- Email: [email protected]
Piezoelectric peristaltic micropump integrated on a microfluidic chip
Sensors and Actuators A 292 (2019) 90–96
Contents lists available at ScienceDirect
Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna
Piezoelectric peristaltic micropump integrated on a microfluidic chip Tuo Ma, Shixin Sun, Baoqing Li ∗ , Jiaru Chu Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China, Hefei, Anhui, 230027, China
a r t i c l e
i n f o
Article history: Received 25 January 2019 Received in revised form 7 March 2019 Accepted 4 April 2019 Available online 8 April 2019 Keywords: Micropump Microfluidic chip Micro mixing
a b s t r a c t Micropump is extremely useful in miniature biochemical analysis systems for its abilities of manipulating and transporting minute liquid. We developed a novel peristaltic micropump to pump fluids in microfluidic chip in this article. The micropump consists of three piezoelectric actuators, and part of a microfluidic chip as a compressible microchannel. The fluid in the microchannel is transported owning to a series of impacting actions on the microchannel by the piezoelectric actuators. The microchannel is integrated fabricated with the microfluidic chip, which ensures a tiny dead volume less than 300 nL. Being beneficial from the plug-and-play design of the microfluidic chip, there is no cross-contamination for the reagent has no contact with the reusable actuators. Calibration shows the micropump can transport liquid at a high resolution of 8.4 nL/s, a maximum flow rate of 102 nL/s, and backpressure of up to 2.0 K Pa. A demonstration of mixing liquids on a chip was carried out using two micropumps, indicating its great potential for manipulating microamount reagent in biochemical application. © 2019 Elsevier B.V. All rights reserved.
1. Introduction Micropump plays an essential role as a driving source for the microfluidic chip due to its ability of precisely manipulating small volume of liquid in the closed microchannel compactly [1–3]. Various micropumps have been widely used in the applications such as point-of-care (POC) testing [4], drug delivery [5,6] and microfluidic analysis [7]. According to its working principle, the micropump can be categorized into mechanical micropump and non-mechanical micropump [1,8,9]. Among the widely-used mechanical micropump, a kind of integrated peristaltic micropumps, which integrates the flexible microchannel in the microfluidic chip, has attracted most interest of researchers in manipulating micro liquid for its characteristics of high precision and low cost. In the 1990s, Jan G Smits et al [10] firstly designed the peristaltic micropump, a silicon micropump provided with piezoelectric membrane actuators, which can be manufactured by standard lithography technology. This novel micropump showed good potential for the integrated peristaltic micropump in bioapplication. However, the fabricating complexity of the piezoelectric membrane actuators greatly restricted its application. Later, Steven Quake’s group [11,12] designed a pneumatic peristaltic micropump which contained pneumatic actuators and pumps entirely out of elastomer, to form a microfluidic large-scale integration
∗ Corresponding author. E-mail address: [email protected] (B. Li). https://doi.org/10.1016/j.sna.2019.04.005 0924-4247/© 2019 Elsevier B.V. All rights reserved.
conception. Using pneumatic actuator as the moving part of the micropump improved the process compatibility during chip fabrication. After that, Albert Folch and others [13,14] designed a new pneumatic peristaltic micropump by using 3D printing technology, which will substantially lower the cost of microfluidic devices and provide a new microfluidic large-scale integration solution. And the design of integrated microchannel in chip could help the peristaltic micropump to reduce dead-volume as well as increase working frequency, as compared to the traditional peristaltic pump. However, this kind of pneumatic pump requires external bulky pneumatic equipment and has complex fabricating process. In the application of “lab on a chip”, miniaturization and large-scale integration, as well as the cost, power consumption and fabricating complexity reduction, are becoming the critical elements of micropump [12,14,15]. Recently some researchers have tried to separate the compressible microchannel from the moving part for the purpose of avoiding cross-contamination, which is especially important in biomedical applications. For example, Huang’s group [16] demonstrated a valve-less microfluidic peristaltic pumping method. Liquid was pumped by squeezing the microchannels embedded in a poly device with rolling cams or bearing. And Xiang’s group [17] designed a linear peristaltic micropump which synchronously compressed the microfluidic channel with a miniature cam-follower system. Leow’s group [18] reported a bidirectional micropump using electromagnetic actuator and separated microchannel. The design strategy of separating the actuators or moving parts from microchannel ensures simple fabricating process and low using
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
91
Table 1 Comparison of peristaltic micropump. Author and year
Driver
Integrated /separated actuators
Frequency (Hz) (of single actuator)
Pmax(Kpa)
Qmax(uL/min)
Jan G Smits 1990 [10] Steven Quake 2000 [11] Jun Xie 2004 [19] Meng Shen 2011 [20] Bonnie Tingting Chia 2011 [21] Chungshao Chao 2011 [22] Huawei Yu 2015 [8] Yanyi Huang 2015 [16] Anthony K. Au 2015 [13] DeVoe Don 2016 [23] Jiwen Xiang 2016 [17] Yuansheng Lee 2018 [14] Our work
Piezoelectric Pneumatic Electrostatic Rotating magnetic Thermo pneumatic Piezoelectric Rotating magnetic Motor Pneumatic Motor Rotating micro-cam Pneumatic Piezoelectric
Integrated Integrated Integrated Integrated Integrated Integrated Separated Separated Separated Separated Separated Separated Separated
15 75 20 12 1.2 40 400 ˜ 0.3
5.9 40 1.6 6.6 0.49 / 1.67 340 / 100 36 0.15 2.0
100 0.141 0.0017 40 20.01 0.6 342.4 300 11.333 40 274 21.6 6
6.7 500 ˜ 2.8 20 100
the potential of the low-cost and high-precision micropump for biochemical analysis. 2. Methods and materials 2.1. Working principle
Fig. 1. Concept of the micropump integrated on a microfluidic chip.
cost of the peristaltic micropump. A comparison of the recently reported micropumps is listed in Table 1, showing a trend of designing micropump with separated actuators instead of integrated ones. On the other hand, it is also a challenging issue in the design of the peristaltic micropump to improve the working frequency, which is directly related to the pulsing nature of peristaltic pump. In this paper, we developed a piezoelectric peristaltic micropump which was integrated on a plug-and-play microfluidic chip. And part of the microfluidic chip, which is a section of flexible microchannel, composes the micropump as pump module along with the separated piezoelectric actuators, as shown in Fig. 1. The rest of the microfluidic chip can be used as other functional modular part. Here a typical serpentine channel representing a microfluidic mixer module is presented in the conceptual illustration. This integrated design makes the micropump has the smallest dead volume and fast response time without cross-contamination, as well as high resolution. Both the pump module and other microfluidic module were one-time fabricated. To achieve a high working frequency, we designed the actuators with piezoelectric cantilevers. Due to the small size of microchannel and utilization of piezoelectric actuators, the working frequency can reach up to 100 Hz, which would effectively reduce the inherent pulsation of peristaltic pump. It is worth noting that this is not the maximum frequency of the piezoelectric actuator, just limited by the material of the microfluidic module. Furthermore, being different from other fixed actuators driven by high-speed electric motor [8,23], each piezoelectric actuator can be programmable controlled by microcontroller, providing real-time adjustable travelling wave applied to the peristaltic pump. The small dead volume, real-time control and programmable of each step, enable it suitable for biochemical microfluidic applications. An on-chip flow mixing experiment was carried out using two integrated micropumps, demonstrating
A prototype of the peristaltic micropump is shown in Fig. 2a, which consists of a microchannel and three piezoelectric actuators, the microchannel has a flexible PDMS (polydimethylsiloxane) membrane as the top layer. Three enlarged chambers are specially designed in alignment with the actuators, as shown in Fig. 2b. When the three piezoelectric actuators are applied voltage signals, the actuators generate impacting actions. It’s the impacting actions that compress the chambers, and then squeeze the liquid out of the chamber into microchannel in two directions. With the three pins impacting on the chambers in sequence as Fig. 2c, liquid is transported in a designed direction. A demo cycle with six steps is listed as Fig. 2d. At the beginning three piezoelectric actuators are driven in high-level voltages and the three cantilevers bend toward the chip, enabling the pins to compress the chambers. Next, the first actuator is driven in low-lever voltage, making the first pin and the flexible membrane retract. Liquid flows into the chamber due to the generated negative pressure. In the rest steps the three actuators are driven at voltages as shown in Fig. 2c, and the three pins stroke the membrane in sequence, which transports the liquid from left to right. We assume that “1” stands for a pin impacting on the chamber, and “0” for non-impacting on the chamber, thus a peristaltic period can be expressed as 111-011-001-101-100-110, as shown in the Fig. 2d. 2.2. Calculation model of the transported volume Liquid in the microchannel shows a displaced flow when the flexible membrane is deformed by the pin (with volume Vp ). As shown in Fig. 3a, instead of a dynamic model of transported volume [24], we supply a simplified fluidic electric model can be used to calculate the transported volume from the outlet of each action (V´i ) as [25] Vi’ =
Ri1 Vp , i = 1, 2, . . ., 6. Ri1 + Ri2
(1)
Ri1 (Ri2 ) is the flow resistance of the microchannel, which is in analogue with Ohm’s resistance of electricity [26,27], is given by R≈
1 12l . wh3 1 − 0.63 h w
(2)
92
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
Fig. 2. a) The picture of the integrated peristaltic micropump and the microfluidic chip; b)the design of the microchannel of the peristaltic micropump; c) electric waveform of each actuator; d) and the pump sequence used for the piezoelectric impacting actuators.
