Copyright  (2013)  American  Institute  of  Physics.  This  article  may  be  downloaded  for  personal  use  only.    Any  other  use  requires  prior  permission  of  the  author  and  the  American  Institute  of  Physics.    The  following  article  appeared  in  Journal  of  Applied  Physics  (Vol.114,  Issue  2)  and  may  be  found  at  http://jap.aip.org/resource/1/japiau/v114/i2/p024701_s1  or    http://dx.doi.org/10.1063/1.4813484      This  Article  is  copyright  the  American  Institute  of  Physics  (AIP)  Publishing  LLC      AIP Publishing LLC 2 Huntington Quadrangle Melville NY USA Email: jrnlpub@aip.org http://journals.aip.org  

Study on an alternating current electrothermal micropump for microneedle-based fluid delivery systemsRumi Zhang, Graham A. Jullien, and Colin Dalton Citation: J. Appl. Phys. 114, 024701 (2013); doi: 10.1063/1.4813484 View online: http://dx.doi.org/10.1063/1.4813484 View Table of Contents: http://jap.aip.org/resource/1/JAPIAU/v114/i2 Published by the AIP Publishing LLC. Additional information on J. Appl. Phys.Journal Homepage: http://jap.aip.org/ Journal Information: http://jap.aip.org/about/about_the_journal Top downloads: http://jap.aip.org/features/most_downloaded Information for Authors: http://jap.aip.org/authors

Study on an alternating current electrothermal micropumpfor microneedle-based fluid delivery systems

Rumi Zhang,1 Graham A. Jullien,1 and Colin Dalton2,a)

1Advanced Technology Information Processing Systems (ATIPS) Laboratory, Electrical and ComputerEngineering, University of Calgary, Calgary, Alberta T2N 1N4, Canada2Advanced Micro/Nanosystems Integration Facility (AMIF), Electrical and Computer Engineering,University of Calgary, Calgary, Alberta T2N 1N4, Canada

(Received 4 January 2013; accepted 23 June 2013; published online 11 July 2013)

In this paper, we report on a modeling study of an AC electrothermal (ACET) micropump with

high operating pressures as well as fast flow rates. One specific application area is for fluid delivery

using microneedle arrays which require higher pressures and faster flow rates than have been

previously reported with ACET devices. ACET is very suitable for accurate actuation and control

of fluid flow, since the technique has been shown to be very effective in high conductivity fluids

and has the ability to create a pulsation free flow. However, AC electrokinetic pumps usually can

only generate low operating pressures of 1 to 100 Pa, where flow reversal is likely to occur with an

external load. In order to realize a high performance ACET micropump for continuous fluid

delivery, applying relatively high AC operating voltages (20 to 36 Vrms) to silicon substrate ACET

actuators and using long serpentine channel allows the boosting of operating pressure as well as

increasing the flow rates. Fast pumping flow rates (102–103 nl/s) and high operating pressures

(1–12 kPa) can be achieved by applying both methods, making them of significant importance for

continuous fluid delivery applications using microneedle arrays and other such biomedical devices.VC 2013 AIP Publishing LLC. [http://dx.doi.org/10.1063/1.4813484]

I. INTRODUCTION

With microfluidics, one of the significant technological

challenges is the precise control and manipulation of microli-

ter or nanoliter volumes of fluids. One of the key compo-

nents is the miniature fluid-dispensing system or micropump.

Over the past few decades, numerous papers have been pub-

lished on the development and fabrication of reciprocating

displacement micropumps, which use the oscillatory move-

ment of mechanical parts to transfer mechanical energy into

fluid movement.1,2 However, a major problem existing in

reciprocating displacement micropumps is that fluid from

these micropumps is delivered in a series of small discrete

volumes, which make up a pulsating flow.1,3 For a pulsating

flow, flow fluctuation is relatively large at low flow rates,

and therefore the accurate adjustment of uniform flow rates

might be difficult.3 Also, mechanical fatigue caused by wear

and stiction of moving parts in reciprocating mechanical

micropumps often restrict the operational environment and

limit the lifetime of these devices.3 Micropump failure can

also be caused by other problems such as lack of tolerance to

bubbles, leaks in the diaphragm, or valve causing pressure

loss, structural complexity, and slow response time.3,4

In previous studies, electrokinetics has been demon-

strated as having great potential for microfluidic actuation in

bio-MEMS and lab-on-a-chip systems, due to its basic imple-

mentation efficiency and reliability from having neither

mechanically moving parts nor valves to drive the fluid

flow.5–9 Compared to mechanical micropumps, advantages of

electrokinetic micropumps include simplicity in design and

fabrication, longer life cycle, and minimal issues with block-

age.10 More importantly, they do not have any moving parts,

and they are therefore capable of creating a pulsation free

flow, i.e., a very precise continuous steady fluid flow.9,14

Electrokinetic techniques with AC electric sources, such as

AC electrothermal (ACET), AC electro osmosis (ACEO), or

dielectrophoresis (DEP), can easily achieve high electric field

strengths for on-chip micropumping and particle manipula-

tion with only a few volts (�10 Vrms).3,11–14 Also, electrolysis

is avoided when an AC source is applied,9 since an AC

source avoids a net movement of ions within the solution.

