Continuous On-Chip Micropumping for Microneedle Enhanced Drug Delivery

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1 Biomedical Microdevices 6:3, 183±190, 2004 # 2004 Kluwer Academic Publishers.Manufactured in The Netherlands. Continuous On-Chip Micropumping for Microneedle Enhanced Drug Delivery Jeffrey D. Zahn, 1,2,3 * Ajay Deshmukh, 2,4 Albert P. Pisano, 2,4 and Dorian Liepmann 2,3,4 1 Department of Bioengineering, Pennsylvania State University 2 Berkeley Sensor and Actuator Center, University of California at Berkeley 3 Department of Bioengineering, University of California at Berkeley 4 Department of Mechanical Engineering, 497 Cory Hall, University of California at Berkeley, Berkeley, CA jdz10@psu.edu Abstract. Microneedles are promising microfabricated devices for minimally invasive drug delivery applications.needles can be integrated into a variety of devices.however, any portable drug delivery device with integrated microneedles will need an equally compact means to deliver therapeutics.this work presents microneedles integrated with an on-chip MEMS positive displacement micropump for continuous drug delivery applications.the generation and collapse of thermally generated bubbles with ow recti ed by directional check valves are used to achieve net pumping through the device.visualization methods have observed net ow rates of water out of a microneedle at approximately 2.0 nl/ s with a pressure of 3.9 kpa. In addition, continuous pumping was achieved for more than 6 hours with the heaters actuating for over 18 hours (15,000 cycles) without failing. Key Words. microneedle, micropump, drug delivery Introduction In modern medical applications, there is a need for very small hypodermic needles that are economical to fabricate. Currently the smallest needles available, stainless steel 30 gauge needles, have a 305 mm outer diameter with a wall thickness of 76 mm. Traditional machining methods make it unfeasible to create needles with a diameter less than 300 mm. Microneedles (Zahn et al., 2000; Talbot and Pisano, 1998; Talbot, 1999; Lin et al., 1993; Chen and Wise, 1994; Lebouitz and Pisano, 1998; Brazzle et al., 1999; Brazzle et al., 2000; Papautsky et al., 1997; Stupar and Pisano, 2001; Henry et al., 1998; McAllister et al., 1999; Stoeber and Liepmann, 2000a) on the other hand can be almost any size and geometry since they are de ned lithographically. Microneedles have been fabricated from many materials including silicon (Zahn et al., 2000; Talbot and Pisano, 1998; Talbot, 1999; Lin et al., 1993; Chen and Wise, 1994; Lebouitz and Pisano, 1998), metals (Brazzle et al., 1999; Brazzle et al., 2000; Papautsky et al., 1997), and polymers (Stupar and Pisano, 2001). There are two microneedle designs commonly used; micro hypodermic injection needles (such as in this work) where the microneedle is fabricated in the plane of a silicon wafer (Zahn et al., 2000; Talbot and Pisano, 1998; Talbot, 1999; Lin et al., 1993; Chen and Wise, 1994; Lebouitz and Pisano, 1998; Brazzle et al., 1999; Brazzle et al., 2000; Papautsky et al., 1997; Stupar and Pisano, 2001), and arrays of many short microneedles normal to the plane of the wafer surface (Henry et al., 1998; McAllister et al., 1999; Stoeber and Liepmann, 2000a, b). Microneedles are designed to be high performance minimally invasive conduits, through which drug solutions may pass into the body. In order to be minimally invasive, the needles are designed to be as small as possible. Needles are also designed to be extremely sharp, with sub-micron tip radii. This allows the needles to be effectively inserted into the skin. The stress on the skin is inversely proportional to the area over which the force is applied. Therefore as tip radii decreases the stress imposed at a constant force increases and allows lower forces to be used for needle insertion. In addition, the small size of the needles cause less compression of the tissue as needles are inserted which leads to less compression of pain receptors and a decrease in insertion discomfort. The small size of the microneedles also decreases the chance that the needle will be inserted close to a pain receptor. Since microneedles allow such compact, portable and precise delivery of therapeutics, they are well suited for any drug delivery regime in which a continuous infusion is preferred to a bolus injection. Microneedles are extremely attractive for delivering therapeutics in a portable intravenous drip style fashion, such as the continuous delivery of insulin to a diabetic patient (Liepmann et al., 1999). The needles would be integrated into a short-term drug delivery device capable of *Corresponding author. 183

