Piotto

P Bruschi

M Dei

M Piotto

SINGLE CHIP SENSING OF MULTIPLE GAS FLOWS

published or not. The documents may come

INTRODUCTION

Detection and measurement of fluid flow is a crucial operation in many complex systems. In the field of automotive engines and household heating boilers, precise control of the combustion mixture allows efficiency improvement and reduction of pollutant emission. High resolution sensing of very small gas flows is an important target in the development of more effective ionic and cold gas thrusters for orbit and attitude control in scientific satellites [START_REF] Noci | Development Activities on an Advanced Propellant Flow Control Unit[END_REF]. Fine monitoring of multiple gas and liquid flow is also a requirement for controlling the correct evolution of chemical reaction. In particular, sensitive and miniaturized flow sensor are beginning to be integrated in microreactors [START_REF] Shin | A micromachined thin-film gas flowsensor for microchemical reactors[END_REF][START_REF] De Mas | Scaled-Out Multilayer Gas-Liquid Microreactor with Integrated Velocimetry Sensors[END_REF], a new breed of silicon based devices devoted to handle small amount of reagents with strict temperature and pressure conditions.

It is interesting to observe that flow control is still preferably performed in an open loop fashion, i.e. relying on calibrated actuators such as, syringe [START_REF] Ferrer | A smart multisyringe flow injection system for analysis of sample batches with high variability in sulfide concentration[END_REF] and peristaltic pumps [START_REF] Bu | A new masking technology for deep glass etching and its microfluidic application[END_REF], with no need of flow meters. The reason stands in the high cost, dimensions and insertion loss of traditional flow sensors.

In principle, such limitations could be overcome by using integrated thermal flow sensors [START_REF] Elwenspoek | Mechanical Microsensors[END_REF], exploiting the advantages of silicon processing technologies to reduce costs, size and power consumption. Unfortunately, in spite of the huge variety of prototypes presented in the literature in last two decades, the number of commercial products currently available is still considerably small. One of the reasons is the necessity of developing proper packages, capable of protecting the fragile micromachined structures and, at the same time, allowing the interaction between the micrometric sensing structures and the gas. The method to convey the gas flow to the sensing structure without damaging the bonding wires and introduce as low as possible insertion losses has to be tailored to the particular application, taking into account specification of sealing, sensitivity, range and miniaturization [START_REF] Bruschi | Postprocessing Technologies, Interface Circuits and Packaging Strategies for CMOS Compatible Gas Flow Sensors[END_REF]. Solutions based on placing the whole chip with its bonding wires inside the gas channel [START_REF] Mayer | Influence of design geometry and packaging on the response of thermal CMOS flow sensors[END_REF][START_REF] Bruschi | A Closed-Loop Mass Flow Controller Based on Static Solid-State Devices[END_REF] are not compatible with liquid flows or aggressive gases; furthermore, the channel cross-section cannot be reduced below several mm 2 , so that, handling of very small flows is limited. On the opposite side, the alternative is placing the chip outside the flow channel, relying on a thermally conducting membrane or simply the pipe walls to allow heat exchange between the sensor elements and the fluid. [START_REF] Van Oudheusden | Silicon thermal flow sensor with a twodimensional direction sensitivity[END_REF][START_REF] Mayer | Flow Sensor[END_REF]. In this case, the need of heating a pipe portion several times longer than the chip size results in sensitivity reduction, higher power consumption and longer time transients.

Recently, we have proposed a packaging method based on a PMMA adapter, thermally bonded to the silicon chip [START_REF] Bruschi | An integrated circuit compatible compact package for thermal gas flow meters[END_REF]. The advantage of this approach is the possibility of conveying the gas flow to selected areas of a completely standard chip, with no need of particular bonding pad configurations, chip size or surface flatness.

In this work we describe the application of this inexpensive technique to the fabrication of a single chip flow meter capable of measuring two independent gas flows. The method is applied to a chip designed earlier than the development of the packaging method, so that the placement of the sensing structures was not optimized to include them in separate flow channels. Nevertheless, it will be demonstrated that the method can be applied to sense two gas flow using sensing structures separated one from the other and from the bonding pads only a few hundreds microns.

II. DEVICE DESCRIPTION

The device is composed of two main elements: (i) a sensor chip and (ii) a plastic adapter used to convey the gas flow to the chip surface. The chip has been fabricated using the STMicroelectronics Bipolar-CMOS-DMOS process "BCD6" with 0.35 µm CMOS devices and three metal layers option. The flow sensing structures are classical thermal anemometers [START_REF] Noci | Development Activities on an Advanced Propellant Flow Control Unit[END_REF], formed by a heater (p-doped polysilicon resistor) and two temperature probes (thermopiles) placed upstream and downstream to the heater. Thermal insulation between the thermopiles and the heater is obtained by suspending the three elements over separate dielectric membranes by means of a standard micromachining ISBN: 978-2-35500-006-5 procedure, briefly described in the next section.

A SEM (scanning electron microscope) micrograph of a sensing structure after the post-processing procedure is shown in Fig. 1, where the heater and thermopile conducting layers are not visible due to the upper dielectric layers.

T1 H T2

Fig. 1. Scanning electron micrograph of a sensing structure after the postprocessing procedure. T1,T2 and H indicate the thermopile and heater membranes, respectively.

The output voltage, proportional to the temperature difference between the downstream and upstream thermopile, is related to the flow velocity by a monotonic relationship over a wide velocity range. The sensing structures were equipped with three different thermopile types: p-polysilicon/Al (PA), p-polysilicon/n-polysilicon (PN) and n-polysilicon/Al (NA). Unfortunately most structures, designed using NA thermopiles, exhibited a near zero sensitivity, probably due to the low value assumed by the NA couple in the BCD6 process. On the other hand, the only two different structures, one of PA and the other of PN type, showed a high sensitivity; however, their placement was not favorable for the following packaging phase, due to their small spacing and vicinity of the bonding pad frame. The situation is clearly shown in Fig. 2, where the chip layout with the sensing structure is depicted.

The gas flow is conveyed to the micrometric sensing structures by means of the plastic adapter, applied to the chip surface in order to obtain sealed channels including the sensing structures. The adapter face, which is placed into contact with the chip, is slightly smaller than the pad frame, so that it can be easily applied after bonding of the chip. Trenches of proper shape and position are milled on the adapter face, in order to make the gas flow interact with selected chip areas. The trenches are accessed by means of cylindrical holes perpendicular to the chip surface. The holes, that terminates on the opposite adapter face, can be connected to gas lines by means of capillary pipes. Sealing of the trenches to the chip surface is obtained using a thermal procedure explained in the next section. Two different configurations of the trenches were used in this work: the first one, shown in Fig. 3 (a) consists of two independent U-shaped trenches, including the PA and PP structure, respectively. The second configuration, shown in Fig. 3(b), consists of a single L-shaped trench, including both sensing structures. The first configuration represent the target two channel device while the second one, similar to the device presented in [START_REF] Bruschi | An integrated circuit compatible compact package for thermal gas flow meters[END_REF], is used for reference purposes. In both cases the trenches have a 0.5 mm × 0.5 mm rectangular cross section. The actual position of the trench on the chip surface in the case of Fig. 3 The micrograph on the left shows an enlargement of the chip at the point the sensing structures are placed. The thermopiles (T 1 and T 2 ) and heater (H) are indicated for both structures. The small spacing between the sensing structures imposed to reduce the thickness of the wall separating the two trenches, requiring a fairly good precision in the mechanical fabrication of the adapter. Another unwanted consequence of the structure position and orientation was the necessity to place them practically inside a 90 degrees curve of the flow channel. Considering that the type of sensing structure used in this work is sensitive only to the gas velocity component along the thermopile-heater direction, a sensitivity reduction can be expected with respect to the case of Fig. 3 (b) where both structures are oriented along the gas streamlines.

III DEVICE FABRICATION

The chip was designed with the BCD6 (Bipolar-CMOS-DMOS) process of STMicroelectronics. Thermal insulation of the heaters and thermopile membranes from the substrate was obtained by suspending them on a cavity etched into the bulk silicon by means of a front side micromachining procedure. To this aim, proper openings are previously opened into the dielectric layers in order to access the bare silicon. The opening geometry was defined by means of a 1 µm resolution photolithographic step. This operation is facilitated by the presence of passivation openings purposely introduced in the layout design [START_REF] Bruschi | An integrated circuit compatible compact package for thermal gas flow meters[END_REF]. The substrate silicon is then etched through the openings by means of a TMAH based anisotropic etching performed at 90 ºC, two steps of 40 minutes each, renewing the solution after the first step. The described post-processing approach requires much shorter etching times than the alternative of removing the silicon by means of back-side micromachining.