In this paper, we carry out a modeling study of an ACET

micropump for the requirements of high operating pressures

as well as fast pumping flow rates. One specific application

area is for fluid delivery using microneedle arrays. In previous

studies of microneedles, reciprocating displacement micro-

pumps excited by piezoelectric or thermopneumatic actuators

have been designed for microneedle fluid delivery sys-

tems.4,18,19 It has been observed in Ref. 18 that net water flow

rates of approximately 2.0 nl/s out of microneedles were

obtained at a pressure of 3.9 kPa for continuous on-chip

micropumping. Other studies20,21 also suggested that micro-

needles be operated in the kPa range of pressures. ACET has

been shown to be very effective in high conductivity fluids

and therefore is useful for biological applications, which fre-

quently involve samples with conductivities higher than 0.1 S/m

(for example, saline solution has a conductivity of 1–2 S/m).14–17

ACET micropumping has been demonstrated at low voltages

(< 15 Vrms) for fluid conductivity of 0.02–1 S/m with

fluid velocity of 100–1000 lm/s.17 With this, ACET micro-

pumps are capable of pumping 0 to 500 nl/s fluid flow, and

the precisely controlled flow velocity can be achieved bya)Electronic mail: cdalton@ucalgary.ca.

0021-8979/2013/114(2)/024701/8/$30.00 VC 2013 AIP Publishing LLC114, 024701-1

JOURNAL OF APPLIED PHYSICS 114, 024701 (2013)

adjusting the AC potential between electrode pairs.

However, we should note that the existing AC electrokinetic

pumps including ACET micropumps usually can only gener-

ate low operating pressures of 1 to 100 Pa, and thus flow

rates will decay rapidly and cannot even prevent flow rever-

sal against an external load.22

Recently, Huang et al. proposed a high-pressure (about

1.3 kPa) ACEO micropump, in which a long serpentine chan-

nel was used to boost its pressure.22 However, ACEO is not

effective in fluids with high electrical conductivities, which

excludes most biofluidic applications. Also, the microchannel

present in Ref. 22 is only 100 lm (W) and 25 lm (H), too

small to have high volume fluid flow rates (Q¼ outlet cross-

section area� velocity), as required for microneedle arrays or

other such microfluidic devices. It has been reported that

experiments using microneedles for intradermal delivery of

sterile salin23 used a flow rate of 0.3 ml/min (5000 nl/s). As

well, Roxhed et al. reported that a flow rate of 2.0 ll/h

(0.56 nl/s) ensured a leak-free delivery for insulin into rat der-

mis.24 Such a study was carried out using an array of 21 hol-

low microneedles on an effective area of 2� 2 mm2 only.

However, as stated in Ref. 24 such a low infusion rate is

impractically low for most human drug delivery applications

(for example, a representative real-time need of 10 units typi-

cally absorbed on the time scale of an hour for a conventional

formulation of 100 units/ml insulin for diabetic patients,

�27.8 nl/s), and, hence, the rate needs to be increased to

make the technique useful.24 As the total infusion rate

depends on the number of needles, using a large array with

an effective area of a few cm2 will allow much higher deliv-

ery rates (14 nl/s with 1� 1 cm2, 56 nl/s with 2� 2 cm2, or

126 nl/s with 3� 3 cm2, respectively) for leak-free delivery.24

In our study, we examine two means for increasing the oper-

ating pressures: applying relatively high AC operating vol-

tages (20 to 36 Vrms) to ACET actuators with silicon

substrates and increasing the channel length with a long ser-

pentine channel. Studies show that operating pressures of

1–12 kPa and pumping flow rates of 102–103 nl/s can be

achieved. This does not account for the reduction of flow

rates expected in the transition from the ACET actuator to the

complex channel geometry of the microneedle array. The

pressure and flow rate drop required to flow fluid through a

microneedle depends on needle geometry and fluid viscosity

and density.25 Needles of micron dimensions can exhibit sig-

nificant resistance to flow.25 Also, flow resistance into tissue

is expected during a microneedle injection into skin, and it

has been reported that the presence of skin tissue decreases

the flow rate by one or more orders of magnitude.25 Since the

flow rate decays linearly with the back pressure, P, improve-

ment in both pumping flow rates and pressures are of signifi-

cant importance for ACET micropumps when implemented

for continuous fluid delivery using microneedle arrays and

other microfluidic applications.