Fig. 3. Characteristics of the micropump: a) flow rates at varied voltages of actuators; b) flow rates at varied working frequency of actuators; c) relationship between flow rate and backpressure.
where, w, l, and h are the width, length and height of the microchannel, is the viscosity coefficient of liquid, and R is the flow resistance of the microchannel, respectively. It’s worth noting that the flow resistance is changed when the acting actuator is at different positions (P1 , P2 , and P3 ). Two mainly factors induce the change, which can be seen from Fig. 3b. One is the geometrical changes in the length of channels, and the other is the changes in height of the three circular chambers caused by pin’s squeeze. The total transported volume can be estimated by a sum of all six movements. In each movement, only one actuator is designed to change its status. We use V‘1 , V‘2 , V‘3 , V‘4 , V‘5 and V‘6 representing the transported volumes of each action out of the outlet, respectively. Then the volume of total transportation V’ in one period is V ’ = V1’ + V2’ + V3’ + V4’ + V5’ + V6’
(3)
Substitute Eq. 2 into Eq. 1, then into Eq. 3, the transported volume in one period can be expressed as
V ’ ≈ f W, w, l1 , l2 , l3 , lp , lr , h · s.
(4)
Details of the derivation are given in the supplementary material from Eq.S11 to Eq.S12. In Eq. 4, l1 , l2 , l3 , l4 are the lengths of different sections, as is shown in Fig. 3b. lp is the equivalent length of the compressed chamber, and lr is the equivalent length of uncompressed chamber. W and w are the widths at different sections, and h is the height of the microchannel, s is the deformation of flexible membrane, which is equal to the bending displacement of the piezoelectric actuators. And the displacement is related to the driving voltage of the piezoelectric actuator. In brief, the transported
liquid volume can be directly controlled by regulating the driving voltages as well as changing of channel sizes. 2.3. Micropump system To squeeze the liquid in the microchannel, three piezoelectric cantilevers (QDTE52, PANT Corp, China) were adopted with three micro pins (1356-1 & 1356-3, Keystone Electronics Corp, USA) adhered at the ends of cantilevers by epoxy adhesive (467 M P, 3 M Corp, USA), respectively. The microfluidic chip was fabricated using the standard soft lithography process [11]. Firstly, a mold was made using SU-8 (SU-8 2025, MicroChemicals Corp, Germany) material by UV lithography. Secondly, the PDMS (Sylgard 184, Dow-Corning Corp, USA) mixed at a ratio of 1:10 (curing agent to base) was casted onto the SU-8 master mold to form microfluidic channel. Then the solidified PDMS with microchannel was bonded to a glass slide by oxygen plasma (PDC-MG, Mingheng Corp, China). The width w, the height h of the channel and the total thickness of the PDMS layer were 355.7 m, 54 m and 355 m, respectively. l3 was 2 mm, and the total length (from inlet to outlet) was 35 mm. Diameters of the pin and the chamber were 0.6 mm and 1.0 mm, respectively. 2.4. Calibration methods To characterize the micropump, flow rate was calibrated by measuring the transported liquid volume through the micro channel during pump working. Here we measured flow volume by monitoring the liquid levels in the glass tubes, which were connected to the inlet and outlet (Fig. 2a), respectively. The inner
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
93
Fig. 4. Characteristics of the micropump: a) flow rates at varied driving voltages; b) flow rates at varied working frequency of actuators; c) relationship between flow rate and backpressure.
diameter of the tube is 2.90 mm. The operating process is as follows: Firstly, DMSO (Dimethyl Sulphoxide) with 5% (w/w) blue dye (GT52, HP Corp, USA) was injected from the inlet tube to the outlet tube, to achieve a same liquid level. By measuring the real-time change of liquid level during working, we can get the flow rate of the micropump. At this moment, the piezoelectric cantilevers were driven by programmable pulse-width-modulation (PWM) square waveforms (range: DC 60 V–180 V), which were generated by an FPGA controller circuitry. The backpressure can also be calibrated through calculating the pressure difference, which was related to the height difference in the two tubes using the same setup, except that a longer outlet tube was required. A high-speed camera (Chronos 1.4, Chronos Corp, Canada) was used to capture the mixing process on a chip. 3. Results and discussion 3.1. Characteristics of the micropump Owing to small size of the microchannel, the dead volume is less than 300 nL. It means that when the micropump is integrated on a microfluidic chip, there is only microamount reagent wasted, which is particularly important in using expensive biological reagents. The measured process of the stop of the micropump using high speed camera shows a waiting time less than 12.5 ms (Fig.S2), indicating a fast response time in the scale of millisecond. A series of calibration experiments were carried out to inspect the characteristics of the peristaltic micropump. As shown in Fig. 4a, the flow rate is monotonously increased with the driving voltage. According to Eq. 4, the transported volume has positive correlation with the displacement of the actuator, which is proportional to the driving voltage. When the voltage is 90 V and the working frequency is 25 Hz, the flow rate is 8.4 nL/s. Meanwhile, the minimum repeatable volume of the transported liquid in one cycle is only about 2.0 nL. This volume resolution can be further improved by using smaller driving voltage or longer channel. Actually, we have ever achieved flow rate resolution of 0.08 nL/s with length (l3 ) of 8 mm and driving voltage of 150 V (Fig.S3). On the other hand, the flow rate is also increased with the driving frequency in the range from 0 to 100 Hz. And the relationship between the flow rate and the frequency is almost linear when driving voltage is larger than 90 V. Since higher frequency means more working cycles within unit time, the micropump can transport more volume of liquid in same time. In Fig. 4b, the measured flow rate at 180 V (red solid line) is very close to the theoretical result (black solid line), especially at low frequency (<50 Hz). The theoretical result is drawn from Eq. 4, representing a maximum transported rate. At the moment, the flexible membrane on the chamber got a maximum deformation, which is equal to the height
of the microchannel. When the driving frequency is larger than 75 Hz, the error between the theoretical and experimental results is obvious because the theoretical result is based on a static model without considering the inertia of the liquid. Peristaltic micropump has a special capability of transporting liquid bi-directionally. By adjusting the impacting sequence in controlling program, our integrated micropump can precisely change flow direction at any working step. Here, we calibrated the flow rates at two directions at 150 V using the chip presented in Fig. 2c. As shown in Fig. 4b, when frequency is less than 100 Hz, the flow rates of both forward and inverse directions are almost equal. They are both increased with frequency. The differences between the two directions at larger frequency could mainly be caused by the installation error, which makes the initial gaps between the three pins and the membrane unequal in vertical position. Observed from the experiments, we found that the influence of the installation error on the symmetry of the bidirectional flow rates was enhanced with the frequency and voltage. Thus, a careful alignment between the pump chip and the three actuators is very essential and important. At last, we studied the flow rate change at the different back pressures. The maximum back pressure can reach 2.0 K Pa, as shown in Fig. 4c. 3.2. Application of the peristaltic micropump: mixing on a chip In biochemical applications, titrations are performed not only for a given compound but also for a mixture of two or more compounds, including their titration [28]. The mixing of biological or chemical species can be applied to digital polymer chain reaction [7,29], biochemical analyses [30], nanomaterial synthesis [31], or emulsification [32]. In order to achieve effective, low consuming, rapid and complete mixing, various types of micro-mixers have been created to improve the mixing in microscale. The enhancement of mixing is usually achieved by increasing the mixing area [33]. In our demo, we conducted a mixing process based on a microfluidic two-phase mixing chip integrated with the peristaltic micropump. Two colored liquids were efficiently mixed at both low and high flow rates. As shown in Fig. 5a, there are two microfluidic pump modules, pump one is for yellow dye, and pump two for blue dye. Two groups of piezoelectric actuators were adopted, forming two peristaltic micropumps together with the two pump modules (Fig. 5b). A zigzag channel with length lz is 10 mm, applied as mixing module to assist mixing for its ability of improving the mixing efficiency [34]. The microfluidic two-phase mixing module was fabricated in one step with the integrated pump modules. The width of inlet of “Y” junction wi is 200 um, and its width of outlet wo is 500 um, the height of microchannel is 54 um. It was easy to generate the initial laminar flow by making the two micropumps work simul-
94
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
Fig. 5. a) A photo of a microfluidic chip showing the pumping area and zigzag mixing area; b) photo of the microfluidic mixing setup.