The dies were glued to 28 pin DIL ceramic packages by means of epoxy resin and bonding of the pads to the package pins is performed. The chip includes also a few electronic blocks that were required to control the chip temperature. This possibility was exploited to optimize the adhesion of the chip surface to the plastic adapter. The adapter was fabricated using a precision milling machine (VHF CAM 100) controlled by a personal computer. The material for the adapter material was chosen considering characteristic of transparency, for alignment purposes, and low glass transition temperature, to facilitate the sealing procedure. Polymethylmethacrylate (PMMA) was chosen as a material for the adapter, considering also its greater hardness with respect to similar plastics [START_REF] Bruschi | An integrated circuit compatible compact package for thermal gas flow meters[END_REF]. Hardness was essential to allow fine milling of trenches with the small separation shown in Fig. 4. Alignment of the adapter to the chip was performed using a guide, schematically shown in Fig. 5(a), consisting in a thin PMMA bar with an opening of the same dimensions of the adapter front face. Firstly, the guide is leaned on the package surface by means of a micrometric displacement stage and coarsely aligned to the chip in order to allow the adapter to be inserted with no risk of hitting the bonding wires. Then, fine alignment, aimed to place the sensing structure into the channels, is performed with the adapter inside the guide. After the alignment, the guide was It is important to observe that this manual procedure can avoided in an industrial process, where it is possible to glue all the chips in a precise position of the package chamber and then simply align the guide to reference marks on the package itself.

At this stage it is not possible to obtain a leak proof connection of the adapter to the chip surface due to the uneven characteristic of the latter. In particular, protrusions of more than 3 µm, deriving from the thick upper metal layers, were measured with a stylus profilometer. Sealing of the flow channels was then obtained maintaining the chip at the PMMA glass transition temperature (110 ºC) for five minutes while applying a force of about 5 N between the chip itself and the adapter. A ∆V BE substrate temperature sensor and a series of power DMOS present on the chip were used to maintain a constant temperature for the duration of the whole sealing procedure. Due to the poor adhesion of PMMA on silicon and silicon dioxide, the structure obtained in this way, though being leak proof, is rather fragile. A robust structure is obtained by carefully pouring epoxy resin inside the package chip chamber through holes in the guide. It is worth observing that, without the thermal sealing procedure, the resin slipped between the chip and the adapter, partially filling the trenches. The final structure of a two channel flow meter is shown in Fig. 5(b).

IV EXPERIMENTAL RESULTS

Characterization of the devices was performed in nitrogen flow by means of a reference line, equipped with a mass flow controller (MKS 1179B) with a 10 sccm (standard cubic centimeters per minute) full scale range. The gas flow was applied to the devices by inserting a stainless steel pipe of 0.7 mm external diameter (syringe needle), into each one of the four upper orifices, in turn. Shifting the needle to one orifice to the other allowed to test the two channel in both flow directions. Before each measurement run, the needle was sealed to the orifice by means of silicone glue, which was easily removed after the measurement run in order to remove the needle and insert it in the next orifice. The thermopile output signal was read by means of a purposely built interface board, including a low noise instrumentation amplifier (AD 620), a second order 10 Hz low pass filter and a constant voltage source to bias the sensor heaters. In all the The response of the two channel sensor is presented in Fig. 6. In details, Fig. 6(a) shows the thermopile output voltage of the channel with the PN structure, while Fig. 6(b) refers to the channel with the PA structure. A significant asymmetry with respect to flow reversal is visible for both channels. To better put in evidence this phenomenon, the linear fit of the data at flow rates less than 1 sccm has been drawn separately for the positive and negative branches of the curves. Inspection of Fig. 6 reveals that the asymmetry is much less important at small flow rates: it can then be ascribed to a non linearity effect, more evident for negative flows. Another aspect is the different sensitivity of the two channels, indicated by the slope of the linear fit reported in the figure distinctly for the negative and positive branch.

To further investigate the actual role of the channel geometry to determine the observed response features, experiments have also been performed with single channel devices. The result for the PA and PN structures is shown in Fig. 7(a) and 7(b), respectively. Note that, in this case, the two thermopile are included in the same channel. The fact that the PN structure still presents a sensitivity nearly 50 % higher than the PA one, clearly indicate that this is an intrinsic property of the thermopile type. It is more important to observe that the single channel sensor is generally more sensitive (almost one order of magnitude) and its response is also much more symmetric. This is a clear consequence of the fact that, for the two channel sensor, it has been necessary to place the sensing structures practically inside a channel elbow. The oblique gas velocity direction with respect to the thermopile-heater line may be the reason of the lower sensitivity of the double channel configuration. The linear fits are calculated for flow rates smaller than 0.8 SCCM, distinctly for the negative and positive branches. The asymmetry, instead, is probably a consequence of turbulences that may arise even at very low Reynolds numbers in the proximity of channel sharp bends.

Finally, it was important to test the cross-sensitivity between the two measurement channels. To this aim, flow rates up to 50 sccm have been injected into one channel while measuring the output signal of the other one. In all the experiments no measurable cross-talking was detected, confirming the excellent sealing of the channels.

V CONCLUSIONS A previously presented technique for conveying gas flows to selected areas of a chip has been improved to allow multiple flow sensing with a single standard silicon chip. The fact that the chip was not designed to this purpose, prevented the placement of the structures into straight segments of the flow channels. As expected, this resulted in sensitivity reduction and increased response asymmetry. Reasonably, the design of a new chip with sensing structures with more favorable placement and orientation would eliminate this problem. However, the effective separation of the two flows, even in this critical configuration, demonstrated that the proposed technique can be successfully applied to probe adjacent areas of standard chips by distinct fluid flows. This suggests that the proposed inexpensive packaging approach can be applied also to the detection of liquid flows, where leakage towards critical chip areas (e.g. bonding pads), has to be strictly avoided.

Fig. 2 .

 2 Fig.2. Placement of the PA and PP sensing structures on the chip layout.

Fig. 3

 3 Fig.3 Perspective view of (a) the proposed device and (b) the prior version.

Fig. 4 .

 4 Fig. 4. Layout of the designed chip with the channel path.

9- 11

 11 April 2008 ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5

Fig. 5 .

 5 Fig. 5. (a) Perspective view illustrating the device assembling and (b) Photograph of the final device.

9- 11

 11 April 2008 ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5 experiments the heaters were supplied by a 2 V constant voltage, corresponding to a power of 4 mW for each sensor.

Fig. 6 .

 6 Fig. 6. Response of the two two-channel sensor after offset removal: (a) channel equipped with the PN structure, (b) channel with the PA structure.The linear fits are calculated for flow rates smaller than 0.8 SCCM, distinctly for the negative and positive branches.

Fig. 7 .

 7 Fig. 7. Response of the two sensing structures placed inside the same Lshaped channel.

ACKNOWLEDGMENT

The authors wish to thank the R&D group of STMicroelectronics in Cornaredo, Milan, for fabricating the chips used in the experiments described in this work. 9-11 April 2008 ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5


Fabien Maury

Arnaud Pothier

Aurelian Crunteanu

F Conseil

Pierre Blondy

RF-MEMS Switched Varactors for Medium Power Applications

published or not. The documents may come   L'archive ouverte pluridisciplinaire

INTRODUCTION

Many researches about MEMS trend to exploit their wellknown and large advantages. Cost and size reduction enables MEMS sensors to be focus on. Some of them have demonstrated their mass production feasibility thanks to their CMOS compatibility. Nowadays, micro actuators are studied to follow the same way in telecommunication domain. Those which are driven by electrostatic actuation are the most investigated. This kind of structures has an easy way to build; a simple difference of potential between two metallic layers is enough to create an attraction force for actuation. The most mature of them have been already commercialized, only tens of volts enable them to switch [START_REF] Rebeiz | RF MEMS: Theory, Design and Technology[END_REF]. Their interesting performances promise to be very suitable for a large kind of control circuits.

For example, MEMS switched capacitors are used for their low series resistances and low-drive power requirements. For wireless communication and particularly radar applications, switched-line phase shifters using MEMS fixed-fixed beams have demonstrated good degree/ loss ratios. Nevertheless, the choice of these MEMS structures is compromised for high and even medium power applications. For microwave signals beyond 1 Watt, most of MEMS switched capacitors fail in hot switching operation. In this operating way, the RF signal is applied continually and, contrary to cold switching, it is not interrupted when the MEMS beams are activated. This is a better manner of working for users who don't have to care about RF signal source control.

The most critical power failure phenomenon appears by this way is called RF latching. In down state position, the electrostatic pressure generated by RF power prevents the beam to recover the up state when the DC actuation signal is stopped. To avoid this event, the mobile structure must have a better stiffness to lead to a restoring force capable of frustrate this electrostatic force [START_REF] Peroulis | RF MEMS switches with enhanced power-handling capabilities[END_REF].

II. DEVICE DESIGN

In the research institute XLIM in Limoges, we have proposed a design of MEMS switched capacitors with an enhanced topology offering a better opposition to RF power forces compared to traditional architecture [START_REF] Palego | RF-MEMS Switched Varactor for High Power Applications[END_REF]. This geometry has been especially studied to consider electrostatic pressures generated by high RF power signal. The actuation part of cantilevers and their contact area are well separated and have been designed to be almost fully independent. That's why, the architecture of the beam consists in two different areas: a large actuation part and a smaller contact tip as shown on Fig. 1. Hence, when a RF signal is applied across the component, due to its electric field, a static attractive electrostatic pressure is induced on the movable beam. In fact, this pressure mainly depends both on the RF signal magnitude, on the localized area geometry of the beam where this field is concentrated (the beam tip in our case) and also on the distance separating this area from the RF line where the signal propagates.