II. ACET FLUID FLOW

ACET refers to electrothermal flow induced by a non-

uniform thermal field generated in a fluid, i.e., a temperature

gradient, rT, in the presence of AC electric fields. The

temperature gradient, rT, can be obtained from Joule heat-

ing in Eq. (1) (Ref. 27)

kr2T þ rjEj2 ¼ 0; (1)

where r and k are the electrical conductivity and the thermal

conductivity of the fluid, respectively. jEj is the magnitude

of the electric field applied over the fluid, determined by the

gradient of the AC potential, V

E ¼ �rV: (2)

The time average electric force on the fluid is described

as27

hfeti ¼1

2

eða� bÞ1þ ðxsÞ2

ðrT � EÞE� 1

4eajEj2 � rT; (3)

where a ¼ 1=eð@e=@TÞ; b ¼ 1=rð@r=@TÞ, in which r and eare the electrical conductivity and permittivity of the me-

dium, respectively. s ¼ r=e is its charge relaxation time,

and x ¼ 2pf is the angular frequency of the applied AC

potential.

For an incompressible fluid of low Reynolds number,

the resulting steady fluid flow in the microchannel follows

the Navier-Stokes equations9,36

�rpþ gr2uþ hfeti ¼ 0; r � u ¼ 0; (4)

where u denotes the velocity field vector. g is the dynamic

viscosity, and p is the pressure, respectively. Equations

(1)–(4) provide the functional relationship between fluid ve-

locity and temperature gradient produced by the applied AC

electric field. For a more detailed explanation of ACET fluid

flow, please refer to Ref. 31.

III. DESIGN OF AN ACET MICROPUMP

As shown in Fig. 1, the new ACET micropump consists

of multiple unequal width electrodes located at the bottom of

a long serpentine channel forming an interdigitated array,

with AC voltages applied between the narrow and the wide

electrodes. In the following sections, we will explain the

detailed design of the new ACET micropump, and how it is

able to achieve faster pumping flow rates as well as higher

operating pressures than previously reported designs.

A. Substrate

Substrate materials for ACET micropumps can be glass,

silicon, or other materials, such as ceramics. Coplanar elec-

trode arrays are positioned on the surface of the substrate,

which acts as a heat sink. The electric fields generated by the

electrode arrays are much stronger around the electrode arrays

than in the bulk fluid, and thus more heat is generated there.17

Part of the heat spreads into the substrate by conduction and

then from the underside surface of the substrate into the ambi-

ent air. Thus, the thermal properties of the substrate have a

significant effect on ACET micropumping abilities. Part of

024701-2 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

the heat can also spread into the walls of the microfluidic

channels, which can be glass, or polymer such as polydime-

thylsiloxane (PDMS). However, this effect is negligible, since

less heat is generated there and passed through the thermally

insulated channel walls.

Due to less heat escaping through a thermally resistant

substrate, a higher temperature rise can be obtained when

even a small AC potential is applied. Glass, with a low ther-

mal conductivity of k � 0:6 W=K � m,30 is a low-cost,

highly efficient thermally resistant material for the tempera-

ture rise acceptable for ACET micropumping. However, a

high thermally resistant substrate can cause a high tempera-

ture rise in certain regions where a strong electric field is

applied, which is sometimes detrimental to ACET fluid

flow. Glass substrate is limited by this excessive tempera-

ture rise, and they are usually used in conditions of very low

AC voltages (<10 Vrms).

Alternatively, thermally conductive materials, such as

silicon, with k � 140 W=K � m,30 exhibit a relatively small

temperature rise in the fluid because the thermally conduc-

tive substrate removes a large amount of heat generated by

the strong electric field. More interestingly, since ACET

devices are generally operated at low voltages, thermally

conductive substrates are able to obtain much faster pumping

flow rates by increasing the amplitude of the AC voltages.

The ability to achieve faster fluid flow, yet maintain a rel-

atively low temperature rise, is important for ACET micro-

pumps in the application area of transdermal or implanted

drug delivery. For this reason, silicon substrates are used for

the ACET micropump design and relatively high AC voltages

are applied to multiple electrode pairs to obtain fast pumping

flow rates with a low temperature rise.

B. PDMS serpentine channel

As stated previously, most existing AC electrokinetic

pumps can only generate relatively low operating pressures

of 1 to 100 Pa, which are insufficient to pump fluid into a

device such as a microneedle array. For any pump operating

in the viscous regime of a low-Reynolds number, the flow

rate decays linearly with the back pressure, P, according to

Eq. (5) (Ref. 26)

Q

Qmax¼ 1� P

Pmax; (5)

where Qmax is the flow rate at P¼ 0, and Pmax is the back

pressure that yields Q¼ 0 and effectively stops the pump.26

Here, we consider flow in a microchannel of rectangular

cross-section with width, W, and depth, H, where H � W.

