ME381: FINAL PROJECT REPORT

Transcription

1 ME381: FINAL PROJECT REPORT Analysis of Thermo-Pneumatic and Piezoelectric Actuation in MEMSbased Micropumps for Biomedical Applications Kenneth D Aquila, Sean Tseng Northwestern University 12/10/07

2 TABLE OF CONTENTS Abstract 3 Project Description 3 Motivation 3 Objective 3 Thermo-Pneumatic Actuation 4 General Description 4 Input Electrical Energy 5 Flow Rate 6 Microfabrication/Materials 7 Piezoelectric Actuation 9 A Brief and Interesting History 9 Fundamentals 9 Piezoelectric Micropumps 11 Actuation Mechanism 11 Microfabrication 12 Governing Equations 13 Variables and Optimization 15 Comparing Piezoelectric to Thermo-Pneumatic 18 Conclusion 18 References 19 Biographies 20 2

3 ABSTRACT The constantly growing demand for more effective drug therapies and improved DNA/protein processing continues to define a role for the distinct advantages of MEMS micropumps. This report investigates thermo-pneumatic and piezoelectric actuation and draws conclusions regarding the unique electrical, flow rate, and fabrication characteristics of each type. For thermo-pneumatics, the quest for adequate performance and minimal costs of fabrication and electrical input has motivated increasingly clever designs. For piezoelectric type micropumps, complex mechanical behavior can be modeled numerically and used to motivate better engineered designs. In the end, it is application and specific device usage that defines whether piezoelectric or thermo-pneumatic is the optimal choice for micropumping actuation. PROJECT DESCRIPTION Motivation Improving drug delivery systems (DDS) is critically important to maximize the efficacy of drug therapies. Increasing controllability of this process helps improve dosage and specificity. Dosage is vital to ensure that drug concentrations are high enough to avoid trivial treatment but low enough to prevent intoxication of the body. Specificity helps position the drug for release into biological systems for optimum effect and helps minimize any collateral damage to healthy tissues. Drug delivery micropumping utilizing micro-electro-mechanical systems (MEMS) offer high controllability in these areas and thus have a great potential for drug therapy. Furthermore, there is a strong technological opportunity for MEMS in the development of micrototal analysis systems (µ-tas). µ-tas, also known as lab-on-a-chip technology, may quickly and efficiently process and analyze fluid containing DNA, proteins, and drug molecules. An essential component to such systems is the micropump, which must propel the sample fluid with a flow rate appropriate to the specific application while expending minimal electrical energy. Objective In this project, we seek to analyze and compare two different micropump actuation mechanisms based on their required electrical energy input, flow rate performance, and fabrication/material considerations. Specifically, thermo-pneumatic and piezoelectric actuation have been chosen as promising routes for the continued development of micropumping in DDS and µ-tas. While both options are certainly viable, each has unique micropumping characteristics. In addition recent improvements in these micropumping characteristics are a direct result of creative design strategies and careful engineering/modeling. The following analysis will first analyze thermopneumatic and piezoelectric actuation and then compare the advantages and disadvantages of each, in light of specific applications for DDS and µ-tas. 3

4 THERMO-PNEUMATIC ACTUATION General Description Figure 1 shows a schematic for the typical components of a thermo-pneumatic micropump. The standard mechanism for thermo-pneumatic micropump actuation based on resistive heating of a solid heating element such as Indium Titanium Oxide (ITO), Cr/Au, or Ti/Al that is located inside an enclosed actuation chamber. As a result, the air in the actuation chamber expands and generates a pressure on the flexible membrane which seals one end of the chamber. When the applied voltage is removed, the heater and chamber air cool down, the pressure is reduced, and the membrane moves back to its original position. By cyclically applying voltage of 1-20 V, the membrane can be made to reciprocate usually around a few Hertz. The major variations of this design pertain to the way fluid flow is rectified (check-valves, nozzle/diffusers, peristaltic action) and the body that actually pushes the fluid (flat membrane, corrugated membrane, membraneless air bubble). The following analysis presents the rationale for these design options in terms of three important micropump parameters: Applied electrical energy, fluid flow rate, and device fabrication/materials. See Table 1 for a summary. Figure 1: Basic Thermo-Pneumatic Micropump Schematic 1 Pumping Chamber Actuation Chamber Table 1: Thermo-pneumatic Micropump Design Summary Name Year Variant Type Input Electrical Flow Rate Materials Jeong Nozzle/Diffuser, 8 V, Corrugated Membrane 40% Duty at 4 Hz 14 µl/min Doped Silicon Jeong Peristaltic, 20 V, Flat Membrane 50 % Duty at 2 Hz 21.6 µl/min PDMS, Cr/Au Jun Surface Tension, µl/min, 3.5 V Air Bubble 116 nl in 5 min PDMS, Ti/Al Van de Pol Check Valves,??? V, 30 µl/min, Flat Membrane 0.5 Hz Silicon Yoo Nozzle/Diffuser, 500 mw, 0.73 µl/min Flat Membrane 1% Duty at 2Hz PDMS, ITO Yoo Nozzle/Diffuser, 500 mw, PDMS, ITO, 1.05 µl/min Flat Membrane 7% Duty at 2Hz Parafilm 4