2 184 Zahn et al. Fig. 1. Proposed integrated drug reconstitution and delivery system. delivering therapeutics intradermally for about 24 hours (Figure 1). Continuous drug delivery may be used to increase drug effectiveness and lower side effects for a large range of therapeutics. However, any microneedle enhanced drug delivery will need a compact means to deliver the correct drug dosages. Microfabrication technology allows tighter control over physical parameters ( ow rates, pressure, for example) for the precise delivery of concentrated drug solutions in order to meet dosage speci cations. By integrating microneedles with planar micropumps, very tight control over injection ow rates at given drug concentrations can be achieved. In addition, a more concentrated drug solution may be used at very low ow rates and administrated as needed by the patient. Thus, the drug concentration in the body may be controlled, to achieve either a constant or time varying drug concentration pro le in the body as needed by a patient. In addition, such a device can also be used for sample collection for analysis. The pumps may be used for uid extraction through the microneedle simply by reversing the valve directions so the needle extracts uid from the body into the device rather than delivering it from the device to the body. Thus, an integrated device could be fabricated to determine glucose levels for diabetics. In this scenario, one needle/pump system could sample interstitial uid to determine glucose level while a second pump could deliver insulin in a controled manner as needed by the patient Microneedles have a potential advantage over other approaches to transdermal drug delivery such as electroporation (Prausnitz, 1995), ultrasonic delivery (Kost, 1993), or chemical modi ers/enhancers (Smith and Malbach, 1995) all of which rely on decreasing the permeation barrier of the stratum corneum, the outermost layer of the skin. Microneedles mechanically penetrate the skin barrier and allow the injection of any volume of uid over time. Microneedles have the ability to be precisely inserted to inject therapeutics any particular distance below the stratum corneum. This allows precise localization of a high concentration drug solution in order to obtain effective absorption into the bloodstream or to stimulate particular clusters of cells in or near the skin. Therefore, the drug delivery does not depend on transient delivery of therapeutics across the skin. The delivery is independent of the drug composition and concentration and merely relies on the subsequent drug absorption into the bloodstream, which occurs at a much faster rate than permeation of a solution across the skin. This also allows complex drug delivery pro les. Since drug is actively injected into a patient the dosage may be varied with time. In addition, by employing multiple needles or effective uid control with mixing of solutions (Deshmukh, 2001; Deshmukh et al., 2000; Evans, 1999), multiple drugs may be injected simultaneously speci c to a patient's personal needs. Needles may also be used to transdermally sample body uids for analysis. The extreme miniaturization of uidic devices enables portable devices for personalized medicine allowing continuous metabolite monitoring with drug delivery in response to metabolite levels. This has tremendous advantages over existing technologies. It allows patients more freedom in their treatment since they are no longer dependent on a bolus injection or a short period of intravenous drug delivery with little or no therapy between treatments. It also allows a lower drug dosage to be injected over a longer period of time to maintain a constant blood concentration. By maintaining a constant blood concentration, side effects associated with a high concentration bolus injection may be reduced. This research assumes that microneedles are capable of being integrated into autonomous, planar microdevices capable of sampling biological uids, analyzing the uid, performing self cleaning and calibration and then delivering therapeutics in response to metabolite levels (Liepmann et al., 1999). Any portable drug delivery or analysis module will need to operate in a compact manner. A microfabricated delivery module (Zahn et al., 2001) with integrated microneedles also allows more freedom in a patient's drug regiment by freeing the patient from bulky belt worn pumps with a large stainless steel needle protruding into the abdomen. Such a device could be worn as a patch style device on an arm or the abdomen, so that the patient hardly realizes they are wearing it. In order to realize such a system, however, many micro uidic components such as micropumps and microvalves for uidic control must also be realized (Gravesen et al., 1993). Although there has been ample research on such components (Forster et al., 1995; Olsson et al., 1996; Grosjean and Tai, 1999; Tsai and Lin, 2000),

3 On-Chip Micropumping for Microneedle Enhanced Drug Delivery 185 one other goal is to develop planar components in order to limit the complexity of the completed device. Thus, all uidic components may be de ned in a single lithography step, with a minimum of assembly. This work attempts to integrate such components into a planar micro uidic device capable of sophisticated uid control for medical analysis and drug delivery in response to metabolite levels. Design and Fabrication Planar micropumps have the advantage of being easily integrated with other planar uidic components. The pumps consist of a bubble chamber and two directional check valves. In order to obtain effective pumping a time varying electrical input must be supplied to the device. The pumping occurs using a cyclic heating cycle. Vapor bubbles are created by the heat dissipated by polysilicon resistors on a transparent quartz substrate. Vapor bubbles (literally steam bubbles) are generated by heat dissipated when an electric current is run through the resistor. Bubbles have long been recognized as thermopneumatic transducers to generate