To recover up state position, when the actuation command is interrupted, the beam must have a restoring force greater than the electrostatic RF power attraction force. Designing a restricted beam tip area allows to reduce the effect of electrostatic pressure induced by the RF signal which is really critical in down state position when the tip is put into contact with RF line. Indeed, contrary to the up state position, the induced RF electric field is much higher in down state since the air gap is almost reduced to zero, only a thin dielectric layer isolates the beam tip from the RF line. That is why hot switching operation is the operating mode where power failure first appears. Thanks to this specific tip geometry, the power handling of the component could be highly improved up to several Watts on 50 Ohms impedance.

The other particularity of the proposed beam geometry is the difference of gaps between the actuation area and the contact tip. Hence in the beam down state position, a higher air gap on the actuation area allows improving the contact force on the tip and avoiding in the same way some collapse phenomenon between the beam and its pull-down electrode after the tip has made contact on the RF line. Thus, the component reliability could be grandly improved since dielectric charging troubles generally observed in contact mode electrostatic actuators is strongly reduced.

III. TEST CELLS

Several switched capacitors geometries have been designed and analytical calculus have predicted that most of them should be able to handle RF signal power of 4 Watts at least. To evaluate the real power handling capabilities under hot switching conditions, fabricated devices have been tested for different level of RF power monitoring their minimum DC release voltages before recovering the cantilever up state. An example of cell which have been tested is shown Fig. 2 and Fig. 3 presents there is no significant voltage change up to 18 Watts under 50 Ohms revealing no failure phenomenon occurs at least until this RF power level. To our knowledge, there is no MEMS based phase shifter in literature able to operate under several Watts in hot switching conditions. The presented switched capacitor concept could be implemented in a DMTL (Distributed MEMS Transmission Line) to reach such power handling. To design such phase shifter a high impedance microwave transmission line is generally used periodically loaded with tens or more of varactors, the principle of a DMTL phase shifter topology is shown Fig. 4 and a fabricated testing cell is presented Fig. 5. The phase shift is achieved by turning on several of the switched capacitors distributed along the line allowing a decrease of the line impedance and so meaningful change on the RF signal velocity. Actually, the available phase shift increases with the number of cantilevers placed along the line. Thus several switched capacitors can be actuated in the same time as a single element (a bit) in order to reach well defined phase shift value: 0°, 90°, 180°, 270° as example on a 2 bit phase shifter. Compared to analog phase shifters, this approach guarantees a very low sensitivity to phase noise and better performance reproducibility.

IV. CONCLUSION

For instance, several elementary phase shifting cells (1-bit topology) have been fabricated. Each cell is based on a coplanar waveguide loaded with 2 cantilevers symmetrically placed at each side. The RF performances and the power handling of these structures are currently under test. The last results will be presented during the conference.

Fig. 1 .

 1 Fig. 1. Principle of proposed cantilever based switched varactor.

Fig. 2 .Fig. 3 .

 23 Fig. 2. Testing cell.

Fig. 4 .

 4 Fig. 4. Principle of a DMTL phase shifter topology.

Fig. 5 .

 5 Fig. 5. Picture of a fabricated phase shifting cell.


H Yang

T.-H Tsai

C.-C Hu

Portable Valve-less Peristaltic Micropump Design and Fabrication

This paper is to describe a design and fabrication method for a valve-less peristaltic micro-pump. The valve-less peristaltic micro-pump with three membrane chambers in a serial is actuated by three piezoelectric (PZT) actuators. With the fluidic flow design, liquid in the flow channel is pumped to a constant flow speed ranged from 0.4 to 0.48 mm/s. In term of the maximum flow rate of the micro-pump is about 365μl/min, when the applied voltage is 24V and frequency 50 Hz. Photolithography process was used to fabricate the micro-pump mold. PDMS molding and PDMS bonding method were used to fabricate the micro-channel and actuator chambers. A portable drive controller was designed to control three PZT actuators in a proper sequence to drive the chamber membrane. Then, all parts were integrated into the portable valve-less peristaltic micro-pump system.

I.

INTRODUCTION

Recent progress in MEMS technology provides micro-fluid system manufacture and application pluralism further. Because the micro-fluid systems have the advantages of tiny size and easy to carry, also have high accuracy and short response time. They have quite great values, no matter in such fields like semiconductor, electronic, machinery, chemical analysis or biomedicine and laboratory chip development. Among the micro-fluid control system components, they include micro-channel, micro-valve, micro-pump, micro-sensor and micro-actuator. Micro-pump is the key component in the micro-fluid control system development, since the micro-fluid controls system requires a power to transport fluids via micro-pump. The micro-pump has to control the flow rate accurately and efficiently.

Micro-pump systems usually include an actuating chamber and a valve. There is always a reciprocating cycle vibration in the actuating chamber and the flow channel, it is due to the volume variation of the chamber via membrane moving. A pressure drop is naturally formed when fluids flowing. In accordance with its different actuating model, it can be divided into piezoelectric, electrostatic, thermomechanical, electromagnetic and shape memory alloy types. In addition, the valve can be divided into check-valve and valve-less types. The main function of the valve is to control the flow in a unique direction. The valve of the check-valve type acted as a blocking slice using a cantilever beam in the port. When applied a micro-actuator, it resulted in the difference between the internal and external pressure inside the micro-pump. The valve blocking slice in the channel turns on or off, making the fluid to flow in one way without a reversing flow. Since the valve is operated frequently, it is easy to cause the valve material fatigue or unable to return its original state which affect micro-pump efficiency and life-span. The valve-less micro-pump is comparatively simple and stable.

Koch et al. in 1998 issued the piezoelectric membrane micropump with size of 8mm× 4mm×70μm which deposits PZT material on silicon wafer as actuator [START_REF] Koch | A novel micromachined pump based on thick film piezoelectric actuation[END_REF]. After silicon process of the heat treatment, polarize and electrode, then PZT actuate membrane can be finished, and other structure including actuating chamber, check-valve of cantilever type and entrances-exits are also made with silicon process technology. Micropump have a maximum flow rate 120μl/ min at inputting voltage 600 Vpp with frequency 200 Hz. The advantages of this micropump is lower costs for mass production, the shortcoming is processing complicatedly for PZT membrane. Zengerle et al in 1995 issued the electrostatic type micropump that membrane and operating electrode are separated by insulating barrier [START_REF] Zengerle | A bidirectional silicon micropump[END_REF]. As both ends have a different electric charge which resulted in membrane out of shape with electrostatic, the flow rate is 300μl/ min with maximum moving frequency 1KHz, but it is up to 150~200 operating voltage. Jang in 2000 issued the electromagnetic type micropump, its operating principle is utilizing magnetic and electric field that direction is vertical each other to produce Lorentz force to drive fluids [START_REF] Jang | Theoretical and experimental study of MHD micropump[END_REF]. Its fluid should have conductivity characteristic that can produce Lorentz force and move under the magnetic and electric field reciprocating, it is a kind of non-mechanical micropump. This micropump physical dimension is 40mm×1mm×200μm, the diameter of the entrance and exit of the pipe is 2 mm, when inputting electric current is 1.8 mA, the flow rate is 63μl/ min. Its characteristic is under the interactive influence of the magnetic and electric field for the working fluid to produce Lorentz force and move, so it does not need actuating device on the structural design and does not need to consider the question of the valve either. Its advantage is suitable for high conductivity working fluid, two-way flow, structure simple and process easy; but the shortcoming is that it need plus a regular permanent magnetic field. Stemme issued bimetal micropump in 2000, this micropump membrane is consisted mainly by silicon and aluminum [START_REF] Stemme | Valveless Diffuser/Nozzle Based FluidPump[END_REF]. As the polycrystalline silicon set up an electric circuit and heated, the membrane is crooked and out of shape because the degree of expansion is different when silicon and aluminum are heated. The advantage is its operating voltage low, the shortcoming is the consuming energy and frequency low just only 0.5Hz. 1997, Benard et al., [START_REF] Benard | A Titanium-Nickel Shape Memory Alloy Actuated Micropump[END_REF] issue memory alloy type micropump, this micropump utilized memory alloy that have different shape characteristics to actuate membrane under different temperature, so it has the function of reciprocating motion micropump under being proper to control, its advantage is difficult to be tired and operate the voltage low for the material, and the shortcoming is operating frequency low and process complicated.

