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Sensors and Actuators B 152 (2011) 267–276

Contents lists available at ScienceDirect

Sensors and Actuators B: Chemical
journal homepage: www.elsevier.com/locate/snb

A microfluidic platform with integrated flow sensing for focal chemical
stimulation of cells and tissue
Jonathan T.W. Kuo a , Li-Yuan Chang a , Po-Ying Li b , Tuan Hoang a , Ellis Meng a,b,∗abBiomedical Microsystems Laboratory, Department of Biomedical Engineering, University of Southern California, Los Angeles, CA, United States
Ming Hsieh Department of Electrical Engineering, University of Southern California, Los Angeles, CA, United Statesa r t i c l e

i n f oArticle history:
Received 8 July 2010
Received in revised form
10 December 2010
Accepted 14 December 2010
Available online 21 December 2010
Keywords:
Parylene C
Focal chemical stimulation
Neural interface
Chemical interface
Fluid deliverya b s t r a c t
A microfluidic platform for precise biochemical control of the extracellular microenvironment was developed. A chemical interface was established with cells or tissues through the precise and focal delivery ofsoluble chemical agents through a pore addressed by a polymer microchannel. Thermal flow sensors were
integrated along the length of the microchannel and monitored internal flow rate. Sensor performancewas characterized in anticipation of future studies with real-time feedback control of focal delivery. The
microfluidic system was characterized by determining the fluid delivery rates through the pores andconcentration profiles of agents delivered. Finally, focal delivery to rat retinal tissue was demonstrated.
© 2010 Elsevier B.V. All rights reserved.

1. Introduction
The cellular microenvironment contains a number of chemical gradients that fundamentally influence cell functions such ascell signaling, nutrient uptake, waste disposal, and gas exchange.
Thus, control of these microenvironments is sought to enable precise studies of cells and tissues both in vitro and in vivo. The
conventional method to control microenvironments is by releasing soluble factors to cells through bulk fluid flow delivered byhand-pipetting or other forms of perfusion and targets only large
cell populations. However, this crude method does not allow for
precise modulation of the cellular microenvironment for a targeted cell or small cell group within a population [1,2]. A focusedchemical stimulus in the targeted microenvironment of a cellwithin a population can evoke a specific response that elucidates
its relationship within the cellular network or tissue organization.
Devices that allow precise and repeatable modulation of the cellular microenvironment would benefit biological applications inwhich precise control over the microenvironment is required suchas stem cell niches, selective differentiation, metaplasia (reversible
transformation of cell from normal to abnormal state), synap-∗ Corresponding author at: Department of Biomedical Engineering, Universityof Southern California, 1042 Downey Way, DRB-140, Los Angeles, CA 90089-1111,
United States. Tel.: +1 213 740 6952; fax: +1 213 821 3897.
E-mail address: ellis.meng@usc.edu (E. Meng).
0925-4005/$ – see front matter © 2010 Elsevier B.V. All rights reserved.
doi:10.1016/j.snb.2010.12.019

togenesis, and chemotaxis [1,3–8]. Microfluidic dosing devices
that enable selective stimulation of cells have been introduced[1,2,9–12]. Devices that produce controlled laminar flows definedby microchannels have demonstrated cellular and sub-cellular
resolution of cell inactivation by changing the localized microenvironment of cells cultured in the microchannels [2,10,11]. For
example, Tourovskaia et al. demonstrated localized stimulationof muscle cells cultured within a main poly(dimethylsiloxane)
channel 2 cm long, 1500 !m wide, and 250 !m high with a hydrodynamically focused agrin stream (∼100–150 !m wide) deliveredby a central inlet channel connected to the main channel [2]. However, these devices are suitable for cell culture studies but are
difficult to use with tissue samples.
Another method of controlling the microenvironment is torelease soluble factors from a small aperture (2–8 !m in diameter)[1,2,9]. This technique can produce repeatable localized chemical
gradients at multiple locations on a device and is applicable fortissue applications. Agrin was focally applied via a microaperturearray to myotubes cultured on top of the array; apertures (ranging 2–8 !m diameter) were etched into a low-stress silicon nitride
membrane and useful in studies of neuromuscular synaptogenesis [1]. Technology for the localized delivery of soluble agents is
also beneficial not only for the study of in vitro or ex vivo systems but also for in vivo therapeutic applications such as drugdelivery and neurotransmitter-based biomimetic neural prostheses [1,2,9,12–14]. A neural prostheses intended for restoring vision
was devised in which a flow control channel and aperture wereused to control neurotransmitter release from each aperture andAuthor's personal copy
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Fig. 1. Top view of a microfluidic platform with three 50 !m wide Parylene
microchannels. Each channel is centrally perforated with a pore having a size indicated by the label on the upper right corner of each channel (5, 10, 20 !m diameter).
Eight Pt thermal flow sensors line each channel and flank the central pore.