The total flow rate due to a mean slip velocity, U, over the

bottom surface is given by Eq. (6) (Ref. 26)

Qmax ¼aHWU

2; (6)

where a � 1� ð1=2ÞðH=WÞ2 corrects for fringe flows in the

limit H � W. The back pressure, Pmax, required to stop the

net forward flow is given by26

Pmax ¼ RBQmax ¼UL

k¼ 6gaUL

H21þ H

W

� �2" #

; (7)

where k is the hydrodynamic Darcy permeability and L is the

length of the microchannel. RB is the back-flow hydraulic

resistance.

From Eq. (7), we see that Pmax can be increased by

reducing the micro-channel size; however, the depth cannot

be reduced too much, due to high flow volume rates required

for a microneedle array.

On the other hand, Pmax can easily be boosted by

increasing the channel length, L, by using a long serpentine

channel, according to Eq. (7). Using a standard PDMS chan-

nel on a single microfluidic chip, the pump length, and thus

the pressure, can be theoretically boosted by two orders of

magnitude.

C. Overall design

The ACET electrode arrays consist of a 20 lm narrow

electrode and a 120 lm wide electrode with a gap of 20 lm.

Multiple electrode pairs are separated by 150 lm. Each

�30 mm long ACET electrode array consists of 96 electrode

pairs. For an ACET micropump with a pump area of

4 cm� 4 cm, as shown in Fig. 1, one can fabricate about

forty 30 mm long ACET electrode arrays (3840 electrode

pairs in total) via a 1200 mm long serpentine channel where

the fluid microchannel is 500 lm wide and 100 lm deep. The

coplanar electrode array design reduces the complexity of

FIG. 1. An ACET micropump consisting of multiple asymmetric electrode

arrays with long serpentine channel. Electrode arrays consist of a 20 lm

narrow electrode and a 120 lm wide electrode with a gap of 20 lm.31 The

electrode pairs are separated by 150 lm. The bottom substrate is silicon

(thickness¼ 1000 lm). On top of the substrate, there are electrode arrays

with a thickness of 100 nm. A PDMS serpentine microchannel forms on

the top of the electrodes by soft lithography, and the channel is 100 lm

deep and 500 lm wide.

024701-3 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

the completed devices. Thus, all fluidic components may be

integrated onto one substrate, with a minimum of assembly

for future applications.

IV. NUMERICAL SIMULATIONS

A. Modeling

In this section, numerical simulations are carried out

using COMSOL MULTIPHYSICS software35 to evaluate the per-

formance of the ACET devices. First, the electric field is

solved using Eq. (2). Then the resulting thermal field is cal-

culated from Eq. (1), and finally the induced flow velocity is

solved from the Navier-Stokes equation, Eq. (4). From these

simulations, flow rates as well as operating pressures can be

obtained from Eqs. (6) and (7).

The schematic diagram of modeling domains and the

boundary conditions is given in Fig. 2. Three domains are

used for this study: the bottom 1000 lm thick substrate, the

100 lm thick fluid inside the channel, and finally a lid of a

100 lm thick PDMS layer.33 The boundary conditions for

the left and the right sides of the channel are set to be peri-

odic for electrostatic, thermal, and fluid flow simulation.

Besides these, for the electrostatic simulation, the electric

potentials on the asymmetric electrode pairs are set as Vrms

and 0, while the remainder of the boundaries is defined as

insulated/symmetry.27–31 For the thermal simulation, we

assume a fixed temperature of T¼ 20 �C (293.15 K) to the

top of the PDMS layer and the bottom of the substrate;33 for

the incompressible Navier-Stocks simulation, a non-slip wall

boundary condition is imposed on the fluid channel walls.

ACET actuators with glass and silicon substrates are

simulated at various operating AC potentials to evaluate the

performance of silicon ACET using relatively high poten-

tials. As with,30 the thermal conductivity, k, used for the

glass is 0:6 W=K �m, for the silicon 140 W=K �m, and for

the fluid, kf, is 0:598 W=K �m. Here, we use 1� phosphate

buffered saline (PBS) as an example, which is a buffer solu-

tion commonly used in biological research. The conductivity

and permittivity of 1� PBS buffer are r¼ 0.224 S/m and

e ¼ 80e0, respectively, where the permittivity of free space,

e0 ¼ 8:854� 10�12 F=m. The viscosity of water, g, taken as

a function of temperature T(K), are used to account for vis-

cosity change with rising temperature

g ¼ 2:414� 10�5 � 10247:8=ðT�140Þ: (8)

The number and type of mesh elements are critically im-

portant to obtain the minimum level of discretization to

achieve accurate results. For our simulation, more than

100 000 triangular elements were used in each simulation.