force (Lin et al., 1991) or for use as a pump actuation source (Grosjean and Tai, 1999; Tsai and Lin, 2000). When the electric current is interrupted, the vapor will recondense and the bubble will collapse. When a vapor bubble is created it acts as a thermopneumatic piston and drives uid out of the pumping chamber. The directional check valves are free oating silicon pieces that are moved by viscous drag. Flow in a positive direction will open the valve, whereas ow in the opposite direction will drag the valve closed. By cycling bubble generation and collapse, a net pumping action occurs. Microneedles are integrated by adding them to the uid ow path (Figure 2). Microneedles are formed by a previously described polysilicon molding process (Zahn et al., 2000). For ease of handling some microneedles were designed with a large base that may be placed into a large etched cavity to Fig. 2. Schematic of an integrated micropump/microneedle device. The top shows a closeup of a planar free oating directional microvalve. couple the needle ow path to that of the micropump. A double polished silicon wafer ( p-type h100i single crystal) is used as a starting material as shown in Figure 3. The wafer is wet oxidized at 1,000 C to grow 1 mm of thermal oxide. The needle shape is patterned using standard lithographic techniques and the patterned oxide is etched with reactive ion etching (RIE). The wafer is then coated with 0.3 mm of low-pressure chemical vapor deposition (LPCVD) low stress silicon nitride. The backside of the wafer is aligned and patterned to the needle design using a Karl Suss contact mask aligner with backside alignment capabilities. The silicon nitride and oxide lms are then etched away using RIE. Afterwards, through holes are etched with potassium hydroxide (KOH). The nitride is then removed in phosphoric acid at 175 C. The needle mold is then etched using deep reactive ion etching (DRIE) in a surface technology systems (STS) etcher. The resulting trench is typically between 100 and 125 mm deep. The wafer is then wet oxidized at 1,000 C to decrease surface roughness. Afterwards 2 mm of phosphosilicate glass (PSG) is deposited onto the mold wafer. A second bare silicon wafer is coated with 2 mm of PSG. These two wafers are pressure bonded at 1,000 C in a N 2 atmosphere. After bonding, 4 mm of polysilicon is deposited onto the mold wafers at 580 C. The molds are then annealed at 1,000 CinaN 2 atmosphere. The polysilicon is annealed between depositions to alleviate lm stress. The deposition and annealing is repeated until the desired thickness is reached, typically 15±20 mm. The polysilicon deposits conformally onto the outside of the wafer and inside the mold through the KOH etched through holes. After deposition, the external polysilicon is removed by RIE. The needles are then released in concentrated HF overnight. The resulting structures are closed uid passages with inlets and outlets on the side of the needle rather right at the tip (Figure 3). The micropump is fabricated using silicon on insulator (SOI) and quartz dice. The SOI wafer has a 2 mm thick buried oxide layer and a 75 mm thick device layer. Two sequential deep reactive ion etches (DRIE) are performed; the rst penetrating through the wafer to form uid access holes, while the second etch only etches through the device layer to create the micropump and microchannels. A 1.3 mm thick wet silicon dioxide layer provides the mask for the microchannel etch while a 9.5 mm thick layer of photoresist serves as the mask for the uid access hole etch. After the DRIE processes, the valves are almost completely freed from the substrate in 5 : 1 BHF. They are then placed in H 2 O 2 to generate a thin oxide layer on the bottom of the valve. Afterwards, a probe tip is used to gently free the valves (Figure 4). The quartz wafer has 3,000 A Ê n doped polysilicon that is patterned into a resistive heater (Figure 5) and 1,000 A Ê

4 186 Zahn et al. Fig. 3. (Top) Microneedle fabrication process ow. (Bottom Left) SEM of a microneedle mold. (Bottom Right) SEM of a representative microneedle. sputtered platinum for electrical access and is passivated with silicon oxide and/or silicon nitride, leaving openings for the electrical connections. In order to bond the two dice together, a low viscosity epoxy (Epotek 301) is spun onto the quartz die at 10,000 rpm. The two dies are aligned and bonded in a ip-chip bonder at pressures between 100±300 kpa per 1 cm 2 die. The epoxy bonds and seals the dies but also traps the valve bodies. An oxygen plasma (200 W in 600 mtorr of O 2 ) is used to remove the epoxy in the uid channels. Through holes allow the plasma to access the uid channels. The oxygen rst removes the epoxy on the top of the channels and then moves slowly outward from the channel. The plasma etching is timed to free the valves without penetrating far from the channel edge. The valves are freed, but suf cient epoxy remains outside the channels to bond and seal the device. This process is described in detail by Deshmukh (2001). A descriptive cross section of the micropump is shown in Figure 6; it does not correspond to any actual cross section in the device but is a combination of the heater and valve structures. A seat for the microneedle is de ned using DRIE. The microneedle is placed into the seat and sealed using a two-part epoxy.