Piezoelectric type micropump which drived by the membrane is developed firstly by Spencer [START_REF] Spencer | An Electronically Controlled Piezoelectric Insulin Pump and Valves[END_REF] in 1978, micropump combines the piezoelectric actuating slice above the membrane that enable piezoelectric material to drive the membrane to produce high frequency vibration, but after reaching certain vibration frequency, the valve is too late to shut down and cause the phenomenon of back flowing, the main reason is that membrane replying speed can not catch up with its vibration frequency, so the flow rate of micropump is decreasing obviously. Later in 1990, Spencer [START_REF] Spencer | Piezoelectric Micropump with Three Valves Working Peristaltically[END_REF] issue Peristaltic micropump, the major principle is utilizing reciprocating motion of membrane to change the volume of chamber, hence there is a pressure drop inside and outside the chamber that promote fluids to move, its characteristic is three phase place to control the motion of membrane to actuate the glass membrane, resulted in transport function of fluid, frequency at 15 Hz can reach flow rate of 100μl/min, the working figure of micro pump. 1990, Pol [START_REF] Van De Pol | Thermopneumatic Micropump Based on Micro-engineering Techniques[END_REF] issue pneumatic type micropump, one of membrane's end connected to gas pump and squeezed into high-pressure air to make membrane deformation, its advantage is simple process and can verify the experiment fast, but need extra pressure sources is its shortcoming. 2000, Unger [START_REF] Unger | Monolithic Microfabricated Valves and Pumps by Multilayer Soft Lithography[END_REF] choose elasticity PDMS as materials, cooperating with outer way of air pressure to drive membrane to be out of shape, it is mainly two molds of different channel shape after PDMS molding and counterpoint joint, then the atmospheric pressure was pushed into the above channel and then pinches the below fluid of channel to reach the function of pushing the fluid, its channel width is 40μm, height is 10μm, it can be look as the function of valve as one group atmospheric pressure channel, if many group atmospheric pressure channel pinches in a proper order, it can be acted as a Peristaltic micropump, but because atmospheric pressure channel with microchannel overlap area is small, so its maximum flow rate is up to 0.14μl/min only. 1995, Olsson [START_REF] Olsson | A valve-less planar fluid pump with two pump chambers[END_REF] et al., issue diffusing type micropump, its operating principle is to utilize geometry design of the micro-structure, using the difference of come in and go out flow that resulted in diffusion effect to actuate microfluid, its advantage is design and process simple, it is apt to combine it on the chip, and only utilize the geometry form of the micro-structure to cause the function of the valve, so it also call valve-less micropump, but its shortcoming causes the non-directional flow in a chamber of microfluid due to diffusion effect, it is apt to cause the effect of mixing.

The vale-less micro-pump is originated from a diffusing channel concept. It utilizes the micro-structure geometric design by using the difference between inlet and outlet flow that resulted in diffusion effect to actuate the micro-fluid. Its advantages include simple geometry, no active component, providing a constant flow rate. It is apt to combine all components on a chip, and only utilizes the geometrical micro-structures to cause the valve function. That is so call valve-less micro-pump. This paper presents the design and manufacture of the portable valve-less peristaltic micro-pump system, utilizing photolithography technology to produce structural pattern in photoresist and using PDMS material as the molding material. The structure after molding can form flowing channel and actuator chamber, then making a portable circuit controller to control three piezoelectric slices on actuate chamber membranes. By controlling three PZT actuators in a proper sequence, the membrane deformation was actuated directly and generated the wriggling motion in the channel. Due to the valve-less design and the wriggle flow, it is a portable valve-less peristaltic micro-pump system. This paper utilizes the ANSYS to imitate the flow behavior in the micro-channel and results in an optimum solution for the micro-pump. Valve-less micropump uses geometry structure of diffuser/nozzle to replace cantilever beam structure of valve to limit the direction of the fluid flow; diffuser/nozzle utilize its difference of conical mouth degree and pressure drop of chamber causing by actuating membrane that make working fluid in diffuser/nozzle have the difference of flow rate and reach the transportation function of the fluid.

II. THEORY

It is mainly the pressure recovery coefficient Cp to decide the design of diffuser/nozzle, the relationship figure is the size of diffuser/nozzle and pressure recovery coefficient Cp, the higher is pressure recovery coefficient Cp value, the well design for expand and contract mouth, the lose coefficient K is correlated with pressure recovery coefficient Cp, the larger the pressure recovery coefficient, the smaller loss coefficient K, such as Eq.(1) shown

Cp A A 1 K 2 out in - ⎟ ⎟ ⎠ ⎞ ⎜ ⎜ ⎝ ⎛ - = (1)
A in : Cross-section area of entrance A out : Cross-section area of exit In order to design better expansion and contraction mouth, it can be got by the relation between K and Cp, it is good design that opening angle 2θ is small, in 5°~20° lose coefficient is

+ ρ = + ρ (3)
The structure of diffuser mouth d, it is the increasing cross section that can reduce the velocity of fluid in order to increase the pressure; The structure of nozzle n, it is the cross section decreasing gradually that can increase the velocity of fluid in order to reduce the pressure. As the fluid flows through the structure of the diffuser and nozzle, the pressures can be expressed respectively as Eqs. ( 4) and ( 5): When the pressure loses coefficient is larger, then the flow rate of volume is relatively smaller; On the contrary, when the pressure loses coefficient is smaller, then the flow rate of volume is relatively larger, hence by good design of the channel length、scale and conical degree to reach the idea flow rate.

2 V P 2 d d d ρ ξ = Δ (4) 2 V P 2 n n n ρ ξ = Δ (5)

III. EXPERIMENTS

The purpose of this research is to design and fabricate simple structure, easy process, low cost and efficient well for portable valve-less peristaltic micropump system, the major experiment course can be divided into four parts. First is mainly focus on the design of micropump working structure, this research selects valve-less type micropump structure. Second is utilizing ANSYS analyses software to simulate the diffuser structure and nozzle profile to analysis the influence of various angles and lengths of the diffuser and nozzle profiles to get the most stable and efficiency of the physical dimension. Third is to fabricate a microchannel and an actuator chamber. Photolithography process in MEMS technology was used to fabricate main structure and combine PDMS technology to duplicate the microchannel and actuator chamber. Forth is to fabricate the portable control circuit board that can actuate piezoelectric slice and drive the chamber membrane for fluid flowing.

The actuator chamber and channel size of micropump are obtained by according to piezoelectric slice dimension and ANSYS analysis; The structure is shown as Fig. 1, which design three groups of actuator chambers, the size is 10 mm in diameter, another is about diffuser and nozzle geometry size that its key size can be divided several parts. First is entrance width W1, it is in size 300μm, second is diffuser and nozzle mouth angle 2θ that angle is 10 degrees, third is diffuser and nozzle length L that size is 3.4 mm. Analysis of the diffuser is mainly to analyses the influence of the fluid velocity efficiency that the angle causes. The angle 2θ is to be better for smaller, when opening angle 5~20 degrees have minimum of lose coefficient, so selecting simulation angle definitely as 5, 10, 15, 20 degrees of angles, the width W1 of entrance size is 300μm and the length L of the diffuser is 3.4mm, the simulating measurement of the diffuser for different angles is showing as Table 1. (1) 2θ =5°

(2) 2θ=10° 2 is the result of imitation for the velocity of flow of the diffuser, from the result of imitating, the width W1 of the entrance of the diffuser is 300μm, length L is 3.4mm, and angle 2θ is 10 degrees that its efficiency is the most stable as shown in Fig. 3. It is the relationship of the opening angle and loss rate of the flow speed; it did not discuss the back flow phenomenon of the diffuser, but the result in simulation is in accord with the theory that is designed, in L/W =11.33 and 2θ=10 degrees, it is one transient stall, its pressure loss is minimum, it is the best performance. This research utilizes photolithography technology to fabricate a resist structure which include microchannel and actuator chamber and combine PDMS molding technology to duplicate out the microchannel, actuator chamber and membrane as illustrate in Fig. 1. It is a micropump sketching figure; With simple and convenient joint technology, the PDMS molding was bonded on the glass and then use piezoelectric slice to actuate the chamber membrane which utilize vibrating principles of upper and lower piezoelectric slice to drive the membrane out of shape, hence driving the fluids of the chamber to move. This experimental design for driving the chamber is 10mm, using designed actuating circuit to control piezoelectric slice that drive the chamber membrane to reach the function of wriggling. The process of this experiment can be divided into two major steps, first part is utilizing photolithography technology to fabricate a resist structure. The second part is molding, using molding principle to fabricate main structure of micropump by using elastic high polymer material PDMS.

This research uses the JSR-THB negative resist to fabricate JSR mold on the silicon wafer, it is fabricated by the multi-layer coating technology for high aspect ratio microchannel protruding structure. As Fig. 4(a) showing, it is JSR mold finished product figure, it is quite intact that can be observed from actuate chamber and microchannel structure, they did not appear fracture and stripped situation caused by stress concentration. In the course of processing, in order to remove the surplus of the resist solvents, it utilizes baking hard for a long time at low temperature, for avoiding PDMS colloidal to response when PDMS molding appeared unsolidification and stuck on mold; As Fig. 4(b) showing, the chamber and microchannel of PDMS structure, it can be found out the structure to be quite intact. This resist structure is measured with the 3D confocal microscopy, it utilizes this microscope to take the 3D surface outline of the structure and measures the thickness of the structure. Fig. 5 is the 3D surface outline picture of microchannel, it can find out that microchannel surface is very smooth by using photography process. Fig. 6 is the thickness measurement of the channel, the thickness of channel is nearly 350μm. During the past studying process, other microchannel structures were made. Their thickness ranges were in 346~403μm, it may be an error caused by experimental equipment or the used of negative resist material.