We have also performed a simple mesh-convergence analy-

sis to ensure the convergence of our simulation results. This

required that after solving for one situation, the mesh was

made finer and the simulation re-solved to check that the so-

lution was within acceptable bounds. Less than 0.5% change

was observed for our simulation results when more than

100 000 triangular elements were used. Stable convergence

was achieved in the numerical iteration in all cases.

Fig. 3 gives the simulation result of the 1-pair section,

part of the silicon ACET actuator, where an AC potential of

25 Vrms was applied to its corresponding structures. In this

figure, surface colors, arrows, and streamlines represent the

temperature, the electric field, and the fluid flow, respec-

tively. The temperature rise for the silicon ACET is

293–310.4 K where the volume flow rate obtained is 154 nl/s

(1.54� 108 lm3/s).

B. Preliminary experiments and accuracy

The performance and reliability of the numerical simula-

tion method used in this paper have been demonstrated in pre-

vious studies, in which COMSOL MULTIPHYSICS has proved to

provide good agreement between numerical simulation and

analytical solutions or available experimental data.9,17,28,30,36

In our previous preliminary experiments with glass substrates,

we used 8 electrode pairs of an electrode array for the study,

each consisting of an 18 lm narrow electrode and a 116 lm

wide electrode with a gap of 24 lm; we provided a separation

of 152 lm between electrode pairs, as shown in Fig. 4. The

PDMS microchannel sealed over the electrodes was 1000 lm

wide and 1000 lm deep.

Pumping was observed with fluids of 0.224 S/m when

applying 5 Vrms at 100 kHz. Fluid flow was determined using

time lapse video of beads traversing the array with an appliedFIG. 2. Domains and boundary conditions for solving the electrical field,

thermal field, and fluid flow inside the channel.31

FIG. 3. Numerical simulation result for the 1-pair section, part of the silicon

ACET actuator, where an AC potential of 25 Vrms was applied.

024701-4 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

field. After analyzing dozens of moving latex beads �100 lm

above the ACET electrode array in (Figs. 5(a)–5(d)), the ex-

perimental data shown that fluid velocity was estimated to be

within 35–40 lm/s, compared to 33.2 lm/s in the same loca-

tion obtained from the simulation. The measured experiment

data thus agree with the simulation result, being within 20%.

Differences in the measured and simulated velocity can be

caused by many factors (for example, fluctuations in the

actual size of the fabricated electrodes and PDMS channel,

the resolution of the optical microscope, the size of time step

used in the video capture, and also viscous drag on the beads).

C. Flow rates

Figs. 6(a) and 6(b) show the effect of the applied AC

potentials, Vrms, on volume fluid flow rates, as well as maxi-

mum temperature, Tmax, occurring in the fluid, for both sili-

con and glass substrates, in which the AC potentials are

operated at a frequency of f¼ 100 kHz. Note that ACET fluid

flow is usually operated at frequencies of � 100 kHz. From

Ref. 36, there are three frequency regions (Coulomb force

dominated region, transition domain, and dielectric force

dominated region), and the fluid flow is independent of fre-

quency except in the transition domain. Both devices have

the same dimensions, sizes, and electrode geometry, the only

difference being the thermal conductivity and density of the

substrate, where k ¼ 0:6 W=K �m and q¼ 3000 kg/m3 are

for glass and k ¼ 140 W=K �m and q¼ 2330 kg/m3 are for

silicon, respectively.

From this simulation, we see that, with the same applied

AC potentials, the ACET device with the glass substrate

achieves much faster flow rates than that with the silicon

substrate. For example, for an AC signal with an amplitude

of 9 Vrms, the ACET with the glass substrate achieves a vol-

ume fluid flow rate of 8.30 nl/s (8.30� 106 lm3/s), approxi-

mately 400% more than that with the silicon substrate (2.33

nl/s or 2.33� 106 lm3/s), at the cost of a temperature

increase of about 23.5 �C (Tmax¼ 45.7 �C (318.9 K) for the

glass substrate and Tmax¼ 22.2 �C (295.4 K) for the silicon

substrate). It is noted that the faster fluid flow results mostly

from higher temperature rise due to the good thermal resist-

ance of the glass substrate. Compared to the thermally resist-

ant materials, a higher electric field, i.e., a high AC potential,

is required for the ACET device built on the thermally con-

ductive substrate in order to maintain the same flow rates as

that on the thermally resistant substrate.