Fig. 6. The mixing process by two peristaltic pumps showing a) the initial state when both fluids were laminar flow at the frequency of 100 Hz and voltage of 150 V. When the frequency was reduced to 10 Hz, pulsation was obviously as follows: b) the mixed fluids was flowing back to the yellow colored channel; c) the mixed fluids was flowing back to the blue colored channel; d) the liquids in the channel after fully mixed by alternative transportation; e) the mixing effects at position A and position B (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article).
taneously, as shown in Fig. 6a. There was only a stable diffusion layer existing between two colored layers where the inefficient mixing was completely realized by molecular diffusion. In this situation, the flow rate was about 83 nL/s for each micropump when the working frequency and applied voltage were 100 Hz and 150 V, respectively.
3.2.1. Mixing at low frequency As mentioned in the section of working principle, fluid is squeezed out of the chamber and flows in both directions when the pin impacts on the membrane. On the other hand, fluid is also sucked back into the chamber in both directions when the pin reverts. During the six-step cycle, the fluid in the pump chip moves back and forth alternately, and lastly moves forward due to the impacting sequence. Usually the pulsation of the peristaltic pump is disadvantageous in the application where a stable flow is required. Nevertheless, mixing using microfluidic chip is benefited from pulsation or disturbance flow, since the fluid is always laminar flow in the microchannel. This pulsation of the designed peristaltic micropump was especially obvious in the experiment at low driving frequency (e.g. 10 Hz). When the work sequences of the two pumps weren’t completely synchronous, the liquid in the Y junction moved back and forth into the two pump channels alternately, generating a turbulent flow, as shown in Fig. 6b & c. Furthermore, due to the low flow rate (5 nl/s) at low driving frequency (10 Hz), fluids were mixed well just behind their confluence (Fig. 6d). A movie in the Supplementary Files (Video-1) shows the mixing process. Here, we describe the mixing effect using relative mixing index (RMI) [35,36], where RMI = 1 and RMI = 0 indicate complete separation and complete mixing, respectively. By calculating RMI, the mixing level is quantitatively evaluated by substituting a standard
deviation, which is a light intensity distribution in a specified region inside a channel as
RMI =
= 0
N 1
(Ii − Im )2
N
i=1
N
(I − Im )2 i=1 oi
N 1
.
(5)
Where, is the standard deviation of the pixel intensities, o is the standard deviation of the pixel intensities in the unmixed case, Ii shows a local pixel intensity, I0i represents the local pixel intensity in the unmixed state, Im shows the average of the pixel intensities in the cross section, N represents the total number of pixels. As shown in Fig. 5e, at position A, turbulent mixing began immediately after the two pumps worked at low-frequency mode. And it became stable in about 10 s when the RMI was 0.278. After another 40 s, these fluids were further mixed with RMS of 0.188 at position B when flowing through the zigzag channel. It can be seen that the peristaltic micropumps can mix the fluids in a short and straight channel quickly, even without the zigzag channel. However, the zigzag channel can help improve the mixing effect. 3.2.2. Mixing at high frequency The pulsation of the peristaltic pump was gradually weakened with the driving frequency increasing. And this attenuating pulsation showed a regressive effect on the aforementioned mixing. For example, when the frequency was 100 Hz, the fluids in the microfluidic chip returned to laminar flow. For mixing in microchannel, some researchers used specially-designed microchannel configurations [33,37,38] to enhance mixing effect. Here, we proposed a new mixing method by generating an arbitrary pulsation at high frequency, without specially-designed microchannel configurations. Being beneficial from the peristaltic pump’s characteristic of bidirectional transportation, we made one pump intermittently
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
95
Fig. 7. The mixing process by two peristaltic pumps at high working frequency: a) pulsation waveform of the flow rate of pump one, the negative and positive value of flow rate represent the direction of transportation; b) pulsation waveform of the flow rate of pump two; c)-e) pump one transports the yellow ink in pulsation and pump two transports liquid continuously, the red arrow shows the direction of the flow; f) the liquids after fully mixing by alternative transporting ; g) the mixing effects at position A and position B. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article).
transport liquid in inverse direction during the normal forward transportation, while the other pump still worked in constant forward direction. A movie in the Supplementary Files (Video-2) shows the mixing process. As shown in Fig. 7a, the liquid was transported in the positive direction in the time of t1 , and then transported reversely in the time of t2 . Here, t1 and t2 are 3.6 s and 0.36 s, respectively. Due to the small size (9 mm) of pump chip, the peristaltic pump can change transporting direction during working in milliseconds by just importing driving signals in different sequence. As can be seen from Fig. 7c-e, the reciprocating flow of fluids by pump one generates a similar artificial turbulent flow as that at low frequency near the Y junction. After 5 periods, the mixing effect was getting better gradually as Fig. 7f. The relative mixing index (RMI) was calculated to describe the mixing effect. As is shown in Fig. 7g, at position A, the reciprocating flow generated by pump one shows a positive effect on the mixing. It decreased the RMI slowly. With the combined effect of the zigzag microchannel, the mixing efficiency was further improved, i.e. the RMI at the position B was reduced from 0.24 to 0.065. Compared to native pulsation of the peristaltic pump at the low-frequency mixing, the new arbitrary pulsation at high frequency shows a better mixing effect, more flexible strategy and larger mixing efficiency. 4. Conclusion In this paper, we presented a peristaltic micropump that was composed of separated microfluidic channel and piezoelectric actuators. The unique design of integrating the channel with other functional microfluidic modules makes it suitable for building a plug-and-play, conceptual pump-on-a-chip. The peristaltic micropump has been successfully designed, fabricated, and calibrated, showing a high resolution of 8.4 nL/s, and dead volume less than of 300 nL. In addition, a demonstration of mixing in one chip using two peristaltic micropumps was carried out. Two working modes at both low and high driving frequencies were presented, which realized an effective mixing in a microfluidic chip with only simple-structure channel. The results of mixing have demonstrated
the excellent performances and high reliability of the micropump, showing great potential for applications such as point-of-care (POC) testing, drug delivery, and chemical analysis. Acknowledgements This research work was supported in part by the National Natural Science Foundation of China (No. 51675505), the Joint Research Fund for Overseas Chinese, Hong Kong and Macao Young Scientists of the National Natural Science Foundation of China (No. 51628502). Appendix A. Supplementary data Supplementary material related to this article can be found, in the online version, at doi:https://doi.org/10.1016/j.sna.2019.04. 005. References [1] D.J. Laser, J.G. Santiago, A review of micropumps, J. Micromech. Microeng. 14 (2004) R35. [2] K.W. Oh, C.H. Ahn, A review of microvalves, J. Micromech. Microeng. 16 (2006) R13. [3] A. Pourmand, S.A.M. Shaegh, H.B. Ghavifekr, E.N. Aghdam, M.R. Dokmeci, A. Khademhosseini, Y.S. Zhang, Fabrication of whole-thermoplastic normally closed microvalve, micro check valve, and micropump, Sens. Actuators B Chem. 262 (2018) 625–636. [4] C.D. Chin, V. Linder, S.K. Sia, Commercialization of microfluidic point-of-care diagnostic devices, Lab Chip 12 (2012) 2118–2134. [5] A. Nisar, N. Afzulpurkar, B. Mahaisavariya, A. Tuantranont, MEMS-based micropumps in drug delivery and biomedical applications, Sens. Actuators B Chem. 130 (2008) 917–942. [6] S. Herrlich, S. Spieth, S. Messner, R. Zengerle, Osmotic micropumps for drug delivery, Adv. Drug Deliv. Rev. 64 (2012) 1617–1627. [7] C. Zhang, D. Xing, Y. Li, Micropumps, microvalves, and micromixers within PCR microfluidic chips: advances and trends, Biotechnol. Adv. 25 (2007) 483–514. [8] H. Yu, W. Ye, W. Zhang, Z. Yue, G. Liu, Design, fabrication, and characterization of a valveless magnetic travelling-wave micropump, J. Micromechanics Microengineering 25 (2015), 065019.