5 Input Electrical Energy The applied electrical energy for a micropumping system faces two slightly conflicting requirements. On one hand, the electrical power must be low for two reasons. First low power devices are more economical and may be powered by portable means such as batteries. Second, low voltages are important for safety since drug delivery systems (DDS) inevitably require human contact. On the other hand, it is crucial for the dosing accuracy of micropumps that pump strokes are powerful enough to resist any naturally variation in back pressure. 2 As seen in Figure 2, back pressure can change the nominal flow rate because a decreased stroke volume results. One way to compensate for this is to use a calibrated membrane stop and apply a large enough voltage so that the membrane will always deflect against the stop despite some variation in back pressure. The bottom line objective for micropumping is to minimize the required electrical power while maintaining consistent volume flow rate. Figure 2: (Left) Flow rate variation with Back Pressure 3 Figure 3: (Right) Deflection/Temperature vs. Voltage 3 Thermo-pneumatic micropumps tend to use significant electrical power because of the cyclical heating/cooling cycles during operation. In many thermo-pneumatics, the applied voltage must resistively heat air to generate a chamber pressure that can cause membrane deflection. So in general, higher voltage will generate greater gas temperature/expansion, and greater volume pump strokes (greater membrane deflection). As shown in Figure 3, Jeong et al. has built micropumps which followed a nearly linear Voltage vs. Temperature relationship with a slope of unity (20V applied voltage corresponds to 20 C of heater temperature rise). For this micropump, there was also a monotonic increase of membrane deflection with voltage (20 V achieved 100 µm membrane deflection). 3 While there have been many micropump design strategies for reducing the required applied voltage, here are three creative examples. Yoo et al. has used a nozzle/diffuser rectified pump. 4 Early thermo-pneumatic pumps utilized check valves which forced the pump membrane to reach at least a threshold pressure before each pump stroke could begin. 5 Like most valve-less pumps, this newer design requires less electrical energy because less pressure is needed to rectify solution flow than with the check values. The added flexibility of a corrugated membrane is another solution to lower voltage. Figure 4 displays how an applied voltage of 6 V deflects the 5

6 corrugated p + silicon membrane 3 times as much as a flat membrane, thus efficiently utilizing electrical energy. 6 Recent efforts to utilize PDMS instead of Si for pump membranes is partly motivated by the efficiency advantage of a more flexible membrane material as well. Finally, some microvalves, in a design by Yoo et al., utilize parafilm inside the actuation chamber. When heated, the parafilm s phase change from solid to liquid is accompanied by a significant volume change which helps maximize the membrane deflection for a given applied voltage. 7 Figure 4: (Left) Increased deflection for a more flexible pump membrane 6 Figure 5: (Right) Schematic of a membrane-less thermo-pneumatic pump based on air bubbles 8 Flow Rate Different types of thermo-pneumatically actuated MEMS for micropumping attain different degrees of continuity and various volume flow rates. Each type may be suitable for specific combinations of application (DDS, µ-tas) and use (slow and precise delivery of small fluid amounts with a disposable system or prolonged or high flow rate delivery of fluids for good throughput) For high flow rate applications/usage, Jeong et al. have developed a membrane-based micropump consisting of three actuation chambers that work together to transport fluid by peristaltic action. Unlike other single actuator pumps that have dedicated pumping and recovery strokes, the combined effect of the three actuator system is that fluid is always being moved forward. This continuous and high flow (21.5 µl/min) comes at the price of higher applied voltage and a larger duty ratio. 3 The duty ratio is fraction of a pumping cycle that the voltage is applied. Large voltage and duty ratio imply more electrical energy is being applied per cycle. For low flow rate applications/usage, Jun Sim and Yang have demonstrated a novel membraneless pumping system that simply utilizes air bubbles, surface tension, and nozzle/diffuser structures. According to Figure 5, air is heated resistively in two separate air chambers, expanded out of the chambers mostly through the diffuser openings, and made to form one bubble on each side of the microchannel. Fluid lying between these two growing bubbles is moved forward down the micro channel and eventually pushed even farther once the two bubbles combine to form an air space. Upon cooling, the air moves back to the two chamber via the nozzle openings and simultaneously sucks more fluid from upstream. Without having to push on a membrane, 6