As previously discussed, temperature rise is the major

concern in designing ACET actuators, due to the fact that ex-

cessive temperature rise might cause the degradation of

many reagents/analytes and even alter the functions of

biological systems.37 It is very difficult to further increase

the AC potential for the glass ACET devices (e.g., Tmax

> 58.4 �C (331.6 K) for an AC signal with an amplitude of

11 Vrms). In contrast, the silicon ACET devices are able to

enhance fluid flow by increasing the applied AC potentials,

since the maximum temperature is only 22.2 �C for an AC

signal with an amplitude of 9 Vrms.

As shown in Fig. 7(a), volume fluid flow rates increase

dramatically when the applied AC potentials increase. Fig.

7(b) depicts the relationship of volume fluid flow rates and

maximum temperature, Tmax (in which AC potentials for a

glass ACET are in the range of 0 to 11 Vrms). AC potentials of

more than 11 Vrms for a glass substrate ACET will cause the

temperature to rise to undesirable levels (above 60 �C). From

Fig. 7(b), the simulation results show that silicon ACET

devices are capable of achieving much faster pumping flow

rates, more than 20 times the rate of ACET devices built

on glass substrate, for all cases, where both devices are

maintained at the same temperature rise (e.g., 330 nl/s

(3.3� 108 lm3/s) at Tmax¼ 44.9 �C (318 K) for silicon, and 8.3

nl/s (8.3� 106 lm3/s) at Tmax¼ 45.7 �C (318.9 K) for glass).

FIG. 4. An ACET array.31,32

FIG. 5. Time lapse video images of beads traversing the array under an

applied AC field (dark areas are asymmetrical electrodes, latex beads were

used for tracking fluid movement). (a). t¼ 0 s; (b) t¼ 1 s; (c) t¼ 2 s; (d)

t¼ 3 s.

024701-5 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

In general, thermally conductive materials are usually

electrically conductive. For example, silicon substrates have a

very low resistivity due to doping, making high resistivity sili-

con expensive. For this reason, it is very important to grow a

thin insulation layer, such as silicon dioxide (SiO2) or silicon

nitride (Si3N4), on top of a silicon substrate, to prevent a short

circuit occurring when applying AC potentials on the adjacent

ACET electrodes. Moreover, since silicon dioxide is a ther-

mally resistant material with a low thermal conductivity of

k � 1:0 W=K �m, a silicon ACET device coated with differ-

ent thickness of silicon dioxide layers can act as a material

with a thermal conductivity of less than 140 W=K �m.

Fig. 8 shows the effect of a silicon dioxide layer fabri-

cated on silicon ACET electrodes on a silicon substrate.

From Fig. 8, we see that by controlling the silicon diox-

ide thickness, faster pumping flow rates can be achieved due

to a higher temperature rise as a result of a thicker insulation

layer.

D. Operating pressure

For an ACET micropump with a pump area of

4 cm� 4 cm, as shown in Fig. 1, a 1200 mm long serpentine

channel can be fabricated on the top, where the fluid micro-

channel is 500 lm wide and 100 lm deep. Conducting traces

connected to the electrode array can be placed under the

spacing between the microchannel, which is greater than

500 lm. Fig. 9 gives the operating pressure, Pmax, obtained

under the applied AC potentials of 20–36 Vrms.

From Fig. 9, we see that it is easy to achieve higher

operating pressures exceeding 1 kPa, using an ACET micro-

pump with a long serpentine channel. This pressure is high

FIG. 6. Effect of the applied AC

potentials, Vrms, on volume fluid flow

rates, as well as maximum tempera-

ture, Tmax, in the fluid for both silicon

and glass substrate.

FIG. 7. The relationship of volume

fluid flow rates, as well as maximum

temperature, Tmax, occurred in the fluid

when higher AC potentials, Vrms, were

applied onto silicon.

024701-6 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

enough to drive a drug solution into microneedle arrays,

according to previous experimental data.18,19

E. Further discussion

Two important parameters of a micropump are the maxi-

mum pressure, Pmax and the maximum flow-rate, Qmax that

the pump can generate. Overall, an silicon ACET micropump,

with relatively high AC operating voltages (20 to 36 Vrms) and

a 1200 mm long PDMS serpentine channel, can achieve

102–103 nl/s (108–109 lm3/s) fluid flow rates with 1–12 kPa

operating pressures. Our new design achieves pressures of

same order of magnitude, but with flow rates that are faster by

two orders of magnitude when compared to previously

reported results (about 3.25 nl/s flow rate and 1.3 kPa) from

Ref. 22. Also, since the microchannel in Ref. 22 was only

100 lm (W) and 25 lm (H), such a device will experience

substantial pressure drop when fluid flows from such a small

size microchannel into the much larger cross-section channel,

such as those in a microneedle array. In the design reported

here, the fluid microchannel is 500 lm wide and 100 lm deep.

