A possible reason for this variation is that the nonuniform heating profile in the LC wireless heater, where the heat is maximum at the center of the coil. Moreover, the heat loss of the LC wireless heater to the surrounding and the unknown parameters in PDMS thermal behavior (such as thermal diffusion of the PDMS structure), where only affinity value could be gained [28] could be another reason for the discrepancy. Nonetheless, it is worth noting that both results show a close agreement where the temperature eventually decreases to room temperature along the microchannel. This temperature profile suggests that the proposed device did not induce temperature on the working fluid at the reservoir, hence making it a promising device for biomedical application. 4. Device fabrication The RF thermo-pneumatic operated micropump consists of three flexible layers: a PDMS based cover lid, a PDMS fluid directing channel and a polyimide LC wireless heater (described in Fig. 3a). Replication moulding (REM) of PDMS Sylgard 184 (Dow Corning Corp., Midland, MI, USA) was implemented for PDMS layer fabrication. Chemical compounds of resin and hardener were first mixed in a 10:1 ratio before degassing in a vacuum desiccator for 15 min to remove the trapped gas bubble. Subsequently, the mixed

Fig. 3. Fabrication results: (a) view of micropump layers (b) completed micropump device.

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Fig. 4. Wireless setup for micropump characterization.

PDMS was poured into a mould which was fabricated via rapidprototyping process followed by thermal curing procedure at 80 ◦ C for 2 h in an oven. A PDMS cover lid with pre-punched outlet (hole diameter 2.5 mm) and liquid loading holes (hole diameter 1 mm) is then attached to the fluid directing channel patterned PDMS layer using 25 ␮m thick silicone transfer film (ARclad® IS-8026, Adhesive Research Inc., PA, USA) to form a complete pump structure [12]. A planar LC wireless heater which was fabricated on a 45 ␮m thick single sided copper-clad polyimide (PI) film was then adhered onto heating chamber (7 × 5 × 1 mm3 ) using silicone transfer film layer. The details of the LC wireless heater micropatterning process is reported in [26]. Fig. 3b displays the complete fabricated micropump device with a close-up view of the PDMS channel and LC wireless heater. Dye solution was added in the fluid directing channel for better visualization. 5. Experimental setup Wireless testing for the characterization of the fabricated micropump was experimentally performed with the measurement setup ascertained in Fig. 4. An external coil (curled in a diameter of ∼6 mm, 0.213 ␮H inductance) was located at a distance of ∼1.5 mm from the micropump device to generate an external magnetic field that excites the LC wireless heater. The resonant frequency of the LC wireless heater was measured to have 81.5 MHz using network analyzer (Agilent E5071C) by detecting the dips in the S11 parameter. The wireless heater is designed to have 81.5 MHz resonant frequency to accommodate the heating chamber with a size of 7 × 5 mm2 which is similar to the coil size. In this case, theoretical values of the inductance and capacitance of the planar wireless heater are 570 nH and 6.69 pF, respectively, which contributes to the resonant frequency of 81.5 MHz. The resonant frequency can be adjusted by simply changing the size of the capacitor plate of the heater while maintaining the coil size. The thermal behavior and temperature distribution on the PDMS heating chamber were monitored using an infrared thermal camera with the resolution of 320 × 240 (ETI 7320, Infrared Cameras Inc., Texas, USA). Thermal characterization on PDMS surface using infrared imaging has been widely applied in [29,30] as the transmittance of the PDMS material reaches ∼95% in the infrared region (with wavelength range 400–800 ␮m) [31]. In our effort, the experiment was conducted without the inclusion of cover lid to have a clear optical access to the inner wall of the heating chamber. A blue dye liquid was added into reservoir to investigate the volumetric flow rate response of the device. The liquid movement in a microchannel over a period of time was recorded using the camera mounted on a microscope. The image frames were then analyzed with an image processing tool. Similar measurement can also be achieved by detecting the pH change of the released amount of buffer solution [24] or calculating the conductivity of the released molar amount of a saline sample

Fig. 5. Dynamic heat transfer behavior of the LC wireless heater to the heating chamber.

