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0-7803-5998-4/01/$10.00 @2001 IEEE 527

IMPROVED MICRO-FLOW REGULATOR FOR DRUG DELIVERY SYSTEMS

P. Cousseau, R. Hirschi, B. Frehner, S. Gamper, D. MailleferDebiotech S.A., Av. de Svelin 28, CH-1004 Lausanne, Switzerland

Email: p.cousseau@debiotech.com, Website: www.debiotech.com

ABSTRACT

The paper reports on the design, manufacturing, and

experimental testing of a micromachined pressure compensatingflow regulator. This device was designed to provide a constant

liquid flow rate of 1 ml/hr within a pressure difference of 100 to

600 mbar. At pressures higher than 600 mbar the device isdesigned to block the flow, preventing an over-delivery of

medicine. Structural and fluidic simulations were used to designthe geometry of the device before manufacturing. After

manufacturing, over 50 devices have been characterizedexperimentally. The experimental results demonstrate thecredibility of the design, the accuracy of the flow rate and the long-

term stability of the device. One application of this device is the

replacement of the flow restrictor in an elastomeric infusionsystem, which will increase the accuracy and safety of the drug

delivery system. This pressure compensating flow regulator ispassive, hence it needs no external energy source. The device is

relatively inexpensive to manufacture and is therefore, potentially a

disposable unit in a microfluidic system. Finally, it is small andlightweight, ideal for portable applications.

INTRODUCTION

A pressure compensating flow regulator maintains a

constant flow rate for pressure differences within the operationalpressure range of the device. The work presented in this paper is an

extension of the flow regulator reported at Eurosensors XII [1].

The flow regulator presented here has been independently designedand manufactured. The flow rate of the realized devices isapproximately twice as accurate as has been previously reported.

Figure 1 shows a diagram of the devices geometry. The device isa stack of 3 layers. The center layer is a silicon micromachined

membrane with a through hole in the center. A spiral channel is

micromachined in the bottom layer; the start of the spiral is directlybelow the hole in the membrane. This layer has been made fromsilicon or glass. The top layer is a micropackaging layer of glass

with inlet and outlet holes ultrasonically drilled. Figures 2 and 3show photographs of a realized device.

The working principle of the device is as follows: as thepressure difference across the device increases, the deformation of

the silicon membrane increases which covers more of the channelbeneath the membrane. The increase in channel length, and

therefore flow resistance, balances the increase in pressure andthus, a constant flow rate is maintained. Changing the depth and

width of the channel alters the flow rate. Changing the thickness ofthe membrane modifies the operational pressure range.

Debiotech is interested in developing this technology toimprove the accuracy and safety of drug delivery for elastomeric

drug infusion systems. This is not the first time a micromachined

device has been proposed for infusion systems. Precision, silicon

micromachined flow restrictors have been studied for medical

infusion therapy [2]. The major difference between a flow

restrictor and a pressure compensating flow regulator is that flow

restrictors do not compensate the flow rate for pressure variations.

Figure 1. Diagram showing the 3-layer architecture of the

micromachined flow regulation device.

Figure 2. Bottom view of a realized device showing the spiral

channel etched in glass.

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Figure 3. Top view of the device showing the top of the siliconmembrane with the outlet connector glued on the bottom left.

DEVICE DESIGN

Before drawing the masks for the device, extensive computersimulations were performed to establish the appropriate geometry

of the device. Theoretically, the geometry of the device can bescaled to provide any desirable flow rate, therefore, a target flow

rate of 1 ml/hr was chosen because of its suitability for infusiondrug delivery. Two types of simulations were performed: structuraland fluidic. The structural simulations were used to predict the

radius of contact between the membrane and the channel wafer as a

function of the pressure difference across the device. The fluidic

simulations were used to verify the flow rate through the device for

specific membrane deflections.The structural simulations were preformed using the

commercially available software package ANSYS, version 5.3. An

axially symmetric model of the membrane was used. The boundaryof the membrane was considered clamped. A pressure was applied

normal to the upper surface of the membrane. This pressure

corresponds to the pressure difference across the whole device.The channel wafer was modeled by rigid contact elements, i.e., the

membrane deflection was constrained to be less than or equal to thegap between the membrane and channel wafer. The nonlinear

effects of stress stiffening and geometric nonlinearities were

included in the simulations. For the simulated geometry, it was

found that these nonlinear effects are important, with up to a 25%difference between the simulated results and a linear analyticsolution. Different mesh densities and different values for the

convergence criteria were used to verify the stability of thesimulated results. The contact radius between the membrane andthe bottom wafer was solved for several different pressures. With

this data the shape of the spiral channel was calculated.

After completing the structural simulations, the deflected shapeof the membrane is known for a given pressure difference and

therefore the complete geometry of the fluidic path of the device isknown. Using this information, a three-dimensional fluidic model

was created. The CAD software Catia was used to construct the

mesh of the flow path, which included the volume above themembrane, the through hole in the membrane, the micro-channel,

the space between the bottom of the membrane and the spiralwafer, and the outlet hole. The program TASCflow was used to

solve the incompressible Navier-Stokes equations for the above

geometry. The fluidic simulations were necessary to verify that the

pressure drop of the device was mainly in the micro-channel andthat there were no other unforeseen large pressure drops, for

example at the small gap where the bottom of the membrane comesin contact with the channel wafer or through the small through hole

in the membrane. The fluidic simulations showed that the otherpressure drops in the device (excluding the micro-channel)

amounted to less than 2% of the total pressure loss and hence did

not affect the flow rate.