As Fig. 7 showing, this experiment is according to the design step of assembling, it takes with actuator controller, piezoelectric slice and micropump structure to assemble and combine to a portable valve-less peristaltic micropump system, because it did not consider the fluid chamber room when design the micropump originally, so transport fluid in testing, prior to entrance dripping 15ml red dyestuffs to aqueous solution to observe whether the red dyestuffs aqueous solution can flow smoothly in microchannel flow as shown in Fig. 8 (a) (b). The result of the test is that the red dyestuffs aqueous solution can be output smoothly and have not overflowing situation, it is indicated that the microchannel joint strength intensity is enough and have no stopped up phenomenon using PDMS to be joint; Because it did not consider the fluid chamber room, the test of flow rate is measured by sending the volume of the fluid with different frequency, then divided by time, this part is the research that structure need to improve again; As Fig. 9 showing, the flow rate have examined at direct current voltage 24volt for every frequency, but at frequency 50Hz under in a proper order control, the flow rate of volume is about 365μl/min for maximum, according with originally expectation that the flow rate of volume can reach the demand above 200μl/min. IV CONCLUSION This research is successful using the multi-layer coating method for ultra-thick negative JSR-THB resist to fabricate actuator chamber and microchannel JSR mold of high aspect ratio structure. The characteristics of this research is the micropump structural design and multi-layer coating processing that can reduce process time and cost, and with the portable driving circuit for piezoelectric slices to cooperate with the operational frequency to make three groups of chamber membrane producing wriggles in a progressive way to achieve the goal of promoting fluids. It is obtained the following conclusions from the above experiments.

9- 11

 11 April 2008 ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5

9- 11

 11 April 2008 ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5 minimum, 20°~40° lose coefficient is maximum. According to Bernoulli equation, for any two cross-sectional in pipe for incompressible flow is shown as Eq. (2pressure, V: Fluid Velocity, Z : Height of streamline, ρ : Fluid density, g: Gravity constant.If the fluid streamline is highly pretty much the same, it can neglect the gravity iterm, then the Bernoulli equation can be written as Eq. (3

ξ

  d : Pressure loss coefficient of diffuser ξ n : Pressure loss coefficient of nozzle ξ d 、ξ n will change with different channel length、scale and conical degree. By the continuity equation, the flow rate of fluid flowing through the diffuser and nozzle structure can be written as Eqs. (6Narrow Cross-section area of diffuser A n : Narrow Cross-section area of nozzle If the designed structure of diffuser and nozzle is same in size, namely the cross section A d = A n = A, the volume flow rate can be shown as Eqs. (8) (9): the narrow mouth velocity of diffuser and nozzle structure can be shown as respectively Eq.(10)(11): 10)(11) into Eq. (8)(9), the volume flow rate of the diffuser and nozzle structure is shown below:

Fig. 1

 1 Fig. 1 Schematic diagram of the portable valve-less peristaltic micropump.

Fig. 2

 2 Fig. 2 Simulation results of the fluid velocity using different angles for the diffuser.

Fig. 2

 2 Fig. 2 is the fluid velocity simulation result from ANSYS for different angle diffuser. Table2is the result of imitation for the velocity of flow of the diffuser, from the result of imitating, the width W1 of the entrance of the diffuser is 300μm, length L is 3.4mm, and angle 2θ is 10 degrees that its efficiency is the most stable as shown in Fig.3. It is the relationship of the opening angle and loss rate of the flow speed; it did not discuss the back flow phenomenon of the diffuser, but the result in simulation is in accord with the theory that is designed, in L/W =11.33 and 2θ=10 degrees, it is one transient stall, its pressure loss is minimum, it is the best performance.

  Fig. 4 Fabrication results of the nozzle and chamber; (a) resist mold in JSR and (b) replicated structure in PDMS.

Fig. 5 Fig. 6

 56 Fig. 5 Three dimensional profile measurement of the flow channel.

Fig. 7 Fig. 8 Fig. 9

 789 Fig. 7 System assembly of the valve-less peristaltic micropump.

( 1 )

 1 PDMS molding and PDMS bonding method are used to reduce the process time of photography process. It can reduce cost in mass production. (2) The processing technique is skillful to control the dimension of the microchannel steadily. (3) Micropump structure is PDMS that have advantages of elastic and bearing falling etc. (4) The wriggling transportation of the fluid is a very smooth flow rate and provides a high accurate flow rate.

Table 1

 1 Simulation of different angles for the diffuser.

	No W1(mm) L(mm)	2θ	W2(mm)	L/W1
	1	0.3	3.4	5°	0.597	11.33
	2	0.3	3.4	10°	0.895	11.33
	3	0.3	3.4	15°	1.195	11.33
	4	0.3	3.4	20°	1.5	11.33

Table 2

 2 Result of velocity simulation for diffusers.

	No Inlet(mm/s)	Vmax(mm/s)	Vout(mm/s)
	1	0.35		0.48	0.261
	2	0.35		0.45	0.192
	3	0.35		0.43	0.153
	4	0.35		0.41	0.120
			Vmax	Loss rate
		0.5			0.8
		0.48			0.7
		0.46			0.6
	V, mm/s	0.44			0.5 0.4	Loss rate, %
		0.42			0.3
		0.4			0.2
		0.38			0.1
		0.36			0
		5°10 °15 °	20 °
			2 θ
	Fig. 3 Relationships of opening angles related to velocity and
			loss rate.

ACKNOWLEDGMENT

This research is supported by the National Science Council of Taiwan under the grand number NSC95-2212-E-005-091.


Ciprian Iliescu

Elena Barbarini

Marioara Avram

Guolin Xu

Andrei Avram

Microfluidic Device for Continuous Magnetophoretic Separation of Red Blood Cells

This paper presents a microfluidic device for magnetophoretic separation red blood cells from blood under continuous flow. The separation method consist of continuous flow of a blood sample (diluted in PBS) through a microfluidic channel which presents on the bottom "dots" of ferromagnetic layer. By applying a magnetic field perpendicular on the flowing direction, the ferromagnetic "dots" generate a gradient of magnetic field which amplifies the magnetic force. As a result, the red blood cells are captured on the bottom of the microfluidic channel while the rest of the blood is collected at the outlet. Experimental results show that an average of 95 % of red blood cells is trapped in the device.

I.

INTRODUCTION

Separation of red blood cells (RBCs) from whole blood is often an essential step before the application of many clinical and molecular diagnostic tests on blood samples. The procedure is also useful in blood transfusion applications as it allows the selective transfusion of particular cell types. Separation can be possible due to hemoglobin. Hemoglobin is a conjugated metal-protein comprising of four polypeptide globins chains containing a ring structure covalently bonded with a central ferrous iron atom (Fe 2+ ), which binds reversibly with oxygen. If is deoxygenated, each of the four iron atoms contains four unpaired electrons, giving the protein and the cell a substantial paramagnetic moment. White blood cells (WBCs) do not contain hemoglobin and are diamagnetic particles. These properties are the starting point for the trapping of RBCs particles using high gradient magnetic separation.

Here, we report a simple microfluidic device for extraction of RBCs from blood under continuous flow. The device consists of a microfluidic channel with ferromagnetic "dots" deposited on its bottom. The geometry of the ferromagnetic layer and the application of an external magnetic field perpendicular to the flow direction generate a magnetic force on the RBCs present in the blood. The device offers other advantages, such as the collection of the RBCs inside the channel. The glass structure allows the complete visibility and analysis of the sample.

II.

DEVICE DESIGN The structure of the device is presented in Fig. 1. A glass die with inlet / outlet holes and a 60µm-depth microfluidic channel is bonded to another glass die, on which a ferromagnetic structure -square "dots" (2x2x2µm) of Niwas patterned. A permanent magnet generates an external magnetic flux of 0.2 T perpendicular on the flow direction. The blood is diluted with PBS and is flown through the microfluidic channel. The ferromagnetic "dots" generate a gradient of magnetic field which amplifies the magnetic force that acts on RBCs. As a result the RBCs are trapped by the ferromagnetic layer while the WBCs are flushed out with the plasma and the other blood components.

ANALYTICAL CONSIDERATIONS

If a uniform external magnetic field is applied normal to the axis of a ferromagnetic wire, it is deformed near the ferromagnetic wire, and generates a high gradient magnetic field, experienced by the magnetic particles moving around the wire. Therefore, blood cells flowing near the ferromagnetic wire experience a magnetic force created by the high gradient magnetic field near the ferromagnetic wire.

The magnetic force of a rectangular wire on blood cells, located around the x-axis in Figure, can be calculated as

a r a H h w r a BC V k r a H h w r a h w k r a BC V k BC F > ∆ - ⋅ ⋅ + ∆ - =                           , 2 0 2 sin 3 2 0 2 2 0 2 cos 2 2 3 2 0 2 φ ϕ χ µ ϕ χ µ with B w B w k µ µ µ µ + - =
where ∆χ is the relative magnetic susceptibility of a blood cell relative to the buffer solution; µw and µB are the permeabilities of the ferromagnetic layer and the buffer ©EDA Publishing/DTIP 2008 ISBN: 978-2-35500-006-5 9-11 April 2008 solution, respectively; VBC is the volume of the blood cell; a is the lateral dimension of the ferromagnetic structure; r and φ are the cylindrical coordinates of the distance and angle; H 0 is the external magnetic field; and a r and a φ are unit vectors for the distance and angle in the cylindrical coordinates.