[32]. The ability of the fabricated pump to oppose external pressure was determined by attaching a differential pressure sensor (MPX 10, Motorola) at the pump outlet. The sensor was operated in a differential mode to eliminate the measurement errors which might occur due to the variation in the atmospheric pressure. The output voltage from the sensor was recorded, using a computer. The gas uptake and degas effects of the PDMS based heating chamber (due to the PDMS’s gas permeable properties) is known to influence the pump performance. It was verified that these effects was negligible for the device’s thick structure (>1 mm) and time period (<4 min) involved in the experiment [33]. Hence any leakage within the diffuser in a pumping operation is physically prevented. All the characterizations were repeated five times to take into account the experimental errors that can be caused by non-uniformity of RF heating and environmental factors, including the surrounding temperature and air flow. 6. Results and discussion In this section, the fabricated micropump was characterized in terms of its thermal behavior, and its pumping performance within a microchannel, operated using RF resonant frequency. 6.1. Dynamic heat transfer profile The micropump was operated wirelessly through the heat conduction from the LC wireless heater to the heating chamber. This leads to a temperature gradient across the intersection structure due to the low thermal conductivity of the PDMS material (∼0.15 Wm−1 K−1 [34]) and air (∼0.024 Wm−1 K−1 ), as well as a time delay in reaching the steady state temperature. This spatial and temporal thermal conduction behavior between the LC wireless heater (measured without the inclusion of fluid directing channel and cover lid) and the polymeric heating chamber (measured without the cover lid) were experimentally examined in Fig. 5 when 81.5 MHz RF signal with power of 0.22 W is used. As can be observed, the temperature of LC wireless heater was raised to steady state temperature of 51.8 ◦ C in 20 s, whereas the temperature at the heating chamber reached equilibrium of 46 ◦ C after 100 s. The rate of heat change, calculated from Eq. (1), achieved a maximum rate of 0.08 W at ∼27th second (Fig. 5). At this point, the rate of volume change of the expanded air in the heating chamber attains maximum value. Beyond this threshold point, the rate of heat conduction is decreased until equilibrium condition is reached. In addition, IR images of the micropump in Fig. 5 (obtained from the thermal camera) illustrate that the heating process is initiated from the center of the heating chamber (corresponding to the center of

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Fig. 6. Micropump pumped volume (a) flow transportation of the developed device from the reservoir to the outlet port (b) transient behavior.

the LC wireless heater). The heating pattern is almost similar to LC wireless heater’s copper coil layout, which is in rectangular shape. It is also worth noting that low elevated temperature in the location around the heating chamber assures several micropumps can be operated in close vicinity without causing any crosstalk. 6.2. Wireless thermo-pneumatic actuation characterization After the thermal characterization, the pump performance was evaluated against its temporal response, volumetric flow rate and back pressure. The field frequency used in this test is set to match the resonant frequency of the LC wireless heater that is 81.5 MHz. Fig. 6a illustrates the transitional states of sample dispensed from a reservoir to an outlet port, qualitatively verifying the pumping operation of the developed device in a wireless manner. The flow profile of the device shows a fully developed laminar flow pattern. Hence, several theories such as Hagen–Poiseuille and incompressible Navier–Stoke equations can be implemented in the micropump design to optimize its geometrical parameters [35]. Next, the transient volume pumped out from the device, as well as its liquid filling behavior (liquid volume pumped into the device) is monitored. The rate of temperature change in the heating chamber is also included in Fig. 6b to compare the pumping behavior with temperature under non-steady state condition. It can be observed that the liquid volume pumped out has a similar trend with the temperature changed in the heating chamber. The micropump pumped 0.096 ␮L of liquid at the minimum temperature of ∼26.14 ◦ C (temperature increase of 1.14 ◦ C from room temperature).