96
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
[9] Y.N. Wang, L.M. Fu, Micropumps and biomedical applications–A review, Microelectron. Eng. 195 (2018) 121–138. [10] J.G. Smits, Piezoelectric micropump with three valves working peristaltically, Sens. Actuators A Phys. 21 (1990) 203–206. [11] M.A. Unger, H.P. Chou, T. Thorsen, A. Scherer, S.R. Quake, Monolithic microfabricated valves and pumps by multilayer soft lithography, Science 288 (2000) 113–116. [12] T. Thorsen, S.J. Maerkl, S.R. Quake, Microfluidic large-scale integration, Science 298 (2002) 580–584. [13] A.K. Au, N. Bhattacharjee, L.F. Horowitz, T.C. Chang, A. Folch, 3D-printed microfluidic automation, Lab Chip 15 (2015) 1934–1941. [14] Y.S. Lee, N. Bhattacharjee, A. Folch, 3D-printed quake-style microvalves and micropumps, Lab Chip 18 (2018) 1207–1214. [15] H. Landari, M.A. Dussault, J. Ruel, A. Begin-Drolet, A. Miled, Biocompatible compact micropump with integrated unidirectional microvalves for low pressure microfluidic applications, Sens. Actuators A Phys. 276 (2018) 246–258. [16] X. Zhang, Z. Chen, Y. Huang, A valve-less microfluidic peristaltic pumping method, Biomicrofluidics 9 (2015), 014118. [17] J. Xiang, Z. Cai, Y. Zhang, W. Wang, A micro-cam actuated linear peristaltic pump for microfluidic applications, Sens. Actuators A Phys. 251 (2016) 20–25. [18] M. Rusli, P.S. Chee, R. Arsat, K.X. Lau, P.L. Leow, Electromagnetic actuation dual-chamber bidirectional flow micropump, Sens. Actuators A Phys. 282 (2018) 17–27. [19] J. Xie, J. Shih, Q. Lin, B. Yang, Y.C. Tai, Surface micromachined electrostatically actuated micro peristaltic pump, Lab Chip 4 (2004) 495–501. [20] M. Shen, L. Dovat, M.A. Gijs, Magnetic active-valve micropump actuated by a rotating magnetic assembly, Sens. Actuators B Chem. 154 (2011) 52–58. [21] 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. [22] C.S. Chao, P.C. Huang, M.K. Chen, L.S. Jang, Design and analysis of charge-recovery driving circuits for portable peristaltic micropumps with piezoelectric actuators, Sens. Actuators A Phys. 168 (2011) 313–319. [23] J. Han, O. Rahmanian, E. Kendall, N. Fleming, D. DeVoe, Screw-actuated displacement micropumps for thermoplastic microfluidics, Lab Chip 16 (2016) 3940–3946. [24] C. Jenke, S. Kager, M. Richter, C. Kutter, Flow rate influencing effects of micropumps, Sens. Actuators A Phys. 276 (2018) 335–345. [25] Y. Ding, E. Huang, K.S. Lam, T. Pan, Microfluidic impact printer with interchangeable cartridges for versatile non-contact multiplexed micropatterning, Lab Chip 13 (2013) 1902–1910. [26] M. Tanyeri, M. Ranka, N. Sittipolkul, C.M. Schroeder, A microfluidic-based hydrodynamic trap: design and implementation, Lab Chip 11 (2011) 1786–1794. [27] A.J. Teo, K.H.H. Li, N.T. Nguyen, W. Guo, N. Heere, H.-D. Xi, C.-W. Tsao, W. Li, S.H. Tan, Negative pressure induced droplet generation in a microfluidic flow-focusing device, Anal. Chem. 89 (2017) 4387–4391. [28] A. Folch, Introduction to bioMEMS, CRC Press2016, Boca Raton, 2016, pp. 191–204.
[29] M. Hashimoto, F. Barany, S.A. Soper, Polymerase chain reaction/ligase detection reaction/hybridization assays using flow-through microfluidic devices for the detection of low-abundant DNA point mutations, Biosens. Bioelectron. 21 (2006) 1915–1923. [30] C. Neils, Z. Tyree, B. Finlayson, A. Folch, Combinatorial mixing of microfluidic streams, Lab Chip 4 (2004) 342–350. [31] D. Ahmed, X. Mao, J. Shi, B.K. Juluri, T.J. Huang, A millisecond micromixer via single-bubble-based acoustic streaming, Lab Chip 9 (2009) 2738–2741. [32] S. Okushima, T. Nisisako, T. Torii, T. Higuchi, Controlled production of monodisperse double emulsions by two-step droplet breakup in microfluidic devices, Langmuir 20 (2004) 9905–9908. [33] C.Y. Lee, C.L. Chang, Y.N. Wang, L.M. Fu, Microfluidic mixing: a review, Int. J. Mol. Sci. 12 (2011) 3263–3287. [34] V. Mengeaud, J. Josserand, H.H. Girault, Mixing processes in a zigzag microchannel: finite element simulations and optical study, Anal. Chem. 74 (2002) 4279–4286. [35] A. Hashmi, J. Xu, On the quantification of mixing in microfluidics, J. Lab. Autom. 19 (2014) 488–491. [36] R. Sakurai, K. Yamamoto, M. Motosuke, Concentration-adjustable micromixer using droplet injection into a microchannel, Analyst 4 (2019) 559–568. [37] C.C. Hong, J.W. Choi, C.H. Ahn, A novel in-plane passive microfluidic mixer with modified Tesla structures, Lab Chip 4 (2004) 109–113. [38] C.Y. Lee, L.M. Fu, Recent advances and applications of micromixers, Sens. Actuators B Chem. 259 (2018) 677–702.
Biographies Tuo Ma received his BS degree in June 2010 from Anhui University. He is currently pursuing the PhD degree at Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China. His research interests include Lab-on-a-chip, integrated microfluidic system and actuators. Shixin Sun received her BS degree in June 2018 from Northeast Agricultural University. She is currently pursuing the MS degree at Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China. Her research interest is Lab-on-a-chip. Baoqing Li received his BEng, MEng and PhD in mechanical engineering and precision instrumentation from University of Science and Technology of China (USTC) in 2000, 2003 and 2008, respectively. He was a research assistant at Hong Kong Polytechnic University in 2004–2006, and a Post-Doctoral Fellow at University of California, Davis in 2013–2015, working on microfluidic technology. Currently he is an associate professor at USTC, and his research interests are Lab-on-a-chip and MEMS. Jiaru Chu received his BEng, MEng and PhD in mechanical engineering and precision instrumentation from University of Science and Technology of China (USTC) in 1985, 1988 and 1997, respectively. He currently is a professor at USTC, and his interests are MEMS, laser fabricating technology and telescope instruments.