If the ferromagnetic structure is magnetically saturated, the first term of the magnetic force in the equation is independent of the external magnetic field, H 0 , and proportional to the square of the saturation magnetization field. The second and third terms of the magnetic force are linearly proportional to the saturation magnetization field and the external magnetic field.

For magnetic particles placed on the x-axis (φ ≈ 0), sin2φ ≈ 0, cos2φ ≈ 1, the ferromagnetic material attracts paramagnetic particles (RBCs) and repels diamagnetic particles (WBCs). For magnetic particles placed on the y-axis (φ ≈ 90), sin2φ ≈ 0, cos2 φ ≈ -1, the material attracts diamagnetic particles (WBCs) and repels paramagnetic particles (RBCs). The first geometric configuration has been called the paramagnetic capture (PMC) mode; the latter has been called the diamagnetic capture (DMC) mode. To summarize, the main parameters that influence the magnetic force applied on the particles affected by and external magnetic field are: the intensity and orientation of the magnetic field H 0 ; the permeability of the buffer solution and the ferromagnetic material; the susceptibilities of the cells and the buffer solution and the dimension of the device.

IV.

FABRICATION PROCESS The fabrication process of the device consists of three important steps: fabrication of the top wafer with the inlet/outlet holes and the microfluidic channel, fabrication of the bottom wafer with ferromagnetic concentrators and assembling of the wafers.

The fabrication of the top wafer is presented in Figure 2. A 4" glass wafer Corning 7740 was used. A Cr/Au/ photoresist masking layer (Figure 2a) was used for fabrication of the microfluidic channel. A wet etching process in HF/HCl was used for 60µm deep-etching in glass (Figure 2b). During the wet etching process the backside of the wafer was protected by a dummy silicon wafer bonded with wax on the glass wafer. The debonding of the wafer was performed on a hot plate. After removing of the Cr/Au/photoresist mask (Figure 2c) using classical resist stripper (NMP) -for photoresist and residual wax and Au and Cr etchants, a second Cr/Au/photoresist mask was applied on the other surface of the glass wafer (Figure 2d). Similar, a second wax bonding process will assure the protection of the microfluidic channel, while the wax served also as etch-stop layer for the deep wet etching process in HF/HCl (Figure 2e). Finally, the glass wafer was debonded from the dummy silicon wafer and the masking layer was removed with a similar procedure as was described before.

The fabrication of the bottom wafer consists of a simple lift off process presented in Figure 3. On a 4" glass wafer (Corning 7740) a 5µm-thick photoresist mask (AZ4620 from Clariant) was deposited (Figure 3a). The metal layer Ti/Ni (50nm/2µm) was deposited in a CHA e-beam evaporator -(Figure 3b). The photoresist masking layer was finally removed in an ultrasonic bath using acetone. For the assembling of the glass wafers a simple adhesive bonding, using SU8-5, photoresist process was used. The technique consists of imprinting of an SU8-5 layer initially spun on a dummy silicon wafer (Figure 4a) on a Teflon cylinder (Figure 4b) and transferring further the adhesive from the cylinder to the top glass wafer (Figure 4c). This contact imprinting technique allowed deposition of a thin adhesive layer only on the bonding regions. In the last step, the wafers are manually align and bonded at 150OC for 30 minutes with an applied pressure of 1000N (Figure 4d). Finally, the wafer is diced. An image of the fabricated device is presented in Figure 5. The dimensions of the chip are 32mm x 9mm. V.

CONCLUSIONS The paper presents a microfluidic device for magnetophoretic trapping of red blood cells under continuous flow, using high gradient of magnetic field. Experimental results show a good trapping efficiency that can be further improved.

Figure 1 :

 1 Figure 1: Schematic view of the microfluidic device for RBCs trapping

Figure 2 :

 2 Figure 2: Main steps of the fabrication process of the top wafer: a) processing of the Cr/Au/ photoresist masking layer, b) wet etching of the microfluidic channel, c) removing of the masking layer, d) processing of the Cr/Au/photoresist masking layer for inlet/outlet holes, e) wet etching of inlet/outlet holes, f) removing of the second masking layer.

Figure 3 :

 3 Figure 3: Main steps of the bottom wafer fabrication process: a) deposition of the photoresist mask, b) deposition of Ti/Ni layer, c) removing of the photoresist mask.

Figure 4 :

 4 Figure 4: Assembling of the wafers: a) deposition of SU8 on a dummy wafer, b) imprinting of the SU8 on a Teflon cylinder, c) imprinting SU8 on the bonding area, d) alignment and bonding.

Figure 5 .

 5 Figure 5. Photo of the microfluidic chip for RBCs trapping V.TESTING Diluted blood (1:20 in PBS) was used for testing purpose. The permanent magnet creates an external magnetic flux of 0.2 T and the flow rate was set in the range between 0.5 and 0.7 ml/h (using a syringe pump). Two connectors fabricated by polymer printing secured the inlet and outlet connections. The results are based on the analysis of the quantity of red cells collected at the output of the device. Experimental results show an average of 5% of red blood cells collected at

Figure 6 .

 6 Figure 6. Image of the field density of the cells on Neubauer hemocytometer before and after flowing the blood sample through the microfluidic device


Chun-Ping Jen

Yao-Hung Huang

Teng-Wen Chen

Cell Trapping Utilizing Insulator-based Dielectrophoresis in The Open-Top Microchannels

published or not. The documents may come

INTRODUCTION

The ability to manipulate particles, especially living cells, is essential to many biological and medical applications, including isolation and detection of rare cancer cells, concentration of cells from dilute suspensions, separation of cells according to specific properties, and trapping and positioning of individual cells for characterization. Several kinds of physical forces can be employed for particle manipulation. Those are optical [START_REF] Chiou | Massively parallel manipulation of single cells and microparticles using optical images[END_REF], mechanical [START_REF] Huang | Continuous Particle Separation Through Deterministic Lateral Displacement[END_REF], magnetic [START_REF] Berger | Design of a microfabricated magnetic cell separator[END_REF][START_REF] Lee | Integrated Cell Manipulation Systems[END_REF], electrical [START_REF] Huang | Dielectrophoretic cell separation and gene expression profiling on microelectronic chip arrays[END_REF][START_REF] Dürr | Microdevices for manipulation and accumulation of micro-and nanoparticles by dielectrophoresis[END_REF], and other manipulations [START_REF] Yoshida | Hydrodynamic micromanipulation of individual cells onto patterned attachment sites on biomicroelectromechanical system chips[END_REF][START_REF] Hawkes | Continuous cell washing and mixing driven by an ultrasound standing wave within a microfluidic channel[END_REF].

Dielectrophoresis (DEP) has been widely used for the manipulation, separation and characterization of cells, DNA, virus, and colloid particles (in both micro-and nano-scales) in various microfluidic platforms. DEP force is generated by electrode structures of appropriately designed such as the pin-plate [START_REF] Huang | Electrode design for negative dielectrophoresis[END_REF] isomotive [START_REF] Marszalek | Experimental verification of a theoretical treatment of the mechanism of dielectrophoresis[END_REF], polynomial [START_REF] Pohl | Electrode geometries for various dielectrophoretic force laws[END_REF], or castellated interdigitated forms [START_REF] Pethig | Positive and negative dielectrophoretic collection of colloidal particles using interdigitated castellated microelectrodes[END_REF]. Early studies of DEP used electrode structures made from thin metal wires, needles, or plates, while modern DEP employs microfabrication technology to produce microelectrode arrays capable of producing sufficiently large DEP forces to induce particle motion with small applied voltages. DEP on micro-fabricated electrodes has been proved especially suitable for its relative ease of micro-scale generation and structuring of an electric field on microchips. Moreover, integrated DEP biochips provide the advantages of speed, flexibility, controllability and ease of application to automation because DEP traps consist of scalable electrode arrays be designed to pattern thousands of cells on a single glass slide. DEP have been using as cell trapping [START_REF] Washizu | Electrostatic manipulation of DNA in microfabricated structures[END_REF], levitation [START_REF] Arnold | Positioning and levitation media for the separation of biologicalcells[END_REF], separation [START_REF] Markx | Dielectrophoretic characterization and separation of micro-organisms[END_REF], and sorting [START_REF] Washizu | Handling biological cells using a fluid integrated circuit[END_REF] on an electrode array by varying electrode shape and arrangement. DEP on micro-fabricated electrodes has been proved especially suitable for its relative ease of micro-scale generation and structuring of an electric field on microchips.

DEP force field could also be generated by using dielectric constrictions or insulating obstacles, therefore, these methods further extend the scope of DEP applications. These approaches have been termed electrodeless dielectrophoresis (EDEP) [START_REF] Chou | Electrodeless dielectrophoresis of single-and double-stranded DNA[END_REF] or insulator-based dielectrophoresis (iDEP) [START_REF] Lapizco-Encinas | Insulator-based dielectrophoresis for the selective concentration and separation of live bacteria in water[END_REF], respectively, in distinguishing from the method conventionally employed by metal electrodes-based DEP for force generation.