The transient characteristic of the liquid suction (pumping in), on the other hand, exhibits an almost linear relationship and reflects a slow response in liquid filling application. The likely cause of this phenomenon is due to the slow cooling rate at the chamber when the RF power is turned off. These results verify that the transient operation of the pump is linearly dependent on the heat induced by the LC wireless heater as Charles’s law predicts. A device test was further carried out to demonstrate the flow rate characteristic of the developed device at different RF power. This was achieved by varying the RF power (range from 0.07 W to 0.89 W) supplied to the external coil. Measurement in Fig. 7a suggests the flow rate of the device improved with higher input RF power. The increase is similar to the trend of temperature increase at different heating power. The RF power is limited to 0.89 W to avoid a further increase in temperature that might damage the device. A maximum flow rate of 5.24 ␮L/min was achieved at 80 ◦ C with the RF power of 0.89 W. In addition, as seen in Fig. 7a, larger volume can be pumped out from the loading reservoir/chamber if higher heating power is used, including pumping out the entire reservoir. However, as mentioned previously, this might cause overheating and damage the device. In light of this aspect, the heating chamber can be designed to have larger volume, where large air expansion could pump out the entire reservoir without causing overheating issue. To validate the time response discussed in Section 6.1, the temporal change in fluid-pumping velocity at 0.22 W RF power is plotted in Fig. 7b. The flow velocity varies with time due to the dynamic heat transfer from the heating chamber. The micropump shows a maximum peak flow velocity at ∼30th second of continu-

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Fig. 7. Flow performance of the developed micropump (a) flow rate characteristic at different RF power (b) flow velocity at RF power 0.22 W.

ous heating under zero back pressure condition. This result agrees well with the transient thermal conduction (discussed in Section 6.1) of the heating chamber, where at this period, a maximum conduction rate is achieved. Beyond ∼30th second, the flow velocity decreases as the rate of heat conduction depletes upon reaching the equilibrium state. In this case, long period of heating at the heating chamber will not contribute to the air expansion and hence, the pressure produced does not have sufficient force to overcome the external pressure. Based on the experimental result, it suggests that the developed device should be operated for <30 s to ensure its optimum condition. To generate continuous pulse flow (similar to the flow characteristic of reciprocating micropump), multiple strokes can be implemented in sequent pumping operation. As a proof of multistroke pumping principle, two consecutive actuation strokes were demonstrated by monitoring the distance of a liquid pumped in the microchannel, presented in Fig. 8. In the first stroke cycle, the RF signal was turned on until the heating chamber’s temperature elevated to 36 ◦ C (the temperature was selected due to the maximum rate of volume change of the air in the heating chamber). The heating chamber’s temperature was left to cool down to 28 ◦ C by turning off the RF signal. Next, similar actuation cycle was repeated for the second stroke. Particularly, the chamber’s temperature was controlled between 28 ◦ C and 36 ◦ C. As shown in Fig. 8, 1.06 ␮L of liquid is pumped out from the reservoir in the first heating cycle (RF turned on from the 10th to 30th seconds). Although the RF signal is deactivated at the 30th seconds, the volume of the pumped liquid continues to increase 1.18 ␮L. This phenomenon is mainly attributed by the slow thermal response from the expanded air. The air requires 20 s to contract and pulls the liquid into the reservoir.