7 the required voltage and is much lower (3.5 V) and since the flow rate is low (0.023 µl/min), this bubble mechanism can achieve precise fluid delivery resolution (116 nl over 5 minutes). 8 Microfabrication/Materials Now that a large variety of micropump designs exist, the ongoing challenge is to develop one with adequate performance but with low power expenditure and minimal fabrication/material cost. Research efforts seem to have recently changed focus from traditionally high performing and expensive fabrication of doped silicon structures to less expensive and more simple fabrication of poly-dimethlysiloxane (PDMS) structures. Following Jeong et al., fabrication of the corrugated membrane micropump begins with a 450 µm thick Si (100) wafer. The first three steps involve thermal oxidation and etching of silicon. A 0.8 µm thermal oxide is grown, 5 µm of front-side silicon are etched using EPW (Ethylendiamine:Pyrocatechol:Water), then another 0.5 µm oxide layer is grown. In the fourth step, lithography is used to define a concentric ring pattern in a resist layer followed by more EPW etching. This produces a pattern of annular groves in the Si which is then infiltrated by Boron doping from a vaporized solid source (1100 C 10 hrs). After removing the thin layer of borosilicate glass (at 1100 C again), the wafer is backside etched which removes Si and leaves the corrugated, boron-doped membrane structure. To complete fabrication, electrode materials and heating elements are patterned by e-beam deposition on a separate piece of glass. Then the patterned glass piece and the silicon membrane structure are joined by anodic bonding (300 C, 800V), leaving a small air space between the corrugated membrane and the glass wafer. All together, this processing required 9 major steps, 5 involving temperatures greater than 300 C and 2 lithography steps. See figure 6 for the fabrication overview. Figure 6: Fabrication overview of thermo-pneumatic micropump with corrugated Si membrane 6 In contrast, the fabrication of PDMS based thermo-pneumatic micropumps involves 11 steps summarized in figure 7. All of them are below 150 C, but they do involve some manual peeling/punching. First the layer containing the actuation chambers is made by spin-coating PDMS polymer mixture and curing at 65 C for 15 min. This layer is then peeled off, hole punched to create the actuation chambers, and then bonded to glass at 65 C. This glass is prepatterned with heater element materials prior to bonding. Next, a separate silicon wafer coated with negative resist is patterned with lithography and covered with the PDMS mixture. Following curing (again at 65 C), the PDMS is peeled off from the resist-si master to reveal the microchannels and pumping chambers in the PDMS. After this, another separate piece of silicon is spin coated with the PDMS mixture to create a 30 µm thin pumping membrane. The previously mentioned piece of PDMS with the microchannels is bonded to this PDMS membrane 7

8 layer (at 65 C). Finally, these joined PDMS layers are peeled off from the silicon and bonded (at 65 C) to the original PDMS-glass structure containing the actuation chambers and heating elements. The whole structure is now further cured at 100 C. Figure 7: Fabrication overview of thermo-pneumatic pump based on PDMS 3 There are some main differences between Si-based and PDMS-based micropump fabrication. As far as creating pumping and actuation chambers, Si-etching has essentially been replaced by punching holes in PDMS. In terms of wafer bonding, high temperature/voltage anodic bonding of silicon wafers is replaced by low temperature layer bonding steps with PDMS. In addition, PDMS polymer mixtures containing PDMS pre-polymer and a curing agent are typically less expensive then single crystal silicon wafers. 8