In this study, a lower conductive material of photoresist (JSR, THB 151N) was adopted as a structure in open-top microchannel instead of a metallic wire to squeeze the electric field in a conducting solution, therefore, creating a high field gradient with a local maximum. The microchip with the open-top microchannels was designed and fabricated herein. The schematic illustration of the trapping device is shown in Fig 1. This trapping microchip is suitable for trapping cells or biological samples and easily proceeding further treatment for cells, such as culturing or contact detection. Biological sample and buffer loading before analysis and cleaning after analysis in the proposed microchip with the open-top microchannels could become easier and faster, hence, it is beneficial in high-speed sequential multiple measurements. Furthermore, the Joule heat generated by applying high voltage in the open-top microchannel could dissipate more effectively than that in the enclosed microchannel, because that it is open to the air. When the fluorescent detection is used, the intensity of the emitted light will not be interfered by the cover, which is absent in the microchip with open-top microchannels.

II. DIELECTROPHORETIC(DEP) FORCE THEORY

The phenomenon of DEP is first defined by [START_REF] Pohl | Electrode geometries for various dielectrophoretic force laws[END_REF] as the motion of neutral particles caused by dielectric polarization effects in non-uniform electric fields that alternating fields of a wide range of frequencies were used. The DEP force F DEP acting on a spherical particle of radius r suspended in a fluid of permittivity m ε is given: where and are complex permittivity of the dielectric particle and the medium, and the magnitude of the electric field, E, may be replaced by E rms , which is root-mean-square of the external field, in an alternating field. , the real (in-phase) part of the Clausius Mossotti factor, , is a parameter of the effective polarizability of the particle which varies as a function of the frequency of the applied field (ω), and the dielectric properties of the particle and the surrounding medium.

) ( )] ( Re[ 2 2 0 3 E K r F m DEP ∇ = * ω ε ε π (1)
* p ε * m ε )] ( Re[ ω * K ) (ω * K III. FABRICATION OF THE MICROCHIP
The microchip of with the open-top microchannels was fabricated by microfabrication techniques, which were depicted in Fig. 2. The electrodes were made by depositing chrome and aurum sequentially on the slides by an electron beam evaporator. Chrome was the first deposited for the purpose of enhancement. The finally deposition thickness of Chrome and Aurum are 50 nm and 100 nm. The JSR photoresist (60 μm in thickness) was spun on the glass slide to pattern micro-channels of constriction for insulator-based DEP. The PDMS prepolymer mixture was cured and cut to form a well of 3 cm by 1.5 cm. The glass slide and the PDMS well were bonded after the treatment of oxygen plasma to avoid the leakage of the sample.

IV RESULTS AND DISCUSSIONS

The contours of the square of electric field at different heights in the open-top microchip were simulated and depicted in Fig. 3. The applied electric field is 3×10 4 V/m; the space and the angle of the tip are 60 μm and 30 degree, respectively. The insulator (JSR photoresist) is 60 μm in height. On the surface and 30 μm above of the glass slide (where is the middle plane of the insulator), the local maximum of the electric fields occurred at the tips which form the constriction of the electric field. When the height increases, the electric field became uniform due to the absence of the insulator-based constriction. The distributions of the gradient of E 2 along the centerline between the tips at different heights (for the gap between tips 60 μm and the angle of the tip 30 degree) were plotted in Fig. 4a. The magnitude of gradient of E 2 is proportional to the DEP force. Comparing to the DEP force at the surface of the glass slid, the magnitude of the DEP was reduced approximately 14% and 53%, when the height was increased to 30 and 60 μm, respectively. The results in Fig 4a revealed that the DEP force decrease with the height, however, the force is still strong enough within the insulator, where the height is less than 60 μm. The shorter the distance is, the larger gradient of E 2 is, which implied stronger DEP force. The scanning electron microscopy (SEM) image of microfabricated trapping chip with the open-top microchannels was shown in Fig. 5. The experimental results were also performed in the present work. The experimental set up was established as shown in Fig. 6 (the tips are 100 μm apart). According to the preliminary results obtained experimentally, the 3T3 (Mouse embryonic fibroblast cell line) cells suspended in the sucrose solution (8.62 % weight percent; and its conductivity is 17.6 μS/cm) could be successfully trapped at the tip when applying the electric field of 3×10 4 V/m, and the AC frequency of 1 MHz (Fig. 7). The insulator-based DEP trapping microchip with the open-top microchannels was designed and fabricated in this work. The cells trapped by DEP force could be further treated or cultured in the open-top microchannel; however, those trapped in the microchip with enclosed microchannels could not be proceeded easily.

ISBN: 978-2-35500-006-5

Figure 4 :

 4 The numerical results of the gradient of the square of electric field along the centerline between the tips: (a) at different heights (the gap between tips 60 μm and the angle of the tip 30 degree); (b-d) for different distance between tips at heights of 0, 30 and 60 μm, respectively. The applied electric field is 3×10 4 V/m.

Figure 1 :

 1 Schematic diagram of cell trapping chips with open-top microchannels.

Figure 2 :

 2 Figure 2: The fabrication processes of the cell trapping microchips: fabrication of (a) electrode on glass slide; (b) microchannel; and (c) bonding.

Figure 5 :Figure 3 :

 53 Figure 5: The scanning electron microscopy (SEM) image of the microfabricated trapping device.

Figure 6 :

 6 Figure 6: The schematic illustration of the experimental set up.

Figure 7 :

 7 Figure 7: The experimental results of cell trapping using insulator-based DEP: (a) AC field is off; (b) AC field is on, the applied electric field is 3×10 4 V/m, the frequency of the AC field is 1 MHz. (a) (b)

ACKNOWLEDGMENT

The authors would like to thank the National Science Council of the Republic of China for financial support of this research under contract No. NSC-96-2221-E-194-053.


Cheng-Chih Hsueh

Gou-Jen Wang

Shan-Hui Hsu

Huey-Shan Hung

Fabrication of Nanostructured PLGA Scaffolds Using Anodic Aluminum Oxide Templates

come   L'archive ouverte pluridisciplinaire

INTRODUCTION

Nanotechnology refers broadly to the use of materials and systems whose unifying subject is the control of matter at the atomic, molecular and supramolecular levels, and the fabrication of devices ranging from 1 to 100 nm. Nanobiotechnology is the branch of nanotechnology that has biological and biochemical applications or uses. Nanobiomaterial is one of the crucial fields in the study of nanobiotechnology. The major applications of nanobiomaterial are the fabrication of scaffolds for tissue engineering [START_REF] Luis | PLGA 90/10 and Caprolactone biodegradable Nerve guides for the Reconstruction of the Rat sciatic nerve[END_REF] [START_REF] Wang | Fabrication of PLGA Microvessel Scaffolds with Circular Microchannels Using Soft Lithography[END_REF] and carriers for drug delivery [START_REF] Sonaje | Development of biodegradable nanoparticles for oral delivery of ellagic acid and evaluation of their antioxidant efficacy against cyclosporine A-induced nephrotoxicity in rats[END_REF][START_REF] Mittal | Estradiol loaded PLGA nanoparticles for oral administration: effect of polymer molecular weight and copolymer composition on release behavior in vitro and in vivo[END_REF][START_REF] Ryu | Fabrication of multi-layered biodegradable drug delivery device based on micro-structuring of PLGA polymers[END_REF][START_REF] Csaba | PLGA:Poloxamer and PLGA:Poloxamine Blend Nanoparticles: New Carriers for Gene Delivery[END_REF]. The former refers to the fabrication of nanostructured patterns on a biomaterial through chemical processes. The efficacies of the subsequent cell culture and tissue growth using the nano patterned biomaterial can thus be enhanced; the latter involves the fabrication of nano-scale biodegradable materials as carriers to deliver a particular drug to a specific tissue.

For the applications of biomaterials in tissue engineering, the scaffolds eventually have to be implanted in a live body. Accordingly biodegradability and biocompatibility are the main considerations. Among those commercially available biomaterials, PLGA has been successfully utilized as a biodegradable material because it undergoes hydrolysis in the body to produce the original monomers, lactic acid and glycolic acid, and hence gives rise to very minimal systemic toxicity. Nano patterned PLGA even has great application potentials in tissue engineering. Savaiano and Webster [START_REF] Savaiano | Altered responses of chondrocytes to nanophase PLGA/nanophase titania composites[END_REF] investigated long-term functions and intracellular responses of chondrocytes on novel nanophase PLGA/titania composites, indicating that nanostructured PLGA/nanophase titania composites possess properties to alter chondrocyte responses. Park et al. [START_REF] Park | Accelerated chondrocyte functions on NaOH-treated PLGA scaffolds[END_REF] used fabricated 2D and 3D NaOH-treated PLGA scaffolds and demonstrated that NaOH-treated PLGA scaffolds enhanced chondrocyte functions compared to non-treated scaffolds. Webster et al. [START_REF] Webster | Adhesion of Pseudomonas fluorescens onto nanophase materials[END_REF] compared the adhesion of pseudomonas fluorescens on nanophase with that on conventional grain size alumina substrates. Results showed that nanostructured surface feature is the key material property to enhance bacterial adhesion. Miller et al. [START_REF] Miller | PLGA nanometer surface features manipulate fibronectin interactions for improved vascular cell adhesion[END_REF] formulated highly-controllable, nanostructured features on PLGA surfaces using a mica mold that was patterned by different sizes of polystyrene nanospheres. Results demonstrate that PLGA with nanometer spherical features promoted fibronectin spreading and subsequent vascular cell adhesion. Min et al. [START_REF] Min | Formation of nanostructured poly(lactic-co-glycolic acid)/chitin matrix and its cellular response to normal human keratinocytes and fibroblasts[END_REF] employed the simultaneous electrospinning process to fabricate a composite matrix with chitin nanoparticles being embedded within a PLGA nanofiber matrix. Results indicated that the PLGA and PLGA/chitin matrices are good matrices for normal human fibroblasts.