At the second heating cycle (100th second), the liquid is observed to be displaced 1.09 ␮L from the reservoir. Specifically, the first stroke (RF turned on in the 10th second) and second stroke (RF turned on in the 100th second) only differ 2.8%, hence demonstrating its good repeatability in multi-stroke pumping operation. This variation, as shown in Fig. 8, is mainly contributed by the presence of liquid backflow during cooling operation (when the temperature cooled from 36 ◦ C to 28 ◦ C). This issue can be practically addressed through incorporating an active valve in the outlet port of the developed device [36]. As an example of its clinical relevance especially for multi-stroke pumping, continuous subcutaneous delivery of insulin medication in a controlled rate for maintaining diabetics’ sugar level could be a relevant application. In addition, pulsed drug release is shown to lower the pressure required in an intradermal injection [37]. Another example is reported in anabolic treatment of osteoporosis, where human parathyroid hormone fragment (hPTH(1–34)) is released in a pulsatile manner to reduce the bone fractures [38]. It should be noted that for higher dosage of drug medication, the loading capacity of the developed device can be easily enhanced by simply extending the reservoir’s volume, while preserving its miniaturized entity with a lateral size of 22 mm. In addition to the flow characterization of the developed device, the capability to oppose external pressure (back pressure) and the frequency selective activation were demonstrated by attaching a pressure sensor (detail of the experiment setup was discussed in Section 5) to the outlet port while tuning the field from 78.5 MHz to 84.5 MHz. Fig. 9 illustrates the pressure output from the developed thermo-pneumatic micropump when the external field frequency is tuned with constant RF power of 0.22 W. The micropump poses a maximum pressure of 406.5 Pa at the RF resonant frequency of 81.5 MHz. This characteristic matches well with the measured resonant frequency of the LC wireless heater using network analyzer (shown

Fig. 8. Two pumping strokes characteristic of the developed micropump.

Fig. 9. Micropump back pressure with RF frequency variation.

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in Section 5). At resonant frequency, maximum heat is conductively transferred to the heating chamber and thus induces higher expansion force of the trapped air. As the field frequency is tuned away from the resonant frequency of the coil, the developed back pressure reduces significantly. This shows the device operation only responds to suitably tuned electromagentic fields which potentially confirm its functionality as a biomedical tool due to its precision and safety operation. To further enhance its pumping capability, selective activation could be achieved if multiple LC wireless heaters with different resonant frequencies are employed. 7. Conclusion In the present study, a wireless powered thermo-pneumatic actuated micropump using external RF resonant frequency has been developed. The micropump was incorporated with a planar LC wireless heater that was activated through field frequency tuning method. The heater effectively generated heat when the exposed field frequency matched its resonant frequency. The thermal heat produced will be transferred to the heating chamber of the micropump, which was fabricated using MEMs-based technologies, and its temporal response was experimentally characterized. The wireless operation of the fabricated micropump successfully demonstrated a maximum flow rate of ∼2.86 ␮L/min at a chamber temperature of 46 ◦ C with an applied power of 0.22 W, and a maximum pump pressure of 406.5 Pa was obtained. The pump’s flow rate, as denoted in the experiment, can be further enhanced with the utilization of high RF heating levels. In addition, numerical simulation was performed to study the influence of thermal conduction to the sample liquid. The measured data matched well with the simulated result that no elevated temperature was observed in the sample liquid. The developed wireless powered micropump has achieved high controllability performance and can be integrated in application fields that require a remote power source such as a drug delivery system for capsule endoscope [39]. Future improvements of the device involve the optimization of heat transfer efficiency, and a detailed study of pump reliability operation. Future works may increase the heat conduction efficiency of the heating chamber. One approach is to modify the material properties by introducing high thermal conductivity nanocomposite into the PDMS material [40]. Acknowledgements This work was supported by the EScience Fund from Ministry of Science Technology and Innovation Malaysia, Prototype Development Research Grant Scheme (PRGS) and Fundamental Research Grant Scheme (FRGS) from Ministry of Education Malaysia. The authors would like to thank Lum Kin Yun and Then Yi Lung for their assistance and advice in LC wireless heater characterization. We also thank Md. Habibur Rahman and Asnida Abd Wahab for their assistances in high speed microscope and thermal camera usage, respectively. Pei Song Chee acknowledges financial support from Universiti Teknologi Malaysia under postdoctoral research scheme. Marwan Nafea’s doctorate study was supported by the Malaysian Technical Cooperation Programme (MTCP). Appendix A. Supplementary data Supplementary data associated with this article can be found, in the online version, at http://dx.doi.org/10.1016/j.sna.2015.06.017 References [1] M.T. Salam, M. Mirzaei, M.S. Ly, N. Dang Khoa, M. Sawan, An implantable closedloop asynchronous drug delivery system for the treatment of refractory epilepsy, IEEE Trans. Neural Syst. Rehabil. Eng. 20 (2012) 432–442.