Contents lists available at ScienceDirect
Sensors and Actuators A: Physical journal homepage: www.elsevier.com/locate/sna
Piezoelectric peristaltic micropump integrated on a microfluidic chip Tuo Ma, Shixin Sun, Baoqing Li ∗ , Jiaru Chu Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China, Hefei, Anhui, 230027, China
a r t i c l e
i n f o
Article history: Received 25 January 2019 Received in revised form 7 March 2019 Accepted 4 April 2019 Available online 8 April 2019 Keywords: Micropump Microfluidic chip Micro mixing
a b s t r a c t Micropump is extremely useful in miniature biochemical analysis systems for its abilities of manipulating and transporting minute liquid. We developed a novel peristaltic micropump to pump fluids in microfluidic chip in this article. The micropump consists of three piezoelectric actuators, and part of a microfluidic chip as a compressible microchannel. The fluid in the microchannel is transported owning to a series of impacting actions on the microchannel by the piezoelectric actuators. The microchannel is integrated fabricated with the microfluidic chip, which ensures a tiny dead volume less than 300 nL. Being beneficial from the plug-and-play design of the microfluidic chip, there is no cross-contamination for the reagent has no contact with the reusable actuators. Calibration shows the micropump can transport liquid at a high resolution of 8.4 nL/s, a maximum flow rate of 102 nL/s, and backpressure of up to 2.0 K Pa. A demonstration of mixing liquids on a chip was carried out using two micropumps, indicating its great potential for manipulating microamount reagent in biochemical application. © 2019 Elsevier B.V. All rights reserved.
1. Introduction Micropump plays an essential role as a driving source for the microfluidic chip due to its ability of precisely manipulating small volume of liquid in the closed microchannel compactly [1–3]. Various micropumps have been widely used in the applications such as point-of-care (POC) testing [4], drug delivery [5,6] and microfluidic analysis [7]. According to its working principle, the micropump can be categorized into mechanical micropump and non-mechanical micropump [1,8,9]. Among the widely-used mechanical micropump, a kind of integrated peristaltic micropumps, which integrates the flexible microchannel in the microfluidic chip, has attracted most interest of researchers in manipulating micro liquid for its characteristics of high precision and low cost. In the 1990s, Jan G Smits et al [10] firstly designed the peristaltic micropump, a silicon micropump provided with piezoelectric membrane actuators, which can be manufactured by standard lithography technology. This novel micropump showed good potential for the integrated peristaltic micropump in bioapplication. However, the fabricating complexity of the piezoelectric membrane actuators greatly restricted its application. Later, Steven Quake’s group [11,12] designed a pneumatic peristaltic micropump which contained pneumatic actuators and pumps entirely out of elastomer, to form a microfluidic large-scale integration
∗ Corresponding author. E-mail address: [email protected] (B. Li). https://doi.org/10.1016/j.sna.2019.04.005 0924-4247/© 2019 Elsevier B.V. All rights reserved.
conception. Using pneumatic actuator as the moving part of the micropump improved the process compatibility during chip fabrication. After that, Albert Folch and others [13,14] designed a new pneumatic peristaltic micropump by using 3D printing technology, which will substantially lower the cost of microfluidic devices and provide a new microfluidic large-scale integration solution. And the design of integrated microchannel in chip could help the peristaltic micropump to reduce dead-volume as well as increase working frequency, as compared to the traditional peristaltic pump. However, this kind of pneumatic pump requires external bulky pneumatic equipment and has complex fabricating process. In the application of “lab on a chip”, miniaturization and large-scale integration, as well as the cost, power consumption and fabricating complexity reduction, are becoming the critical elements of micropump [12,14,15]. Recently some researchers have tried to separate the compressible microchannel from the moving part for the purpose of avoiding cross-contamination, which is especially important in biomedical applications. For example, Huang’s group [16] demonstrated a valve-less microfluidic peristaltic pumping method. Liquid was pumped by squeezing the microchannels embedded in a poly device with rolling cams or bearing. And Xiang’s group [17] designed a linear peristaltic micropump which synchronously compressed the microfluidic channel with a miniature cam-follower system. Leow’s group [18] reported a bidirectional micropump using electromagnetic actuator and separated microchannel. The design strategy of separating the actuators or moving parts from microchannel ensures simple fabricating process and low using
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
91
Table 1 Comparison of peristaltic micropump. Author and year
Driver
Integrated /separated actuators
Frequency (Hz) (of single actuator)
Pmax(Kpa)
Qmax(uL/min)
Jan G Smits 1990 [10] Steven Quake 2000 [11] Jun Xie 2004 [19] Meng Shen 2011 [20] Bonnie Tingting Chia 2011 [21] Chungshao Chao 2011 [22] Huawei Yu 2015 [8] Yanyi Huang 2015 [16] Anthony K. Au 2015 [13] DeVoe Don 2016 [23] Jiwen Xiang 2016 [17] Yuansheng Lee 2018 [14] Our work
Piezoelectric Pneumatic Electrostatic Rotating magnetic Thermo pneumatic Piezoelectric Rotating magnetic Motor Pneumatic Motor Rotating micro-cam Pneumatic Piezoelectric
Integrated Integrated Integrated Integrated Integrated Integrated Separated Separated Separated Separated Separated Separated Separated
15 75 20 12 1.2 40 400 ˜ 0.3
5.9 40 1.6 6.6 0.49 / 1.67 340 / 100 36 0.15 2.0
100 0.141 0.0017 40 20.01 0.6 342.4 300 11.333 40 274 21.6 6
6.7 500 ˜ 2.8 20 100
the potential of the low-cost and high-precision micropump for biochemical analysis. 2. Methods and materials 2.1. Working principle
Fig. 1. Concept of the micropump integrated on a microfluidic chip.
cost of the peristaltic micropump. A comparison of the recently reported micropumps is listed in Table 1, showing a trend of designing micropump with separated actuators instead of integrated ones. On the other hand, it is also a challenging issue in the design of the peristaltic micropump to improve the working frequency, which is directly related to the pulsing nature of peristaltic pump. In this paper, we developed a piezoelectric peristaltic micropump which was integrated on a plug-and-play microfluidic chip. And part of the microfluidic chip, which is a section of flexible microchannel, composes the micropump as pump module along with the separated piezoelectric actuators, as shown in Fig. 1. The rest of the microfluidic chip can be used as other functional modular part. Here a typical serpentine channel representing a microfluidic mixer module is presented in the conceptual illustration. This integrated design makes the micropump has the smallest dead volume and fast response time without cross-contamination, as well as high resolution. Both the pump module and other microfluidic module were one-time fabricated. To achieve a high working frequency, we designed the actuators with piezoelectric cantilevers. Due to the small size of microchannel and utilization of piezoelectric actuators, the working frequency can reach up to 100 Hz, which would effectively reduce the inherent pulsation of peristaltic pump. It is worth noting that this is not the maximum frequency of the piezoelectric actuator, just limited by the material of the microfluidic module. Furthermore, being different from other fixed actuators driven by high-speed electric motor [8,23], each piezoelectric actuator can be programmable controlled by microcontroller, providing real-time adjustable travelling wave applied to the peristaltic pump. The small dead volume, real-time control and programmable of each step, enable it suitable for biochemical microfluidic applications. An on-chip flow mixing experiment was carried out using two integrated micropumps, demonstrating
A prototype of the peristaltic micropump is shown in Fig. 2a, which consists of a microchannel and three piezoelectric actuators, the microchannel has a flexible PDMS (polydimethylsiloxane) membrane as the top layer. Three enlarged chambers are specially designed in alignment with the actuators, as shown in Fig. 2b. When the three piezoelectric actuators are applied voltage signals, the actuators generate impacting actions. It’s the impacting actions that compress the chambers, and then squeeze the liquid out of the chamber into microchannel in two directions. With the three pins impacting on the chambers in sequence as Fig. 2c, liquid is transported in a designed direction. A demo cycle with six steps is listed as Fig. 2d. At the beginning three piezoelectric actuators are driven in high-level voltages and the three cantilevers bend toward the chip, enabling the pins to compress the chambers. Next, the first actuator is driven in low-lever voltage, making the first pin and the flexible membrane retract. Liquid flows into the chamber due to the generated negative pressure. In the rest steps the three actuators are driven at voltages as shown in Fig. 2c, and the three pins stroke the membrane in sequence, which transports the liquid from left to right. We assume that “1” stands for a pin impacting on the chamber, and “0” for non-impacting on the chamber, thus a peristaltic period can be expressed as 111-011-001-101-100-110, as shown in the Fig. 2d. 2.2. Calculation model of the transported volume Liquid in the microchannel shows a displaced flow when the flexible membrane is deformed by the pin (with volume Vp ). As shown in Fig. 3a, instead of a dynamic model of transported volume [24], we supply a simplified fluidic electric model can be used to calculate the transported volume from the outlet of each action (V´i ) as [25] Vi’ =
Ri1 Vp , i = 1, 2, . . ., 6. Ri1 + Ri2
(1)
Ri1 (Ri2 ) is the flow resistance of the microchannel, which is in analogue with Ohm’s resistance of electricity [26,27], is given by R≈
1 12l . wh3 1 − 0.63 h w
(2)
92
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
Fig. 2. a) The picture of the integrated peristaltic micropump and the microfluidic chip; b)the design of the microchannel of the peristaltic micropump; c) electric waveform of each actuator; d) and the pump sequence used for the piezoelectric impacting actuators.