thus achieve localized neurotransmitter stimulation of neurons in
the visual pathway [12].
Despite the advances summarized above, no microfluidic focaldelivery devices currently include integrated flow rate monitoring
for accurate determination of flow rate or volume delivered. Thus,
a novel microfluidic platform was developed that is capable of controlled fluid delivery from a “pore” to a cellular microenvironment
with both passive (diffusion) and active (ejection) modes. Each poreis addressed by its own dedicated microchannel. This microfluidic
delivery platform is the first to integrate thermal flow sensors formonitoring the flow rate of each microchannel and so feedback control of focal delivery for each pore is achievable. Arrays of thermal
flow sensors made of resistive elements have proven to be simple
to fabricate in MEMS processing, useful in many modes of operation (hot-film, calorimetric, and time-of-flight), and can be used
to determine the thermal properties of the working fluid [15–18].
The thermal flow sensors in the microfluidic platform are capableof detecting flow rates of nL/min. Fluid delivery from different sizedpores was investigated and preliminary focal delivery to tissue wasdemonstrated.
2. Design
The microfluidic platform consists of an array of three surface
micromachined Parylene C microchannels on a soda lime glass substrate (Fig. 1). The microchannels measure 6 mm × 100 !m × 4 !m(L × W × H) with 2 !m thick walls. Each microchannel is perforated
by a centrally located pore that is flanked by a linear array of platinum thermal flow sensors (Fig. 2). Pore sizes of 5, 10, or 20 !m indiameter were used. SU-8 structures provide inlet and outlet accessports (1 mm in diameter) for macro-world fluidic access from the
top of the chip and eliminate the need for through-wafer etching. A
funnel structure connects the access ports to the microchannel andits top wall is supported by Parylene posts for increased mechanicalrobustness to prevent collapse of the wide, free-standing structure.
Parylene C was chosen as the structural material of the
microchannels for its proven bioinert property as well as in vitro
cytocompatibility [17,19–23]. Since the microchannels will beplaced in direct contact with cells/tissue, biocompatibility is a priority for stable long-term studies. Parylene C is also compatible
with low temperature microfabrication processes, including thoserequired to easily form surface micromachined channels [24–26].
The transparency of Parylene C and the supporting glass substrate
allow for real-time visual inspection of flow within the microchannel, localized ejection of chemicals at the pore, and the biologicalresponse of the cells/tissue to the focal chemical gradient using
optical microscopy techniques.
Platinum thermal flow sensors made of thin film resistors werechosen for their ease of fabrication and proven flow sensing properties: high corrosion and oxidation resistance, large temperature
range, linear resistance versus temperature relationship, and hightemperature coefficient of resistivity (TCR) [17,27–32]. A Parylene C
insulation layer over the thermal flow sensors prevents undesirable
reactions with conductive solutions.
3. Theory
3.1. Thermal flow sensing
Thermal flow measurement relies on convective heat loss from aheated resistive element into the surrounding fluid flow. The resistive element or flow sensor is electrically heated above the fluid
temperature and experiences a loss of heat proportional to the fluid
thermal conductivity and experienced flow velocity. Six different
operational modes are possible with thermal flow measurementby controlling either the heating power or temperature and monitoring the heater temperature, power, or temperature difference[33]. Hot-film mode was investigated for this study for its ease of
implementation.
Sensors are located adjacent to the fluid flow and experience
convective heat loss that increases with fluid velocity in hot-filmmode. Monitoring the resistance or voltage change of the sensor
yields an indirect measure of flow rate that is first calibrated against
a known imposed flow rate. For typical thermal flow sensor materials, the resistance relationship to temperature is given by:R(T ) = R(T0 )[1 + ˛(T − T0 )]where R(T) is the resistance at temperature T and ˛ is the TCR. TCR
can be determined experimentally by:aR =Fig. 2. Close up view of one of the Parylene microchannels (100 !m wide). The Pt
thermal flow sensors and pore (20 !m diameter) can be clearly seen.

(1)

R(T ) − R(T0 )= ˛(T − T0 )R(T0 )

or

!R
= ˛!T
R

(2)

in which aR is the resistance overheat ratio and determined by measuring the change in resistance of the material at two different
temperatures. Hot-film sensors can be operated in constant current or constant temperature modes [34]. In constant current (CC)mode, the current bias imposed on the sensor is held constant andthe flow rate is measured by monitoring the change in resistance (orvoltage) due to the resulting temperature change of the sensor. In
constant temperature (CT) mode, feedback circuitry is used to holdthe temperature of the sensor constant and the increase in heating
power needed to maintain that temperature in response to convective heat loss to the adjacent fluid flow is measured. CT mode,
although more difficult to implement, can offer better resolution
and frequency response over CC mode [18].