The published studies regarding the fabrication of nanostructured features on PLGA surfaces can be categorized into the NaOH based chemical etching process and the physical molding process using a nanoparticle patterned replica. The chemical etching process encounters the problems of disorderly nanostructure and NaOH remaining. The physical molding process however requires more complicated procedures. The objective of the present study was to present a simply physical process for the fabrication of orderly nanostructured PLGA membranes. In this novel fabrication approach, an anodic aluminum oxide (AAO) film was use as the template; the PLGA solution was then cast on it; the vacuum air-extraction process ISBN: 978-2-35500-006-5 was applied to transfer the nano porous pattern from the AAO membrane to the PLGA membrane and form nanostructures on it. The pattern and size of the nanostructures can be well controlled through the AAO template.

II. MATERIALS AND METHODS

Materials

The physical processes involve in this study are the PLGA solution casting and the vacuum air-extraction. An AAO membrane was employed as the template for casting. A prepared PLGA solution was then cast on it. After the solvent was completely evaporated, the sample was put in a vacuum oven. The vacuum air-extraction was subsequently conducted to suck the semi-congealed PLGA into the nanopores of the AAO membrane. Stripping off the AAO template, orderly nanorods of PLGA could be obtained. In this manner, the preparation of materials includes the fabrication of the AAO templates with different pore diameters and the producing of the suitable PLGA solution.

A. AAO film preparation

The AAO films were prepared by using the well known anodizing process. Aluminum foils (99.9995%, 175 μm thick) were cleansed using ethanol and acetone in sequence, followed by annealing at 400°C for 3 hours in a vacuum. Electropolishing was then applied by using a sulfuric acid (96 wt %) and phosphoric acid (85 wt %) mixing solution as the electrolyte under a constant voltage of 20 V at 40 °C for 14 min to further polish the surfaces of the foil. An AAO film with nanopore diameter around 50 nm and thickness of 20 μm was obtained by anodizing aluminum foils in a 0.3 M phosphoric acid solution under 140 V of applied voltage at 0°C for 10 hours. An oxalic acid anodization process was also conducted under the conditions of 60 V of applied voltage at 0°C for 4 hours to obtain an AAO film with nanopore diameter around 20 nm. The remaining aluminum beneath the barrier layer was dissolved in an aqueous CuCl 2 HCl solution that was prepared by dissolving 13.45 g powder of CuCl 2 into a 100 ml of 35 wt% hydrochloric acid solution. Following, immerse the bottom surface into the 30 wt% phosphoric acid for 120 minutes to etch out the barrier layer.

Figure 1 shows the SEM morphology of the phosphoric acid and oxalic acid processed AAO films. The nanopore diameters are around 50nm and 20nm, respectively. The AAO films were available as the templates for the vacuum air-extraction.

B. PLGA solution preparation

The PLGA solution was prepared by dissolving 85/15 POLY, IV : 1.6-1.99 (dl/g), Mw:350000-500000 (Da) (Bio Invigor Corp., Taiwan) in acetone in a 1:4 w/w ratio, by stirring the mixture with a magneto agitator at 60°C for 60 min. The PLGA solution was then shaken by an ultrasonic shaker for 15 min to remove bubbles created during mixing. Others solvents such as chlorinated solvents, tetrahydrofurn, acetone or ethyl acetate, and dioxane can also be implemented to dissolve PLGA.

Orderly nanostructure imprinting

Figure 2 schematically illustrates the nanostructure imprinting scheme using an AAO template. The PLGA solution was cast on the AAO template, lasting for 1-2 hours until the solvent was completely evaporated and the surface had solidified. Following, the sample was placed in an oven undergoing the vacuum air-extraction with conditions of vacuum pressure being 76cm-hg at 40°C for 24 hours. The semi-congealed PLGA was sucked into the nanopres of the AAO template to form a nanorods array of PLGA.

AAO

PLGA

Vacuum sucking

Fig. 2. Schematic illustration of the nanostructure imprinting scheme Phosphoric acid can be applied to etch off the AAO template. However, we found that the PLGA membrane could be easily separated from the AAO template in cold water. Specifically, the sample was immersed in cold water for 2-3 minutes and then the PLGA membrane was peeled off. Figure 3 displays the AFM image of a nanopatterned PLGA membrane using an oxalic acid processed AAO template. The AFM image of a nanopatterned PLGA membrane produced by a phosphoric acid processed AAO template is shown in Figure 4. From both images, we can easily observe the nanostructures on the PLGA membrane. Bovine endothelial cells (BEC) were used in the experiments to examine the adhesion capability of the presented nanostructured PLGA membrane. Cells were cultured in standard cell culture conditions (i.e., a 37°C, humidified, 5% CO 2 /95% air environment).

Cell density

Bovine endothelial cells were seeded at a density of 2×10 4 cells/well onto the presented nanopatterned PLGA scaffolds (phosphoric acid processed AAO template produced) and were cultured for 24 and 48 hours. For comparisons, BECs were also seeded on a flat PLGA scaffold. Experimental results shown in Figure 5 indicate that there were almost two-fold of cells on the presented nanopatterned PLGA scaffold after 24 and 48 hours of culture. Results of the present study demonstrated the exhilarating adhesion capability of the presented nanostructured PLGA membrane.

Cell cytoskeleton labeling

The major challenge in observing the seeding cells on a PLGA scaffold is that the PLGA scaffold progressively turned opaque due to absorption of the cultivation medium into the PLGA and hydrolysis of the PLGA. Commonly, a PLGA scaffold becomes completely milky and opaque within 24 hours, making it difficult to monitor the progress of cell seeding by an optical microscopy (OM). To overcome the inability to effectively monitor the seeding cells, a suitable labeling process has to be conducted. The labeling procedures adopted in this study are briefly described as follows.

IV. CONCLUSION

A simply physical process for the fabrication of orderly nanostructured PLGA membranes was presented in this study. An anodic aluminum oxide (AAO) film was use as the template together with the vacuum air-extraction process to imprint orderly nanostructures on a PLGA substrate. The cell culture experiments of the bovine endothelial cells demonstrated that the nanostructured PLGA membrane can largely increase the cell growing rate. Compared to the conventional chemical-etching process, the physical fabrication method proposed in this research not only is simpler but also does not modify the characteristics of the PLGA.

  images of the AAO films; (a) Phosphoric acid processed (b) Oxalic acid processed

Fig. 3 . 5 Fig. 4 .

 354 Fig. 3. AFM image of a nanopatterned PLGA membrane by an oxalic acid processed AAO template

Fig. 5 .

 5 Fig. 5. The cell adhesion test

  (i) Cell fixing: Gently rinsed the PLGA scaffold with PBS to remove any nonadherent cells. The adherent cells were then fixed with a 4% Paraformaldehyde (Sigma, P-6148, US) for 15 min. The sample was rinsed three times with PBS to completely wash off the culture medium. (ii) Cell punching: The sample was then putting into reactions with a 0.5% Triton X-100 (Sigma,T-9284) for 10 min at room temperature. Aspirate the mixing solution from the well followed by rinsing the well three times with PBS. Holes had been punched on cells, enabling stain to penetrate into the cytoplasm. (iii) Mixed a 1% BSA (bovine serum albumin;United States Biochemical, USB, 10857, US) with the PBS. Injected the mixing solution into the well and processed the blocking reaction for 30 min. (iv) Added stain phalloidin (1: 1000, Sigma) to the well and lasted for 30 min. (v) The last step was pouring a preservation solution into the well and wrapped the well with an aluminum film. The sample was ready for further examination.The fluorescence microscopy (Eclipse 80i, Nikon) images of the cytoskeleton labeling experiments for the flat PLGA scaffold, oxalic acid processed AAO template imprinted PLGA scaffold, and phosphoric acid processed AAO template imprinted PLGA scaffold are shown in Figure6-8, respectively. BECs could adhere and grow healthily on those there PLGA scaffolds, especially on the scaffold imprinted by a phosphoric acid processed AAO template.

Fig. 6 .

 6 Fig. 6. Fluorescence microscopy image of the cytoskeleton labeling experiments for the flat PLGA scaffold

Fig. 7 .Fig. 8 .

 78 Fig. 7. Fluorescence microscopy image of the cytoskeleton labeling experiments for the oxalic acid processed AAO template imprinted PLGA scaffold