5 On-Chip Micropumping for Microneedle Enhanced Drug Delivery 187 Fig. 4. A free oating, directional micro uidic check valve. Flow to the bottom right of the picture will open the valve, whereas ow towards the top left closes it. Fig. 5. A polysilicon heater over an etched bubble chamber. Current is supplied through the hexagonal polysilicon lines. When resistive heat is applied to the heater water evaporates into a bubble which acts as a thermopneumatic piston to pressurize uid in the channel. Fig. 7. A working micropump. (Top) Flow channel with a trapped bubble. (Middle) Motion of the tracking bubble after power has been supplied. (Bottom) Acceleration of the bubble as power is maintained. Fig. 6. Description cross section of the device. Experimental Details Successful integration of microneedles with micropumps has been accomplished. As shown in Figure 7, a working micropump has been developed. A tracking bubble in the series of pictures can be used to visualize uid ow. The tracking bubble is stationary until the heater is activated. As the heater dissipates heat, a vapor bubble is formed that acts as a thermopneumatic piston and uid within the microchannel begins moving towards the outlet. The uid velocity in the channel is not constant since the bubble generation and collapse creates a time oscillating ow pro le. The velocity of the bubble shown is estimated to be 15 mm/sec, a peak velocity for the pump. Once dies are assembled they are wirebonded onto a circuitboard package to allow the application of

6 188 Zahn et al. Fig. 8. (Left) Fluorescein at the microneedle outlet when no bubble is present in the bubble chamber. (Middle) Change in radius of curvature and collection of uid at the outlet when the pump is actuated. The uid has been pushed out through the outlet and collects on the needle surface as can be seen by the increase in uorescence. (Right) Change in the radius of curvature when the bubble collapses. The uid is now pulled back into the needle so uorescence collects at the corners of the needle outlet. The side view of the uid bubble at the outlet is shown below the pictures. It shows how the radius of curvature changes between bubble generation and collapse. electrical power to the heaters. As stated before, a time varying electrical input is required to obtain effective pumping. The heaters can create bubbles at different rates by varying the heater voltage and time applied. In these experiments, the various bubble creation/collapse rates have been achieved through a computer control mechanism. A power supply is set to a given DC potential, and supplied to the device through a relay. The relay is activated by a HP-VEE output through a D/A circuitboard. The HP-VEE interface controls the time power is supplied to the device, while the power supply controls the voltage. In order to visualize uid delivery from the microneedle using an actuating micropump, the outlet of the needle was visualized using an inverted epi uorescent microscope (206 objective 0.40 NA) and appropriate uorophores. Due to the planar nature of the devices, the microneedle outlet was positioned normal to the light path. This makes ow quanti cation more dif cult, since uid at the needle outlet is normal to the light path. Flow visualization was also dif cult when the needle was submerged under water. In Figure 8 is shown the radius of curvature of 1% uorecein dye in water at the outlet of the microneedle changes due to surface tension. As the thermopneumatic bubble is generated, the bubble pushes uid out of the needle, but surface tension at the outlet constrains the uid from being pumped out of the needle but the pressure causes the radius of curvature at the outlet to increase. Then as the thermopneumatic bubble collapses after power is no longer supplied to the heater, the collapsing bubble pulls uid back into the needle before the microvalve is completely seated. This collapsing bubble causes the uid surface to become concave until pressure is allowed to equilibrate again and the surface returns to its convex conformation. At the microneedle outlet no net pumping is seen, since the surface tension from the small radius of curvature opposes the pressure generated by the vapor bubble. The radius of curvature at the