FABRICATION

The flow regulator is a stack of three wafers (Fig. 1). Thecentral wafer is the silicon membrane made with a three masksprocess (shown in Fig. 4). First, the silicon oxide is dry etched

(Fig. 4b) on the bottom in order to pattern the membrane area. Aphotolithography step is then processed, followed by dry etching ofthe central hole (Fig. 4c) in the membrane. It is important that the

depth of this step be deeper (by at least 15%) than the finalmembrane thickness to end with a through hole in the membrane.

After removing the photoresist, the membrane area is patterned by

fluorine-based plasma techniques (Fig. 4d). The upper part is thenprocessed after an oxide has been grown on the wafer. For this

purpose we pattern (Fig. 4e) the oxide by dry etching, using TMAH

to etch the silicon (Fig. 4f). First, a timed-etch is used to roughlyreach the membrane thickness. Additional small duration etches

were used each followed by measurement to reach a precisemembrane thickness.

For the spiral wafer, we use only a dry etching step (Fig. 5) in

silicon. The purpose is to pattern a spiral channel. And we choseplasma etching in order to keep a maximum design freedom. The

oxide layer is removed from both silicon wafers (membrane andspiral) that are then bonded together (Fig. 5c) using Silicon-Direct-

Bonding. The alignment accuracy is 5 m in order to guarantee

that the through hole in the membrane is directly above the start ofthe micro-channel. The top wafer is in fact only a Pyrex wafer with

the inlet and outlet holes drilled by ultrasound, which is bonded to

the stack of membrane and spiral wafers using anodic bonding.The Pyrex wafer allows for metallic fluid connects to be glued

directly onto the device. The majority of the wafer processing forthese devices was done at microFAB in Bremen, Germany;

additional processing was done at the CMI (Center ofMicrotechnology) at the EPFL (Swiss Federal Institute ofTechnology) in Lausanne, Switzerland.

EXPERIMENTAL RESULTS

Several different batches of devices with different membrane

thicknesses and channel profiles have been realized. Over 50devices have been characterized experimentally. For measuring theflow rate of the device as a function of pressure, the gravimetric

method was used. The standard test setup consist of a bottle of

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application, a small membrane filter can be included in the system

or an on-chip particle filter can be added to the device, for example

the particle filter reported in MEMS-99 [3].

.

Figure 7. Comparison of the flow rate from a pressurecompensating flow regulator (nominal flow rate of 1.3 ml/hr) and a

flow restrictor for a sinusoidal pressure between 200 and 400 mbar

with a period of 300 s.

Figure 8. Long-term stability of the flow rate of the flow

regulator (nominal flow rate of 0.54 ml/hr) over a 26 day period.

If a valve is placed in series with the flow regulator, then a

dosing system independent of the reservoir pressure is created.Figure 9 shows the amount of liquid dosed when the valve is open

for 100 s for different pressures. Within the operating pressurerange of the flow regulator, the amount of liquid dosed is

independent of pressure. This is not the case when a flow restrictor

is used; the amount of liquid dosed varies linearly with pressure.Using this system, a lightweight portable dosing system can berealized which dispenses a predicable volume of liquid without the

need to measure the pressure in the reservoir or the outlet flow rate

Figure 9. Comparison of the amount of liquid dosed in 100 sbetween a pressure compensating flow regulator and a flow

restrictor for different reservoir pressures.

CONCLUSIONS

This paper has presented a realized and fully operational

microfluidic pressure compensating flow regulator. It has beendemonstrated that this device maintains a constant flow rate for

pressures between 200 and 500 mbar. Also, the long-term stability

of the device has been demonstrated by 26 days of continuousoperation with no degradation of the flow rate. The device can be

used with a valve for liquid dosing applications. Also, the valve

could be duty cycled to generate a variable flow rate device. In the

future, manufacturing the device in plastic and reducing the size in

silicon will be studied. It is believed that flow regulators couldpotentially have a place next to microvalves, channels, mixers and

pumps in Tas applications and drug delivery systems, especially

portable devices, which would benefit from the lightweight andsmall size. Debiotech is interested in using this micromachined

flow regulator to replace the less accurate conventional glass

capillary flow restrictors used in portable elastomeric drug infusionsystems

REFERENCES

[1] Ch. Amacker, Y.-S. Leungki, V. Pasquier, Ch. Madore, M.

Haller, and Ph. Renaud, Passive Micro-Flow Regulator for

Drug Delivery System, Proceedings of Eurosensors XII,September 13-16, 1998, pp. 591-594.

[2] J. Drake, and H. Jerman, A Precision Flow Restrictor for

Medical Infusion Therapy, Proceeding of the 8th International

Conference on Solid-State Sensors and Actuators andEurosensors IX, Stockholm, Sweden, June 25-29, 1995, pp.373-376.

[3] D. Maillefer, H. van Lintel, G. Rey-Mermet, and R. Hirschi, A

High-Performance Silicon Micropump for an Implantable DrugDelivery System, Proceedings of the 12th IEEE Internal

Conference on Micro Electro Mechanical Systems, Orlando,

Florida, USA, January 17-21, 1999, pp. 541-546.