Fig. 3. Characteristics of the micropump: a) flow rates at varied voltages of actuators; b) flow rates at varied working frequency of actuators; c) relationship between flow rate and backpressure.
where, w, l, and h are the width, length and height of the microchannel, is the viscosity coefficient of liquid, and R is the flow resistance of the microchannel, respectively. It’s worth noting that the flow resistance is changed when the acting actuator is at different positions (P1 , P2 , and P3 ). Two mainly factors induce the change, which can be seen from Fig. 3b. One is the geometrical changes in the length of channels, and the other is the changes in height of the three circular chambers caused by pin’s squeeze. The total transported volume can be estimated by a sum of all six movements. In each movement, only one actuator is designed to change its status. We use V‘1 , V‘2 , V‘3 , V‘4 , V‘5 and V‘6 representing the transported volumes of each action out of the outlet, respectively. Then the volume of total transportation V’ in one period is V ’ = V1’ + V2’ + V3’ + V4’ + V5’ + V6’
(3)
Substitute Eq. 2 into Eq. 1, then into Eq. 3, the transported volume in one period can be expressed as
V ’ ≈ f W, w, l1 , l2 , l3 , lp , lr , h · s.
(4)
Details of the derivation are given in the supplementary material from Eq.S11 to Eq.S12. In Eq. 4, l1 , l2 , l3 , l4 are the lengths of different sections, as is shown in Fig. 3b. lp is the equivalent length of the compressed chamber, and lr is the equivalent length of uncompressed chamber. W and w are the widths at different sections, and h is the height of the microchannel, s is the deformation of flexible membrane, which is equal to the bending displacement of the piezoelectric actuators. And the displacement is related to the driving voltage of the piezoelectric actuator. In brief, the transported
liquid volume can be directly controlled by regulating the driving voltages as well as changing of channel sizes. 2.3. Micropump system To squeeze the liquid in the microchannel, three piezoelectric cantilevers (QDTE52, PANT Corp, China) were adopted with three micro pins (1356-1 & 1356-3, Keystone Electronics Corp, USA) adhered at the ends of cantilevers by epoxy adhesive (467 M P, 3 M Corp, USA), respectively. The microfluidic chip was fabricated using the standard soft lithography process [11]. Firstly, a mold was made using SU-8 (SU-8 2025, MicroChemicals Corp, Germany) material by UV lithography. Secondly, the PDMS (Sylgard 184, Dow-Corning Corp, USA) mixed at a ratio of 1:10 (curing agent to base) was casted onto the SU-8 master mold to form microfluidic channel. Then the solidified PDMS with microchannel was bonded to a glass slide by oxygen plasma (PDC-MG, Mingheng Corp, China). The width w, the height h of the channel and the total thickness of the PDMS layer were 355.7 m, 54 m and 355 m, respectively. l3 was 2 mm, and the total length (from inlet to outlet) was 35 mm. Diameters of the pin and the chamber were 0.6 mm and 1.0 mm, respectively. 2.4. Calibration methods To characterize the micropump, flow rate was calibrated by measuring the transported liquid volume through the micro channel during pump working. Here we measured flow volume by monitoring the liquid levels in the glass tubes, which were connected to the inlet and outlet (Fig. 2a), respectively. The inner
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
93
Fig. 4. Characteristics of the micropump: a) flow rates at varied driving voltages; b) flow rates at varied working frequency of actuators; c) relationship between flow rate and backpressure.
diameter of the tube is 2.90 mm. The operating process is as follows: Firstly, DMSO (Dimethyl Sulphoxide) with 5% (w/w) blue dye (GT52, HP Corp, USA) was injected from the inlet tube to the outlet tube, to achieve a same liquid level. By measuring the real-time change of liquid level during working, we can get the flow rate of the micropump. At this moment, the piezoelectric cantilevers were driven by programmable pulse-width-modulation (PWM) square waveforms (range: DC 60 V–180 V), which were generated by an FPGA controller circuitry. The backpressure can also be calibrated through calculating the pressure difference, which was related to the height difference in the two tubes using the same setup, except that a longer outlet tube was required. A high-speed camera (Chronos 1.4, Chronos Corp, Canada) was used to capture the mixing process on a chip. 3. Results and discussion 3.1. Characteristics of the micropump Owing to small size of the microchannel, the dead volume is less than 300 nL. It means that when the micropump is integrated on a microfluidic chip, there is only microamount reagent wasted, which is particularly important in using expensive biological reagents. The measured process of the stop of the micropump using high speed camera shows a waiting time less than 12.5 ms (Fig.S2), indicating a fast response time in the scale of millisecond. A series of calibration experiments were carried out to inspect the characteristics of the peristaltic micropump. As shown in Fig. 4a, the flow rate is monotonously increased with the driving voltage. According to Eq. 4, the transported volume has positive correlation with the displacement of the actuator, which is proportional to the driving voltage. When the voltage is 90 V and the working frequency is 25 Hz, the flow rate is 8.4 nL/s. Meanwhile, the minimum repeatable volume of the transported liquid in one cycle is only about 2.0 nL. This volume resolution can be further improved by using smaller driving voltage or longer channel. Actually, we have ever achieved flow rate resolution of 0.08 nL/s with length (l3 ) of 8 mm and driving voltage of 150 V (Fig.S3). On the other hand, the flow rate is also increased with the driving frequency in the range from 0 to 100 Hz. And the relationship between the flow rate and the frequency is almost linear when driving voltage is larger than 90 V. Since higher frequency means more working cycles within unit time, the micropump can transport more volume of liquid in same time. In Fig. 4b, the measured flow rate at 180 V (red solid line) is very close to the theoretical result (black solid line), especially at low frequency (<50 Hz). The theoretical result is drawn from Eq. 4, representing a maximum transported rate. At the moment, the flexible membrane on the chamber got a maximum deformation, which is equal to the height
of the microchannel. When the driving frequency is larger than 75 Hz, the error between the theoretical and experimental results is obvious because the theoretical result is based on a static model without considering the inertia of the liquid. Peristaltic micropump has a special capability of transporting liquid bi-directionally. By adjusting the impacting sequence in controlling program, our integrated micropump can precisely change flow direction at any working step. Here, we calibrated the flow rates at two directions at 150 V using the chip presented in Fig. 2c. As shown in Fig. 4b, when frequency is less than 100 Hz, the flow rates of both forward and inverse directions are almost equal. They are both increased with frequency. The differences between the two directions at larger frequency could mainly be caused by the installation error, which makes the initial gaps between the three pins and the membrane unequal in vertical position. Observed from the experiments, we found that the influence of the installation error on the symmetry of the bidirectional flow rates was enhanced with the frequency and voltage. Thus, a careful alignment between the pump chip and the three actuators is very essential and important. At last, we studied the flow rate change at the different back pressures. The maximum back pressure can reach 2.0 K Pa, as shown in Fig. 4c. 3.2. Application of the peristaltic micropump: mixing on a chip In biochemical applications, titrations are performed not only for a given compound but also for a mixture of two or more compounds, including their titration [28]. The mixing of biological or chemical species can be applied to digital polymer chain reaction [7,29], biochemical analyses [30], nanomaterial synthesis [31], or emulsification [32]. In order to achieve effective, low consuming, rapid and complete mixing, various types of micro-mixers have been created to improve the mixing in microscale. The enhancement of mixing is usually achieved by increasing the mixing area [33]. In our demo, we conducted a mixing process based on a microfluidic two-phase mixing chip integrated with the peristaltic micropump. Two colored liquids were efficiently mixed at both low and high flow rates. As shown in Fig. 5a, there are two microfluidic pump modules, pump one is for yellow dye, and pump two for blue dye. Two groups of piezoelectric actuators were adopted, forming two peristaltic micropumps together with the two pump modules (Fig. 5b). A zigzag channel with length lz is 10 mm, applied as mixing module to assist mixing for its ability of improving the mixing efficiency [34]. The microfluidic two-phase mixing module was fabricated in one step with the integrated pump modules. The width of inlet of “Y” junction wi is 200 um, and its width of outlet wo is 500 um, the height of microchannel is 54 um. It was easy to generate the initial laminar flow by making the two micropumps work simul-
94
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
Fig. 5. a) A photo of a microfluidic chip showing the pumping area and zigzag mixing area; b) photo of the microfluidic mixing setup.