outlet is approximately the same as the channel depth (37 mm). Therefore the pressure generated in the bubble (approximately 3.9 kpa given by 2s/r where s is the surface tension of water and r is the minimum bubble radius of curvature) is opposed by the outlet free surface so no net uid ow out of the microneedle is seen. Surface tension is diminished by allowing a drop of uid to collect around the needle end so that uid ow out of the needle is seen. The most successful demonstration of uid delivery out of the microneedle by micropumping was with the application of 40 V to the heater. The heater was on for 2 seconds during which time the bubble grew and off for 3 seconds to allow bubble collapse. This corresponds to a power dissipation of 0.5 W. Under these conditions, continuous pumping was observed for a period of 6 hours. The heater was left functioning for over 15,000 cycles. As shown in Figure 9b, a net pumping is seen when the ow is visualized by uorescent beads. When power

7 On-Chip Micropumping for Microneedle Enhanced Drug Delivery 189 Fig. 9a. A series of four video frames (1±4) when the pump is off. The beads stay in focus for all four frames showing the slow background ow. The black rectangles indicate the boundaries of the needle outlet. The arrows point to three representative beads in the series of frames. The image has been inverted for easier visualization of the beads. Fig. 9b. A series of four frames (1±4) once the pump has been turned on. Notice that each frame is dissimilar showing the increased bead velocity and ow rate when the pump is on. The arrows show one representative bead that was tracked showing a large displacement, D, between frames 3 and 4. The white rectangles represent the needle outlet. The image has been inverted for easier visualization of the beads. is not applied a background ow of approximately 0.15 nl/s is present due to a head pressure from uid in the lling lines. However, once the heater is turned on, a net pumping of 1 nl/s can be observed. Flow rates are estimated by estimating the volume ux out of the needle using the beads. The microscope objective has a depth of eld of 5 mm, and the needle outlet is mm 2. Thus when the pump is off the beads can be recorded and seen to stay in focus for 0.16 second. The bead velocity is therefore depth of field time in focus : A ow rate can be estimated as Q ˆ outlet area 6 bead velocity : As shown in Figure 9a, when the pump is off, each video frame looks very similar to each other. However, when power is applied to the pump, each frame in sequence is quite different from each other. When power is supplied to the heater few beads stays in focus for more than one frame. When a bead can be seen in more than one frame (which is rare and usually the bead is much more out of focus between frames) a larger displacement can be seen between frames and a linear velocity may be estimated. This implies a much higher ow rate than the background ow and net pumping is observed. When heater power is turned off and the bubble collapses, the initial collapse and pulling on the uid completely opposes the background ow in the uid. However, once the bubble is gone, the background ow starts up again. Based on these estimates of ow rate 2 nl of uid will be delivered for each 5 second cycle leading to a delivery of 1.44 ml/hour if the pumps are operated continuously. Conclusions This work presents a compact MEMS based positive displacement system to deliver drugs through minimally invasive microneedles. Microneedles could also be integrated with micromixers (Deshmuhk et al., 2000). The systems would then be a portable system capable of reconstituting lyophilized drug, and dispensing it at the correct dosage as needed by a patient. The ow rates observed in this study were about an order of magnitude lower than the 12.5 nl/s previously observed (Deshmukh et al., 2000). This is more than likely due to an increased ow resistance from the microneedle. The microneedle contributes more viscous losses to the pump. In addition,