9 PIEZOELECTRIC ACTUATION A Brief and Interesting History The term piezo comes from the Greek word for pressure. The first discovery of piezo effect was credited to two brothers, Jacques and Pierre Curie, in They realized that specially prepared quartz produced measurable charge when subjected to mechanical stress. This was known as the direct piezoelectric effect. It wasn t until a mathematical model utilizing the fundamentals of thermodynamics created by Gabriel Lippmann that the inverse piezoelectric effect was shown to exist. However, due to the difficult mathematics to understand piezoelectric effect, and the realization of electro-magnetic technology during that time, which produced highly-visible and remarkable machines, little work was done to further the study of piezoelectricity. It took until World War I when a group of French researchers developed an ultrasonic submarine detector using a quartz crystal transducer. This did not go unnoticed by the industrialized nations and research during World War II in the U.S., Japan and Soviet Union led to the development of the barium titanate piezoelectric family, which also led to the all important lead zirconate titanate (PZT) family. Today, researchers are finding and creating new materials displaying even larger piezoelectric effect, and making them more rugged and cost-effective to produce. 11 Fundamentals As mentioned in the previous section, the PZT ceramics family is the most important piezoelectric material because of its wide use in industry. Figure 8 shows the unit cell of PZT above (top figure) and below its Curie temperature (bottom figure) ). The top figure showss a cubic and completely symmetrical unit cell, but it no longer has a net dipole moment resulting in the loss of piezoelectric effect. The bottom figure shows a PZT unit cell that is naturally deformed tetragonally when below the Curie temperature, allowing for piezoelectric effect. Piezoelectric effects are anisotropic by nature so it is beneficial to utilize tensor mathematics. The axes are identified as shown in Figure 2. Axes 1, 2, and 3 represent the classical notation of X, Y, and Z, and axes 4, 5, and 6 represent the rotational and shear terms. Note that the direction of polarization is traditionally shown in the axis 3 direction. 9 Figure 8: (1) PZT unit celll above Curie Temperature. (2) PZT unit cell below Curie temperature. 1 0

10 The following are equations used to describe the piezoelectric effect using tensor notation 12 : Eq. (1) where charge displacement piezoelectric constant C/N stress permittivity at constant stress electric field Eq. (2) where strain pizoelectric constant m/v electric field the compliance at constant field stress Figure 1 (1) PZT unit cell above Curie temperature. (2) PZT unit cell below Curie temperature. [3] Figure 9: Axes definition for tensor notation with the polarization in axis 3 direction. 10 Note that the subscripts, ik, in the piezoelectric constants refer to the relationship of the mechanical and electrical parameter (e.g., d 31 refers to the electric field applied along axis 3 causing a deflection on axis 1, as seen in Figure 9). The following equation can be used to find the change in dimensions after an electric field is applied: Eq. (3) where change in dimension along L stress initial length along L electric field V/m piezoelectric constant m/v Figure 10 - Expansion and contraction due to an applied electric field. 10 See Figure 10 for the graphical representation of the contraction and expansion when an electric field is applied. 10

11 Piezoelectric Micropumps There are substantial benefits in using piezoelectric actuation in micropumps. The first is the unlimited theoretical resolution in piezoelectric transducers. In reality, noise from the applied electric field, electromagnetic interference, mechanical design and mounting flaws limit the resolution. Nonetheless, sub-nano resolutions are achievable using commercially available transducers. Another benefit is the lack of moving parts in the classical sense. In other words, there are no sliding or rotational parts to cause frictional wear, which helps these micropumps last longer. This is especially critical in the case of drug therapy where reliability is highly valued. In general, pumps can be classified into two categories, positive displacement and rotary. Although piezoelectric actuation can be incorporated into both types, there has been more research conducted with positive displacement pumps due to simplicity in design. Positive displacement pumps, also known as reciprocating pumps, involve fluid flow into a pump chamber through an inlet valve, and flow out through an outlet valve by compression of the fluid, by means of a moving surface. This type of pump usually results in high pressure but low flow rate. Reciprocating pumps can in turn be categorized into diaphragm or piston pumps. This paper will focus on diaphragm pumps due to the extensive work already completed in designing and testing these pumps. 13 Diaphragm pumps force fluid into the pump chamber by expanding the chamber itself and contracting the pump chamber to force fluid out. Note that this will result in a pulsating characteristic with only periodic flow possible. Actuation Mechanism Older designs of diaphragm pumps use cantilever valves to control the direction of fluid flow. Figure 11 shows such a design. The PZT plate utilizes the lateral piezoelectric effect to contract and expand the volume of the pump chamber. When the chamber expands, fluid rushes through the inlet valve while the outlet valve is forced shut. When the chamber shrinks, the pressure forces the inlet valve to close while fluid is forced out through the outlet valve. The arrows in Figure 11 show the direction of flow. The disadvantages of using cantilever valves or other moving valves are the potentials for blockage due to solid particles, and fatigue/wear on the moving parts that will affect the flow rate over time. Figure 11 - Diaphragm pump using cantilever valves 14 11