Fig. 6. The mixing process by two peristaltic pumps showing a) the initial state when both fluids were laminar flow at the frequency of 100 Hz and voltage of 150 V. When the frequency was reduced to 10 Hz, pulsation was obviously as follows: b) the mixed fluids was flowing back to the yellow colored channel; c) the mixed fluids was flowing back to the blue colored channel; d) the liquids in the channel after fully mixed by alternative transportation; e) the mixing effects at position A and position B (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article).
taneously, as shown in Fig. 6a. There was only a stable diffusion layer existing between two colored layers where the inefficient mixing was completely realized by molecular diffusion. In this situation, the flow rate was about 83 nL/s for each micropump when the working frequency and applied voltage were 100 Hz and 150 V, respectively.
3.2.1. Mixing at low frequency As mentioned in the section of working principle, fluid is squeezed out of the chamber and flows in both directions when the pin impacts on the membrane. On the other hand, fluid is also sucked back into the chamber in both directions when the pin reverts. During the six-step cycle, the fluid in the pump chip moves back and forth alternately, and lastly moves forward due to the impacting sequence. Usually the pulsation of the peristaltic pump is disadvantageous in the application where a stable flow is required. Nevertheless, mixing using microfluidic chip is benefited from pulsation or disturbance flow, since the fluid is always laminar flow in the microchannel. This pulsation of the designed peristaltic micropump was especially obvious in the experiment at low driving frequency (e.g. 10 Hz). When the work sequences of the two pumps weren’t completely synchronous, the liquid in the Y junction moved back and forth into the two pump channels alternately, generating a turbulent flow, as shown in Fig. 6b & c. Furthermore, due to the low flow rate (5 nl/s) at low driving frequency (10 Hz), fluids were mixed well just behind their confluence (Fig. 6d). A movie in the Supplementary Files (Video-1) shows the mixing process. Here, we describe the mixing effect using relative mixing index (RMI) [35,36], where RMI = 1 and RMI = 0 indicate complete separation and complete mixing, respectively. By calculating RMI, the mixing level is quantitatively evaluated by substituting a standard
deviation, which is a light intensity distribution in a specified region inside a channel as
RMI =
= 0
N 1
(Ii − Im )2
N
i=1
N
(I − Im )2 i=1 oi
N 1
.
(5)
Where, is the standard deviation of the pixel intensities, o is the standard deviation of the pixel intensities in the unmixed case, Ii shows a local pixel intensity, I0i represents the local pixel intensity in the unmixed state, Im shows the average of the pixel intensities in the cross section, N represents the total number of pixels. As shown in Fig. 5e, at position A, turbulent mixing began immediately after the two pumps worked at low-frequency mode. And it became stable in about 10 s when the RMI was 0.278. After another 40 s, these fluids were further mixed with RMS of 0.188 at position B when flowing through the zigzag channel. It can be seen that the peristaltic micropumps can mix the fluids in a short and straight channel quickly, even without the zigzag channel. However, the zigzag channel can help improve the mixing effect. 3.2.2. Mixing at high frequency The pulsation of the peristaltic pump was gradually weakened with the driving frequency increasing. And this attenuating pulsation showed a regressive effect on the aforementioned mixing. For example, when the frequency was 100 Hz, the fluids in the microfluidic chip returned to laminar flow. For mixing in microchannel, some researchers used specially-designed microchannel configurations [33,37,38] to enhance mixing effect. Here, we proposed a new mixing method by generating an arbitrary pulsation at high frequency, without specially-designed microchannel configurations. Being beneficial from the peristaltic pump’s characteristic of bidirectional transportation, we made one pump intermittently
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
95
Fig. 7. The mixing process by two peristaltic pumps at high working frequency: a) pulsation waveform of the flow rate of pump one, the negative and positive value of flow rate represent the direction of transportation; b) pulsation waveform of the flow rate of pump two; c)-e) pump one transports the yellow ink in pulsation and pump two transports liquid continuously, the red arrow shows the direction of the flow; f) the liquids after fully mixing by alternative transporting ; g) the mixing effects at position A and position B. (For interpretation of the references to colour in this figure legend, the reader is referred to the web version of this article).
transport liquid in inverse direction during the normal forward transportation, while the other pump still worked in constant forward direction. A movie in the Supplementary Files (Video-2) shows the mixing process. As shown in Fig. 7a, the liquid was transported in the positive direction in the time of t1 , and then transported reversely in the time of t2 . Here, t1 and t2 are 3.6 s and 0.36 s, respectively. Due to the small size (9 mm) of pump chip, the peristaltic pump can change transporting direction during working in milliseconds by just importing driving signals in different sequence. As can be seen from Fig. 7c-e, the reciprocating flow of fluids by pump one generates a similar artificial turbulent flow as that at low frequency near the Y junction. After 5 periods, the mixing effect was getting better gradually as Fig. 7f. The relative mixing index (RMI) was calculated to describe the mixing effect. As is shown in Fig. 7g, at position A, the reciprocating flow generated by pump one shows a positive effect on the mixing. It decreased the RMI slowly. With the combined effect of the zigzag microchannel, the mixing efficiency was further improved, i.e. the RMI at the position B was reduced from 0.24 to 0.065. Compared to native pulsation of the peristaltic pump at the low-frequency mixing, the new arbitrary pulsation at high frequency shows a better mixing effect, more flexible strategy and larger mixing efficiency. 4. Conclusion In this paper, we presented a peristaltic micropump that was composed of separated microfluidic channel and piezoelectric actuators. The unique design of integrating the channel with other functional microfluidic modules makes it suitable for building a plug-and-play, conceptual pump-on-a-chip. The peristaltic micropump has been successfully designed, fabricated, and calibrated, showing a high resolution of 8.4 nL/s, and dead volume less than of 300 nL. In addition, a demonstration of mixing in one chip using two peristaltic micropumps was carried out. Two working modes at both low and high driving frequencies were presented, which realized an effective mixing in a microfluidic chip with only simple-structure channel. The results of mixing have demonstrated
the excellent performances and high reliability of the micropump, showing great potential for applications such as point-of-care (POC) testing, drug delivery, and chemical analysis. Acknowledgements This research work was supported in part by the National Natural Science Foundation of China (No. 51675505), the Joint Research Fund for Overseas Chinese, Hong Kong and Macao Young Scientists of the National Natural Science Foundation of China (No. 51628502). Appendix A. Supplementary data Supplementary material related to this article can be found, in the online version, at doi:https://doi.org/10.1016/j.sna.2019.04. 005. References [1] D.J. Laser, J.G. Santiago, A review of micropumps, J. Micromech. Microeng. 14 (2004) R35. [2] K.W. Oh, C.H. Ahn, A review of microvalves, J. Micromech. Microeng. 16 (2006) R13. [3] A. Pourmand, S.A.M. Shaegh, H.B. Ghavifekr, E.N. Aghdam, M.R. Dokmeci, A. Khademhosseini, Y.S. Zhang, Fabrication of whole-thermoplastic normally closed microvalve, micro check valve, and micropump, Sens. Actuators B Chem. 262 (2018) 625–636. [4] C.D. Chin, V. Linder, S.K. Sia, Commercialization of microfluidic point-of-care diagnostic devices, Lab Chip 12 (2012) 2118–2134. [5] A. Nisar, N. Afzulpurkar, B. Mahaisavariya, A. Tuantranont, MEMS-based micropumps in drug delivery and biomedical applications, Sens. Actuators B Chem. 130 (2008) 917–942. [6] S. Herrlich, S. Spieth, S. Messner, R. Zengerle, Osmotic micropumps for drug delivery, Adv. Drug Deliv. Rev. 64 (2012) 1617–1627. [7] C. Zhang, D. Xing, Y. Li, Micropumps, microvalves, and micromixers within PCR microfluidic chips: advances and trends, Biotechnol. Adv. 25 (2007) 483–514. [8] H. Yu, W. Ye, W. Zhang, Z. Yue, G. Liu, Design, fabrication, and characterization of a valveless magnetic travelling-wave micropump, J. Micromechanics Microengineering 25 (2015), 065019.