8 190 Zahn et al. the small spacing between the needle and the seat will contribute to a very large drag on the uid which limits ow rates. A better coupling of the needle to a chip could minimize these losses. Acknowledgments The authors would like to thank the members of the Liepmann/Pisano labs for technical discussions and assistance. This project has been funded by the DARPA microflumes program and by Becton± Dickinson Technologies. All devices were fabricated at the U.C. Berkeley Microfabrication Facilities. References J. Brazzle, I. Papautsky, and A.B. Frazier, IEEE Engineering in Medicine and Biology Magazine 18(6), IEEE, 53±58 (Nov.±Dec., 1999). J. Brazzle, D. Bartholomeusz, R. Davies, J. Andrade, R.A. Van Wageman, and A.B. Frazier, Proceedings 2000 Solid State Sensor and Actuator Workshop, Hilton Head, S.C., 199±202 (June, 2000). J. Chen and K.D. Wise, Solid State Sensor and Actuator Workshop, Hilton Head, S.C., 256±259 (June, 1994). A.A. Deshmukh, Continuous Micro uidic Mixing Using Pulsatile Micropumping (Ph.D. Thesis, University of California, Berkeley, CA., 2001). A.A. Deshmukh, D. Liepmann, and A.P. Pisano, Proceedings 2000 Solid State Sensor and Actuator Workshop, Hilton Head, S.C., 73±76 (June, 2000). J.D. Evans, Planar Micro uidics: Towards Large Scale Integration of Channels, Pumps, Valves, and Fluid Mixers in Microelectromechanical Systems (MEMS) (Ph.D. Thesis, University of California, Berkeley, CA, 1999). F. Forster, R. Bardell, M. Afromowitz, N. Sharma, and A. Blanchard, Proceedings of the ASME Fluids Engineering Division, 1995 IMECE, 234, 39±44 (1995). P. Gravesen, J. Branebjerg, and O.S. Jensen, J. Micromech. Microeng. 3, 168±182 (1993). C. Grosjean and Y.C. Tai, 1999 International Conference on Solid-State Sensors and Actuators (Transducers'99), Sendai, Japan (June, 1999). S. Henry, D.V. McAllister, M. Allen, and M. Prausnitz, Proceedings of the IEEE Eleventh Annual International Workshop on MEMS, Heidelberg, Germany, 494±498 (January, 1998). J. Kost, J. Controlled Release 24, 247±255 (1993). K.S. Lebouitz and A.P. Pisano, Proceedings Microstructures and Microfabrication Systems IV, 194th meeting of the Electrochemical Society, Boston, MA (Nov. 1±6, 1998). D. Liepmann, A. Pisano, and B. Sage, Diabetes Technology and Therapeutics, 1(4), 469±476 (1999). L. Lin, A.P. Pisano, and R.S. Muller, Proceedings 7th International Conference on Solid State Sensors and Actuators (Transducers'93), Yokohama, Japan, 237±240 (1993). L. Lin, A.P. Pisano, and A.P. Lee, 1991 International Conference on Solid-State Sensors and Actuators. Digest of Technical Papers (Transducers'91), 1041±1044 (1991). D.V. McAllister, F. Cros, S.P. Davis, L.M. Matta, M.R. Prausnitz, and M.G. Allen, Proceedings 10th International Conference on Solid State Sensors and Actuators (Transducers'99), 1098±1101 (1999). A. Olsson, G. Stemme, and E. Stemme, Proceedings IEEE Workshop On Micro Electro Mechanical Systems (MEMS'96), 378±383 (Feb., 1996). I.E. Papautsky, J.D. Brazzle, H. Swerdlow, and A.B. Frazier, IEEE International Conference on Engineering in Medicine and Biology Conference, Chicago, IL (October, 1997). M.R. Prausnitz, in Electronically controled Drug Delivery, edited by B. Berner and S.M. Dinh (CRC Press, Boca Raton, FL, 1995). E.W. Smith and H.L. Malbach, Percutaneous Penetration Enhancers (CRC Press, Boca Raton, FL, 1995). B. Stoeber and D. Liepmann, Proceedings of the 1st IEEE-EMBS Special Topic Conference on Microtechnology in Medicine & Biology, Lyon (2000a). B. Stoeber and D. Liepmann, Proceedings of the ASME MEMS Division, 2000 IMECE, 1, 355±359 (2000b). P. Stupar and A.P. Pisano, Proceedings 11th International Conference on Solid State Sensors and Actuators (Transducers'01), Munich, Germany (2001). N. Talbot and A.P. Pisano, Proceedings 1998 Solid State Sensor and Actuator Workshop, Hilton Head S.C., 265±268 (June, 1998). N. Talbot, Polysilicon Micromolding of Closed-Flow Passages for the Fabrication of Multifunctional Microneedles (Ph.D. Thesis, University of California, Berkeley, CA, 1999). J.H. Tsai and L. Lin, Late News Poster Session Supplemental Digest of Solid-State Sensor and Actuator Workshop, Hilton Head S.C., 13±14 (2000). J.D. Zahn, A.A. Deshmukh, A.P. Pisano, and D. Liepmann, 14th Annual IEEE International MEMS-01 Conference, Interlaken, Switzerland, (January 21±25, 2001). J.D. Zahn, N.H. Talbot, D. Liepmann, and A.P. Pisano, Biomedical Microdevices 2(4), 295±303 (2000).

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