12 More recent studies have focused on valveless construction. The main difference between the two types is the mechanism that controls flow. The basic design of a valveless piezoelectric diaphragm pump is shown in Figure 12. Figure 12: (a) Top view and (b) side of a valveless diaphragm pump 15 A nozzle is defined with a decreasing cross-sectional area with the direction of flow, and a diffuser is defined with an increasing cross-sectional area with the direction of flow. The pressure loss is much lower through the diffuser direction relative to the nozzle direction. This causes less resistance for flow through the diffuser and creates a net flow in that direction. In supply mode, the PZT disc will deflect the pump membrane to expand the pump chamber. The inlet acts as a diffuser and the outlet acts as a nozzle resulting in a net flow into the pump chamber through the inlet. In pump mode, the pump membrane deflects to decrease the pump chamber volume. The inlet acts as a nozzle and the outlet acts as a diffuser resulting in a net flow out through the outlet. This mechanism allows for directional flow despite allowing some flow in both directions. Microfabrication Typical micro-machining techniques can be used to make micropumps. A design such as Figure 11 uses three silicon wafers. Layers #1 and #2 are identical to each other but assembled front to back. The canonical shapes in these two layers are anisotropically etched using KOH. The cantilevers are made the same way but with a B + anisotropic etch stop. The cantilevers can also be made with anisotropic etching and a cantilever-shaped mask. Layer #3 is made by timecontrolled KOH anisotropic etching with a LPCVD silicon nitride mask. Next, an electrode ink, such as a gold cermet, is printed on top of the membrane, dried, and heated. Koch et al. then uses another ink consisting of 95% PZT-5H powder and 5% leadboronsilicate powder binder to print the piezoelectric transducer on the gold cermet electrode. 14 An electric field of 3 MV/m is applied to the PZT at 130 C for 24 h to polarize and to make it an actual piezoelectric transducer. A final layer of gold cermet is printed on the PZT, dried, and heated to complete the micropump. The fabrication of valveless diaphragm pumps varies depending on each design. Usually, the pump chamber consists of a cylindrical volume with a thin membrane either on the top or both top and bottom. The cylindrical volume can be made by precision turning 17 or deep reactive ion etching (DRIE). 18 The membranes can be any material such as aluminum, copper and even glass. 12

13 It is usually bought from a supplier due to the difficulty in manufacturing consistent thin membranes (<1 mm). The piezoelectric actuators are also usually bought from a supplier but can be cut to size by excimer laser machining. The actuators are then bonded by conductive epoxy glue. 17 The diffuser/nozzles are laser machined and the inlet and outlets are created by anisotropic KOH etching. Governing Equations The pressure loss coefficient through a diffuser/nozzle is given by Eq. 4. Eq. (4) where pressure loss coefficient Δ presure drop in direction of diffuser or nozzle fluid density mean velocity of fluid flow According to work completed by Olsson et al., the pressure loss coefficient is the sum of the pressure drops of the three separate parts of the diffuser/nozzle. 16 These parts are defined as the sudden contraction at the entrance, the expansion at the exit, and the gradual expansion or contraction along the length of the element (see Figure 13). Figure 13: Close-up side view of the diffuser/nozzle element. Note that it depends only on the size opening, W, the length, L, and the diffuser angle, θ