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[2] Q. Cui, C. Liu, X. Zha, Study on a piezoelectric micropump for the controlled drug delivery system, Microfluid. Nanofluid. 3 (2007) 377–390. [3] W. Ryu, Z. Huang, F.B. Prinz, S.B. Goodman, R. Fasching, Biodegradable micro-osmotic pump for long-term and controlled release of basic fibroblast growth factor, J. Controlled Release 124 (2007) 98–105. [4] H. Gensler, R. Sheybani, P.-Y. Li, R. Mann, E. Meng, An implantable MEMS micropump system for drug delivery in small animals, Biomed. Microdevices 14 (2012) 483–496. [5] C.-H. Tai, S.-K. Hsiung, C.-Y. Chen, M.-L. Tsai, G.-B. Lee, Automatic microfluidic platform for cell separation and nucleus collection, Biomed. Microdevices 9 (2007) 533–543. [6] C.-H. Wang, G.-B. Lee, Automatic bio-sampling chips integrated with micro-pumps and micro-valves for disease detection, Biosens. Bioelectron. 21 (2005) 419–425. [7] P.A. Willis, B.D. Hunt, V.E. White, M.C. Lee, M. Ikeda, S. Bae, et al., Monolithic Teflon® membrane valves and pumps for harsh chemical and low-temperature use, Lab Chip 7 (2007) 1469–1474. [8] M. Yakut Ali, Cuifang Kuang, Jamil Khan, G. Wang, A dynamic piezoelectric micropumping phenomenon, Microfluid. Nanofluid. 9 (2009) 385–396. [9] B. Ma, S. Liu, Z. Gan, G. Liu, X. Cai, H. Zhang, et al., A PZT insulin pump integrated with a silicon microneedle array for transdermal drug delivery, Microfluid. Nanofluid. 2 (2006) 417–423. [10] C. Yamahata, C. Lotto, E. Al-Assaf, M.A.M. Gijs, A PMMA valveless micropump using electromagnetic actuation, Microfluid. Nanofluid. 1 (2005) 197–207. [11] B. Le Drogoff, L. Clime, T. Veres, The influence of magnetic carrier size on the performance of microfluidic integrated micro-electromagnetic traps, Microfluid. Nanofluid. 5 (2008) 373–381. [12] P.S. Chee, R. Arsat, T. Adam, U. Hashim, R.A. Rahim, P.L. Leow, Modular architecture of a non-contact pinch actuation micropump, Sensors 12 (2012) 12572–12587. [13] K. Lei, K.-H. Chen, Y.-C. Chang, Protein binding reaction enhanced by bi-directional flow driven by on-chip thermopneumatic actuator, Biomed. Microdevices 16 (2014). [14] J.C. Yoo, G.S. La, C.J. Kang, Y.S. Kim, Microfabricated polydimethylsiloxane microfluidic system including micropump and microvalve for integrated biosensor, Curr. Appl. Phys. 8 (2008) 692–695. [15] B.T. Chia, H.-H. Liao, Y.-J. Yang, A novel thermo-pneumatic peristaltic micropump with low temperature elevation on working fluid, Sens. Actuators A: Phys. 165 (2011) 86–93. [16] Y.H. Y.C. Ou, Y.H. Tang, C.C. Lin, N.N. Yang, S.Y. Hsiao Chu, et al., A passive biomimic PDMS valve applied in thermopneumatic micropump for biomicrofluidics, Nano/Micro Engineered and Molecular Systems (NEMS), 2011 IEEE International Conference on 2011 (2011) 325–328. [17] Y.