96
T. Ma et al. / Sensors and Actuators A 292 (2019) 90–96
[9] Y.N. Wang, L.M. Fu, Micropumps and biomedical applications–A review, Microelectron. Eng. 195 (2018) 121–138. [10] J.G. Smits, Piezoelectric micropump with three valves working peristaltically, Sens. Actuators A Phys. 21 (1990) 203–206. [11] M.A. Unger, H.P. Chou, T. Thorsen, A. Scherer, S.R. Quake, Monolithic microfabricated valves and pumps by multilayer soft lithography, Science 288 (2000) 113–116. [12] T. Thorsen, S.J. Maerkl, S.R. Quake, Microfluidic large-scale integration, Science 298 (2002) 580–584. [13] A.K. Au, N. Bhattacharjee, L.F. Horowitz, T.C. Chang, A. Folch, 3D-printed microfluidic automation, Lab Chip 15 (2015) 1934–1941. [14] Y.S. Lee, N. Bhattacharjee, A. Folch, 3D-printed quake-style microvalves and micropumps, Lab Chip 18 (2018) 1207–1214. [15] H. Landari, M.A. Dussault, J. Ruel, A. Begin-Drolet, A. Miled, Biocompatible compact micropump with integrated unidirectional microvalves for low pressure microfluidic applications, Sens. Actuators A Phys. 276 (2018) 246–258. [16] X. Zhang, Z. Chen, Y. Huang, A valve-less microfluidic peristaltic pumping method, Biomicrofluidics 9 (2015), 014118. [17] J. Xiang, Z. Cai, Y. Zhang, W. Wang, A micro-cam actuated linear peristaltic pump for microfluidic applications, Sens. Actuators A Phys. 251 (2016) 20–25. [18] M. Rusli, P.S. Chee, R. Arsat, K.X. Lau, P.L. Leow, Electromagnetic actuation dual-chamber bidirectional flow micropump, Sens. Actuators A Phys. 282 (2018) 17–27. [19] J. Xie, J. Shih, Q. Lin, B. Yang, Y.C. Tai, Surface micromachined electrostatically actuated micro peristaltic pump, Lab Chip 4 (2004) 495–501. [20] M. Shen, L. Dovat, M.A. Gijs, Magnetic active-valve micropump actuated by a rotating magnetic assembly, Sens. Actuators B Chem. 154 (2011) 52–58. [21] 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. [22] C.S. Chao, P.C. Huang, M.K. Chen, L.S. Jang, Design and analysis of charge-recovery driving circuits for portable peristaltic micropumps with piezoelectric actuators, Sens. Actuators A Phys. 168 (2011) 313–319. [23] J. Han, O. Rahmanian, E. Kendall, N. Fleming, D. DeVoe, Screw-actuated displacement micropumps for thermoplastic microfluidics, Lab Chip 16 (2016) 3940–3946. [24] C. Jenke, S. Kager, M. Richter, C. Kutter, Flow rate influencing effects of micropumps, Sens. Actuators A Phys. 276 (2018) 335–345. [25] Y. Ding, E. Huang, K.S. Lam, T. Pan, Microfluidic impact printer with interchangeable cartridges for versatile non-contact multiplexed micropatterning, Lab Chip 13 (2013) 1902–1910. [26] M. Tanyeri, M. Ranka, N. Sittipolkul, C.M. Schroeder, A microfluidic-based hydrodynamic trap: design and implementation, Lab Chip 11 (2011) 1786–1794. [27] A.J. Teo, K.H.H. Li, N.T. Nguyen, W. Guo, N. Heere, H.-D. Xi, C.-W. Tsao, W. Li, S.H. Tan, Negative pressure induced droplet generation in a microfluidic flow-focusing device, Anal. Chem. 89 (2017) 4387–4391. [28] A. Folch, Introduction to bioMEMS, CRC Press2016, Boca Raton, 2016, pp. 191–204.
[29] M. Hashimoto, F. Barany, S.A. Soper, Polymerase chain reaction/ligase detection reaction/hybridization assays using flow-through microfluidic devices for the detection of low-abundant DNA point mutations, Biosens. Bioelectron. 21 (2006) 1915–1923. [30] C. Neils, Z. Tyree, B. Finlayson, A. Folch, Combinatorial mixing of microfluidic streams, Lab Chip 4 (2004) 342–350. [31] D. Ahmed, X. Mao, J. Shi, B.K. Juluri, T.J. Huang, A millisecond micromixer via single-bubble-based acoustic streaming, Lab Chip 9 (2009) 2738–2741. [32] S. Okushima, T. Nisisako, T. Torii, T. Higuchi, Controlled production of monodisperse double emulsions by two-step droplet breakup in microfluidic devices, Langmuir 20 (2004) 9905–9908. [33] C.Y. Lee, C.L. Chang, Y.N. Wang, L.M. Fu, Microfluidic mixing: a review, Int. J. Mol. Sci. 12 (2011) 3263–3287. [34] V. Mengeaud, J. Josserand, H.H. Girault, Mixing processes in a zigzag microchannel: finite element simulations and optical study, Anal. Chem. 74 (2002) 4279–4286. [35] A. Hashmi, J. Xu, On the quantification of mixing in microfluidics, J. Lab. Autom. 19 (2014) 488–491. [36] R. Sakurai, K. Yamamoto, M. Motosuke, Concentration-adjustable micromixer using droplet injection into a microchannel, Analyst 4 (2019) 559–568. [37] C.C. Hong, J.W. Choi, C.H. Ahn, A novel in-plane passive microfluidic mixer with modified Tesla structures, Lab Chip 4 (2004) 109–113. [38] C.Y. Lee, L.M. Fu, Recent advances and applications of micromixers, Sens. Actuators B Chem. 259 (2018) 677–702.
Biographies Tuo Ma received his BS degree in June 2010 from Anhui University. He is currently pursuing the PhD degree at Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China. His research interests include Lab-on-a-chip, integrated microfluidic system and actuators. Shixin Sun received her BS degree in June 2018 from Northeast Agricultural University. She is currently pursuing the MS degree at Department of Precision Machinery and Precision Instrumentation, University of Science and Technology of China. Her research interest is Lab-on-a-chip. Baoqing Li received his BEng, MEng and PhD in mechanical engineering and precision instrumentation from University of Science and Technology of China (USTC) in 2000, 2003 and 2008, respectively. He was a research assistant at Hong Kong Polytechnic University in 2004–2006, and a Post-Doctoral Fellow at University of California, Davis in 2013–2015, working on microfluidic technology. Currently he is an associate professor at USTC, and his research interests are Lab-on-a-chip and MEMS. Jiaru Chu received his BEng, MEng and PhD in mechanical engineering and precision instrumentation from University of Science and Technology of China (USTC) in 1985, 1988 and 1997, respectively. He currently is a professor at USTC, and his interests are MEMS, laser fabricating technology and telescope instruments.