The cross-section area of the microchannel (5� 104 lm2) is

nearly half that of a microneedle array (cross-section area of

about 13� 104 lm2) reported in Ref. 34. Such an ACET

micropump design can provide sufficient pumping flow rates

as well as operating pressures for microneedle fluid delivery.

Wider and deeper fluid microchannels can increase

pumping flow rates; however, this will also increase the diffi-

culty in obtaining a high operating pressure. A prime con-

cern throughout this research is in developing microfludic

systems with relatively simple fabrication processes and

with improved fluidic control. In our design, the gap between

narrow and wide electrode are chosen to be more than 20 lm

so that the size of the electrode arrays can be accurately con-

trolled by standard photolithography. However, we should

note that faster pumping flow rates can be further enhanced

by fabricating smaller gaps between the narrow and wide

electrodes.

Reducing the microchannel depth can also increase the

operating pressure; however, the operating pressure can not

be reduced by too much, in order to maintain the high flow

volume rates (Q¼ outlet cross-section area� velocity)

required for microneedle fluid delivery. Operating pressure

can be further increased by using a larger pump area with a

longer serpentine channel or using multi-layer pumps con-

nected in series.

V. CONCLUSIONS

In this paper, we have investigated an ACET micropump

that has high pressures as well as fast pumping flow rates and

is targeted towards biological fluid delivery using microneedle

arrays. Fast pumping flow rates (102–103 nl/s) as well as high

operating pressures (1–12 kPa) have been achieved using rela-

tively high AC operating voltages (20 to 36 Vrms) applied to a

silicon ACET micropump with a long PDMS serpentine

FIG. 8. Effect of silicon dioxide layer

with different thicknesses.

FIG. 9. Operating pressure obtained under potentials of 20–36 Vrms for an

ACET micropump with 40� 30 mm long serpentine channel where the fluid

microchannel is 500 lm wide and 100 lm deep.

024701-7 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)

channel. Compared to most existing AC electrokinetic pumps

with low operating pressures of 1–100 Pa, such a dramatic

improvement in operating pressure is of significant impor-

tance for continuous fluid delivery applications using micro-

needle arrays. It is hoped that this research will expand the

application scope of electrokinetics in microfluidic chips

where high pressures and fast pumping flow rates with a rela-

tively low temperature rise are necessary, and where ease of

integration and fabrication is preferred.

ACKNOWLEDGMENTS

The authors wish to acknowledge financial support from

the Alberta Ingenuity Fund, the Natural Sciences and

Engineering Research Council of Canada (NSERC), and

iCORE (Alberta). We also wish to acknowledge CMC

Microsystems for providing software support and access to

their microfluidic carrier platform.

1P. Woias, “Micropumps-summarizing the first two decades,” Proc. SPIE

4560, 39–52 (2001).2A. Acevedo, “Creation of a dual chamber micropump using rapid proto-

typing,” Proceedings of the 20th NCUR (2006), published online, http://

www.ncur20.com/default.asp.3W. E. Morf et al., “Partial electroosmotic pumping in complex capillary

systems Part 1: Principles and general theoretical approach,” Sens.

Actuators B 72, 266–272 (2001).4D. J. Laser and J. G. Santiago, “A review of micropumps,” J. Micromech.

Microeng. 14, R35–R64 (2004).5A. Castellanos, A. Ramos, A. Gonzalez, N. G. Green, and H. Morgan,

“Electrohydrodynamics and dielectrophoresis in microsystems: Scaling

laws,” J. Phys. D: Appl. Phys. 36, 2584–2597 (2003).6N. G. Green et al., “Electrothermally induced fluid flow on micro-

electrodes,” J. Electrost. 53, 71–87 (2001).7J. C. Eijkel, C. Dalton et al., “A circular ac magnetohydrodynamic micro-

pump for chromatographic applications,” Sens. Actuators B 92, 215–221

(2003).8D. Li, Electrokinetics in Microfluidics (Elsevier, 2004).9M. Lian et al., “AC electrothermal manipulation of conductive fluids and

particles for lab-chip applications,” IET Nanobiotechnology 1(3), 36–43

(2007).10F. E. H. Tay, Microfluidics and BioMEMS Applications (Kluwer

Academic Publisher, 2002).11J. T. Lin and J. T. Yeow, “Enhancing dielectrophoresis effect through

novel electrode geometry,” Biomed. Microdevices 9, 823–831 (2007).12J. Wu, “Interactions of electric fields with fluids: laboratory-on-a-chip

applications,” IET Nanobiotechnology 2(1), 14–17 (2008).13F. Malloggi et al., “Electrowetting-controlled droplet generation in a

microfluidic flow-focusing device,” J. Phys.: Condens. Matter 19(46),

462101 (2007).14A. Ramos, H. Morganz, N. G. Greenz, and A. Castellanosy, “AC electroki-

netics: A review of forces in microelectrode structures,” J. Phys. D: Appl.