-J. Yang, H.-H. Liao, Development and characterization of thermopneumatic peristaltic micropumps, J. Micromech. Microeng. 19 (2009). [18] S.W. O.C. Jeong, S.S. Park, J.J. Pak Yang, Fabrication of a peristaltic PDMS micropump, Sensors Actuators A: Phys. 123–124 (2005) 453–458. [19] D.J. Laser, J.G. Santiago, A review of micropumps, J. Micromech. Microeng. 14 (2004) R35–R64. [20] W.H. Song, J. Lichtenberg, Thermo-pneumatic, single-stroke micropump, J. Micromech. Microeng. 15 (2005) 1425. [21] D.S. Lee, Y.B. Shin, Y.H. Choi, M.Y. Jung, H.C. Yoon, A self-contained microfluidic platform combining pumps and valves for biochemical applications, Biochip J. 3 (2009) 230–236. [22] K. Handique, D.T. Burke, C.H. Mastrangelo, M.A. Burns, On-chip thermopneumatic pressure for discrete drop pumping, Anal. Chem. 73 (2001) 1831–1838. [23] D.W. Dissanayake, S. Al-Sarawi, T.F. Lu, D. Abbott, Finite element modelling of surface acoustic wave device based corrugated microdiaphragms, Smart Mater. Struct. 18 (2009). [24] M.S. Mohamed Ali, K. Takahata, Wireless microfluidic control with integrated shape-memory-alloy actuators operated by field frequency modulation, J. Micromech. Microeng. 21 (2011) 75005. [25] P.S. Chee, R.A. Rahim, U. Hashim, R. Arsat, P.L. Leow, Low cost diffuser based micropump using pinch actuation, Adv. Mater. Res. 422 (2012) 397–400. [26] M.S. Mohamed Ali, K. Takahata, Frequency-controlled wireless shape-memory-alloy microactuators integrated using an electroplating bonding process, Sens. Actuators A: Phys. 163 (2010) 363–372. [27] G. Liang, G.S. Guvanasen, L. Xi, C. Tuthill, T.R. Nichols, S.P. DeWeerth, A PDMS-based integrated stretchable microelectrode array (isMEA) for neural and muscular surface interfacing, biomedical circuits and systems, IEEE Trans. 7 (2013) 1–10. [28] D. Erickson, D. Sinton, D. Li, Joule heating and heat transfer in poly(dimethylsiloxane) microfluidic systems, Lab Chip 3 (2003) 141–149. [29] P. Yi, A.A. Kayani, A.F. Chrimes, K. Ghorbani, S. Nahavandi, K. Kalantar-zadeh, et al., Thermal analysis of nanofluids in microfluidics using an infrared camera, Lab Chip 12 (2012) 2520–2525. [30] H.-S. Chuang, S. Wereley, Design, fabrication and characterization of a conducting PDMS for microheaters and temperature sensors, J. Micromech. Microeng. 19 (2009) 045010. [31] H. Ren, S.-T. Wu, Introduction to Adaptive Lenses, John Wiley & Sons, New Jersey, 2012.