14 Therefore, the pressure loss coefficients in the diffuser and nozzle direction are given by the following:,,, Eq. (5),,, Eq. (6) where A 1 = narrow cross-sectional area A 2 = large cross-sectional area The diffuser efficiency is the ratio of the pressure loss coefficient in the nozzle direction to the coefficient across the diffuser direction as shown in Eq. 7. Eq. (7) Eq. 7 says that if the pressure loss coefficient is larger in the nozzle direction than in the diffuser direction, then there is net flow from the inlet (i.e., η>1). The governing equation for the transverse deflection of the pump membrane shown in Figure 5 is given by the following 15 : Eq. (8) where transverse deflection of pump membrane Young's Modulus membrane thickness Poisson ratio membrane density time piezoelectric actuating force dynamic pressure of fluid on membrane The analytical solution to Eq. 8 is difficult to find due to a couple of reasons. The first is the dynamic flow of the fluid making it non-steady-state. The second is the coupling effects between the piezoelectric actuator with the membrane, and the membrane with the fluid. Therefore, a numeric analysis must be applied for practical reasons and is discussed in the next section. Variables and Optimization Due to the difficulties in developing an analytical solution for Eq. 8, many of the papers referenced thus far have utilized finite element analysis along with computer software to simulate a piezoelectric micropump. One of the popular software used is ANSYS, which allows the user to conduct simulations from a CAD model of the pump. An example of a micropump in a finite element model using ANSYS is shown in Figure 14. When simulating a model in ANSYS, the user must define all parameters such as membrane material, thickness of membrane, input voltage and its frequency, etc. 14

15 Figure 14: Mesh model in ANSYS 17 Some researchers focus on the displacement or deflection of the membrane as the important response. In turn, the area under the maximum deflection is the stroke volume (Figure 16). Figure 15 shows some of the simulation results from Mu et al.. 17 It turns out that the best choice of material for a membrane is aluminum compared with the other materials simulated (Figure 16.a). Deflection also increases as the thickness of the membrane decreases, as expected. Therefore, it is best to choose the thinnest possible membrane for maximum flow rate within the materials limits of elastic deformation (Figure 16.b). One surprising result is that the thicker piezoelectric element does not always lead to larger membrane deflections (Figure 16.c). Micropump designers must take this into account to maximize the use of their piezoelectric actuator. The last factor considered is the input voltage and this varies as expected (Figure 16.d). However, not all applications allow for high voltages, such as implantable micropumps, but the performance of the pump can now be predicted relative to the voltage. Figure 15 - Area under the membrane deflection (Uz) is equal to the stroke volume 17 15

16 (a) (b) (c) (d) Figure 16: (a) Membrane deflection vs. membrane position for various materials. (b) Membrane deflection vs. membrane thickness for varying materials. (c) Membrane deflection vs. piezoelectric actuator thickness. (d) Membrane deflection vs. input voltage

17 After all desired parameters are defined, the maximum stress in the finite element model during the simulation must fall below some stress limit, which is a factor of the material used, such as the von Mises criterion. ANSYS can produce such a model with color-coded regions displaying the maximum stress during actuation (Figure 17). Figure 17: Maximum stress model from ANSYS 17 For drug delivery, stroke volume is not as important as flow rate. The main difference is that flow rate also depends on the frequency of the pump cycle or resultantly, the frequency of the input voltage. Cui et al. take a look at how each factor affects the flow rate. 13 They found very similar qualitative results to Figure 8.a-d allowing membrane deflection to be replaced by flow rate. COMPARING PIEZOELECTRIC TO THERMO-PNEUMATIC Piezoelectric actuating micropumps have very distinct advantages over competing actuating mechanisms. The first is the reliability due to the lack of rotational and frictional wear. Another advantage is the ability to control flow more precisely than most mechanisms due to the high resolution that can be achieved through the piezoelectric transducer. Much research has already gone into using piezoelectric actuation, including multiple simulations allowing designers of micropumps to predict the performance based on design parameters. Various flow rates have been achieved as a result of understanding the effects of these parameters. Cui et al. have achieved flow rates between µl/min, Koch et al. achieved µl/min, and Wan 17