Phys. 31, 2338–2353 (1998).15G. Fuhr et al., “Traveling wave-driven microfabricated electrohydrody-

namic pumps for liquids,” J. Micromech. Microeng. 4, 217–226 (1994).

16M. Sigurdson, D. Z. Wang, and C. D. Meinhart, “Electrothermal stirring

for heterogeneous immunoassays,” Lab Chip 5, 1366–1373 (2005).17J. Wu, M. Lian, and K. Yang, “Micropumping of biofluids by alternat-

ing current electrothermal effects,” Appl. Phys. Lett. 90(23), 234103

(2007).18J. D. Zahn et al., “Continuous on-chip micropumping for microneedle

enhanced drug delivery,” Biomed. Microdevices 6(3), 183–190 (2004).19Q. Cui, C. Liu, and X. Zha, “Study on a piezoelectric micropump for the

controlled drug delivery system,” Microfluid. Nanofluid. 3(4), 377–390

(2007).20R. Zhang, P. Zhang, C. Dalton, and G. A. Jullien, “Modeling of drug deliv-

ery into tissues with a microneedle arrays using mixture theory,” Biomech.

Model. Mechanobiol. 9(1), 77–86 (2010).21K. Kim and J. B. Lee, “High aspect ratio tapered hollow metallic micro-

needle arrays with microfluidic interconnector,” Microsyst. Technol. 13,

231–235 (2007).22C. C. Huang, M. Z. Bazant, and T. Thorsen, “Ultrafast high–pressure AC

electro-osmotic pumps for portable biomedical microfluidics,” Lab Chip

10(1), 80–85 (2010).23J. Gupta, S. S. Park, B. Bondy, E. I. Felner, and M. R. Prausnitz, “Infusion

pressure and pain during microneedle injection into skin of human sub-

jects,” Biomaterials 32, 6823–6831 (2011).24N. Roxhed, B. Samel, L. Nordquist, and G. Stemme, “Painless drug deliv-

ery through microneedle-based transdermal patches featuring active

infusion,” IEEE Trans. Biomed. Eng. 55(3), 1063–1070 (2008).25W. Martanto, “Microinjection into skin using microneedles,” Ph.D. disser-

tation (Georgia Institute of Technology, 2005).26A. Ramos, Electrokinetics and Electrohydrodynamics in Microsystems,

CISM Courses and Lectures Vol. 530 (Springer, New York, 2011).27E. Du and S. Manoochehri, “Enhanced ac electrothermal fluidic pumping

in microgrooved channels,” J. Appl. Phys. 104(6), 064902 (2008).28E. Du and S. Manoochehri, “Optimal design of microgrooved channels

with electrokinetic pumping for lab-on-a-chip applications,” IET

Nanobiotechnology 4(2), 40–49 (2010).29E. Du and S. Manoochehri, “Microfluidic pumping optimization in micro-

grooved channels with ac electrothermal actuations,” Appl. Phys. Lett.

96(3), 034102–034105 (2010).30J. Oh, R. Hart, R. Capurroa, and H. Nohc, “Comprehensive analysis of par-

ticle motion under non-uniform AC electric fields in a microchannel,” Lab

Chip 9(1), 62–78 (2009).31R. Zhang, C. Dalton, and G. A. Jullien, “Two-phase AC electrothermal flu-

idic pumping in a coplanar asymmetric electrode arrays,” Microfluid.

Nanofluid. 10(3), 521–529 (2011).32R. Zhang, “Experimental and modeling studies of drug delivery into tissue

using microneedle arrays,” Ph.D. dissertation (University of Calgary,

2011).33R. Perch-Nielsen, N. G. Green, and A. Wolff, “Numerical simulation of

travelling wave induced electrothermal fluid flow,” J. Phys. D: Appl. Phys.

37, 2323–2330 (2004).34P. Zhang, C. Dalton, and G. A. Jullien, “Design and fabrication of MEMS-

based microneedle arrays for medical applications,” Microsyst. Technol.

15(7), 1073–1082 (2009).35See http://www.comsol.com for Multiphysics modeling and simulation

software, Comsol, Inc., 2011.36D. F. Chen and H. Du, “Simulation studies on electrothermal fluid flow

induced in a dielectrophoretic microelectrode system,” J. Micromech.

Microeng. 16, 2411–2419 (2006).37S. Lim, H. Park, E. Choi, and J. Kim, “Long-circulating, temperature-

sensitive and EGFR-targeted liposomes for drugs delivery,” Key Eng.

Mater. 342–343, 537–540 (2007).

024701-8 Zhang, Jullien, and Dalton J. Appl. Phys. 114, 024701 (2013)