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P.S. Chee et al. / Sensors and Actuators A 233 (2015) 1–8

[32] J. Fong, Z. Xiao, K. Takahata, Wireless implantable chip with integrated nitinol-based pump for radio-controlled local drug delivery, Lab Chip (2015). [33] M. Ichiyanagi, K. Sakai, S. Kidani, Y. Kakinuma, Y. Sato, K. Hishida, Evaluation methodology of gas permeable characterization in a polymer-based microfluidic device by confocal fluorescence imaging, J. Micromech. Microeng. 22 (2012) 065023. [34] T. Luo, K. Esfarjani, J. Shiomi, A. Henry, G. Chen, Molecular dynamics simulation of thermal energy transport in polydimethylsiloxane, J. Appl. Phys. 109 (2011) 74321–74324. [35] E. Orabona, I. Rea, I. Rendina, L.D. Stefano, Numerical optimization of a microfluidic assisted microarray for the detection of biochemical interactions, Sensors 11 (2011) 9658–9666. [36] S. Rahimi, E.H. Sarraf, G.K. Wong, K. Takahata, Implantable drug delivery device using frequency-controlled wireless hydrogel microvalves, Biomed. Microdevices 13 (2011) 267–277. [37] R. Schoonmaker, M., Bruehwiler, I., Spool, M. Rosen, Multi-stroke delivery pumping mechanism for a drug delivery device for high pressure injections, Google Patents 2010. [38] R. Farra, N.F. Sheppard, L. McCabe, R.M. Neer, J.M. Anderson, J.T. Santini, et al., First-in-human testing of a wirelessly controlled drug delivery microchip, Sci. Transl. Med. 4 (2012) 122ra21. [39] F. Munoz, G. Alici, W. Li, A review of drug delivery systems for capsule endoscopy, Adv. Drug Deliv. Rev. 71 (2014) 77–85. [40] P. Yi, R.A. Awang, W.S.T. Rowe, K. Kalantar-zadeh, K. Khoshmanesh, PDMS nanocomposites for heat transfer enhancement in microfluidic platforms, Lab Chip 14 (2014) 3419–3426.

Biographies

Pei Song Chee received his B.Eng. degree with honours in Control and Instrumentation Engineering from Universiti Teknologi Malaysia (UTM), Malaysia in 2010. He received his Ph.D. in Faculty of Electrical Engineering, Universiti Teknologi Malaysia (UTM) in 2014 and undergoes his postdoctoral training at the same university. Currently, he is an assistant professor with the Faculty of Engineering and Science (FES), Universiti Tunku Abdul Rahman (UTAR). His research interests are lab on chip, microfluidic technology including sensor and actuator integration, microfabrication and finite element modelling.

Marwan Nafea Minjal received the B.Eng. degree in electronics engineering from Sebha University, Brak, Libya, in 2010, and M.Eng. degree in electrical engineering from Universiti Teknologi Malaysia, Skudai, Malaysia, in 2013. He is currently working towards the doctorate degree at University Teknologi Malaysia, Skudai, Malaysia. His research interests are in the areas of MEMS, including piezoelectric actuators and wireless microdevices, with a partial emphasis on the field of particle swarm optimization.

Pei Ling Leow received her BEng in electrical engineering in control and instrumentation in 2003 and her MEng in electrical engineering (mechatronics and automation control) in 2005 from Universiti Teknologi Malaysia. She received her PhD in bioengineering from Imperial College London, UK, in 2009. She is currently lecturing in Universiti Teknologi Malaysia, and her research interests are microfluidic devices, sensors and instrumentation.

Mohamed Sultan Mohamed Ali received the B.Eng. and M.Eng. degrees in electrical engineering from Universiti Teknologi Malaysia, Skudai, Malaysia, in 2006 and 2008, respectively, and the Ph.D. degree in electrical and computer engineering from the Department of Electrical and Computer Engineering, The University of British Columbia, Vancouver, BC, Canada, in 2012. From 2001 to 2007, he held various engineering positions at Flextronics International Ltd. and Jabil Circuit, Inc. He is currently a Senior Lecturer with the Faculty of Electrical Engineering, Universiti Teknologi Malaysia. Dr. Mohamed Sultan also serving as senior consultant engineer for Flextronics (PTP) Malaysia in PCBA failure analysis division. His research interests are in the areas of MEMS, nanotechnology, and micro/nanofabrication technologies, including wireless microdevices, integration of microstructures, and microrobotics.