18 et. al achieved 900 µl/min. 13,14,18 Unfortunately, piezoelectric actuation also has drawbacks. From the various designs, the voltage required to achieve the flow rates mentioned varied from 10 to 500 V. Depending on the application, this may or may not be acceptable. Also, applications requiring constant flow will not be able to use piezoelectric pumps since they have a pulsating characteristic that results in periodic flow. To summarize the designs of micropumps based on thermopneumatic actuation, consider the following advantages and disadvantages. Most thermopneumatic pumps generate a large membrane displacement per stroke but have limited stroke frequency due to slow cooling periods. 9 This fact implies that fatigue problems are likely to be less of a problem for these lowfrequency thermopneumatic pumps. On the other hand, the fact that this mechanism generates heat by definition will unavoidably warm the fluid to some extent. The simple structure of a thermopneumatic pump (substrate, heater material, air space, membrane) allows for compact designs with small actuation chamber volume. Furthermore, its simple design facilitates nonsilicon materials such as PDMS and easier fabrication that involves only a few steps and no temperatures above 100 C. However, these pumps require several manual steps such as peeling off layers after curing, placement of PDMS layers before bonding, and punching inlet/outlet/chamber holes. CONCLUSION Micropumps have potential for use in drug delivery and in lab-on-a-chip applications. The two actuating mechanisms studied in this paper, piezoelectric and thermopneumatic, are particularly interesting because of the large amount of work already put into studying them and their contrasting capabilities. Creative thinking including valve-less and membrane-less thermopneumatic designs continues to aid researchers on the quest for low cost, but high performance micropumps. Despite the complicated mathematics in micropumping and microfluidics, the combination of finite element analysis and computer software, such as ANSYS, has allowed users to predict the performance of piezoelectric micropumps quite easily using numerical solutions. Still, choosing one actuation mechanism over another is highly dependent on the application. For example, if low power consumption or low voltage is required, thermopneumatic pumping would easily be the better choice. Another case is if very precise fluid output or dosage is required, thermo-pneumatics would again be the better choice due to smaller flow rates and lower frequencies. However, if higher voltage is acceptable and maximizing the flow rate is the goal, then piezoelectric actuation should be used. 18

19 REFERENCES 1. Sue, C; Tsai N. Sensors and Actuators A Woias, P. Sensors and Actuators B Jeong, O; Park, S; Yang S; Pak, J. Sensors and Actuators A Yoo, J; Moon, M; Choi, Y; Kang C; Kim, Y. Microelectronic Engineering Van de Pol, F; Van Lintel H; Elwenspoek M; Fluitman JHJ; Sensors and Actuators A Jeong O; Yang S. Sensors and Actuators Yoo, J; Choi Y; Kanga, C; Kim Y. Sensors and Actuators A Jun D; Sim W; Yang S. Sensors and Actuators A Nguyen, N; Huang, X; Chuan, T. Transactions of the ASME Espinosa, H. D., Piezoelectricity, ME381 Introduction to MEMS course notes, Last accessed: Dec Cui, Q. F., Liu, C. L. and Zha, X. F., "Simulation and optimization of a piezoelectric micropump for medical applications," Int. Journal of AMT, Koch, M., Harris, N., Evans, A.G.R., White, N.M., Brunnschweiler, A., A novel micromachined pump based on thick-film piezoelectric actuation, Solid State Sensors and Actuators, TRANSDUCERS '97 Chicago., 1997 International Conference on Volume 1, June 1997 Page(s): vol Cui, Q. F., Liu, C. L. and Zha, X. F., Study on a piezoelectric micropump for the controlled drug delivery system, Microfluid Nanofluid Olsson, A., Stemme, G., and Stemme, E., Diffuser-element design investigation for valve-less pumps, Sensors and Actuators A: Physical 57. Issue 2, Pages , November Mu, Y. H., Hung, Y.P., and Ngoi, K. A., Optimisation Design of a Piezoelectric Micropump, Int J Adv Manuf Technol Wan, Z. L., Wu, D. M., Cruz, D., Lazarev, A., Piezoelectric Micropump For Drug Delivery, UCLA, EE Department,

20 BIOGRAPHIES Ken D Aquila is a 4 th year undergraduate at Northwestern University s McCormick School of Engineering and studying Materials Science and Engineering, Ken is currently engaged in a senior research project involving collaboration between the Dravid Research group at Northwestern University and NanoInk. Inc. in Skokie, IL. The project s goal is to develop a write-able tin oxide sol-gel and determine suitable parameters for optimal and reproducible dip-pen nanolithography (DPN). During previous research experiences, Ken has used softebl for fabricating nanostructures and has gained proficiency with the associated characterization techniques such as SEM, TEM, and AFM. Sean Tseng will receive his B.S. in Materials Science and Engineering at Northwestern University in He is currently a Co-op Engineer at ITW Technology Center in the Advanced Technology department located in Glenview, IL. 20