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3.2. Diffusion modeling
Diffusion of fluid from the pore in the microchannel can be modeled with a constant point source diffusion concentration profile
given by:pu(t, x) = C1 [1 − erf(x/
∂2 u

ıu
=D 2,
ıt
∂x

4Dt)]

t > 0, x ∈ (0, ∞),(3)u(t, 0) = C1 , lim u(t, x) = 0,x→∞u(0, x) = 0,x ∈ (0, ∞)where u is concentration, D is the diffusion coefficient, x is the distance from the aperture, t is time, and C1 is the initial concentration.
4. Fabrication
Devices were fabricated on piranha (H2 SO4 :H2 O2 4:1) cleaned3(( soda lime substrates (Fig. 3). First, Pt thermal sensors were
defined using a standard liftoff process. AZ4400 (AZ Electronic
Materials, Branchburg, NJ) photoresist was applied (3 krpm, 30 s),
softbaked (90 ◦ C for 3 min), and exposed. Following development
in AZ351 (diluted 1:4 with deionized H2 O, AZ Electronic Materials,
Branchburg, NJ), a short descum step was performed (O2 plasma:
100 mTorr, 60 W, and 1 min). Ti/Pt (200 Å /2000 Å) were depositedby electron-beam evaporation and subsequently defined by liftoff
in sequential immersions with gentle mechanical agitation in acetone, isopropyl alcohol, and deionized water. Following a short
descum to clean the surface (O2 plasma: 100 mTorr, 60 W, and1 min), adhesion promoter (A-174, Specialty Coating Systems, Indianapolis, IN) was applied following the manufacturer’s standard
recipe and a Parylene C film (2 !m) was vapor deposited (Labcoter2010, Specialty Coating Systems, Indianapolis).

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Next, microchannels were fabricated using surface micromachining techniques. Sacrificial photoresist (4 !m AZ4400) was spunon (3 krpm, 30 s), exposed, and developed (diluted AZ351) to defineboth the channel pattern and small pits that eventually form support posts along the channel length to prevent collapse. A short
descum was performed to clean the surface (O2 plasma: 100 mTorr,
60 W, and 1 min) and then channels were completed by depositing
a second layer of Parylene C (4 !m). This Parylene layer was etchedto obtain fluidic access to the channel, form the central fluid delivery pore, and create contact pad openings to the thermal sensor
elements. AZ4400 (6 !m) was applied (2 krpm, 30 s), exposed, anddeveloped (diluted AZ351). Parylene was etched in oxygen plasma
in a reactive ion etcher (200 mTorr and 150 W) [20]. Any remaining photoresist was removed and the surface was descumed (O2plasma: 100 mTorr, 50 W and 30 s).
Finally, the SU-8 fluidic ports were fabricated at the inlet andoutlet to facilitate fluidic access to the Parylene channels and prevent mechanical damage to the structure during packaging [35].
SU-8 2100 (Microchem Corp., Newton, MA) was spun on (2 krpm,
30 s, 75 !m), softbaked (7 min at 65 ◦ C and 30 min at 90 ◦ C bothwith a 3 ◦ C/min ramp), exposed (360 mJ/cm2 ), post exposure baked(5 min at 65 ◦ C then 8 min at 95 ◦ C both with a 3 ◦ C/min ramp),
developed, and hardbaked (30 min at 100 ◦ C with a 3 ◦ C/min ramp).
Individual dies were diced and separated. To remove the sacrificial
photoresist in the channels, dies were immersed in isopropyl alcohol (IPA) over several days then rinsed clean in deionized water
before use. IPA was used instead of acetone to remove the sacrificial photoresist. Although the dissolution process is significantly
slower, this eliminates SU-8 cracking or delamination from theinlet/outlet ports due to large stresses resulting from exposure toacetone.
5. Packaging
A custom made acrylic jig was fabricated using a laser drill(Mini/Helix 8000, Epilog, Golden, CO) to package the microfluidic
platform to allow for simultaneous fluidic and electrical connections (Fig. 4). Sylgard 184 (Dow Corning, Midland, MI) was used
to make gaskets to create a seal around the inlet and outlet portsof the microchannel. Glass capillaries (5 !L, Alltech, Deerfield, IL)
were inserted through the top of the jig into the PDMS gaskets tocreate fluidic connections into and out of the channel.
The microfluidic platform was first bonded to a custom PCB
using standard epoxy. The electrical contact pads on the microflu-

Fig. 3. Fabrication process for creating the microfluidic platform.

Fig. 4. Packaged microfluidic platform in a custom made acrylic jig with gold wirebonds connecting the embedded Pt thermal flow sensors to the PCB. Wires were
soldered to the PCB and covered with epoxy to make robust external electrical connections. Glass capillaries aligned to and clamped on top of the SU-8 inlet/outlets
provided a fluidic connection to the microchannel.

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idic platform were then wire bonded (Au) to the PCB. Finally, wires
were soldered to the PCB for external electrical connections to theembedded flow sensors.
The microfluidic platform bonded to the PCB was then carefullyaligned between the top and bottom pieces of the acrylic jig to
make sure the position of the glass capillaries matched the inlet
and outlet ports of the device. The two pieces of the jig were thentightened together by screws, securing the device.
PEEK tubing was used to connect to the glass capillaries witha short piece of silicone tubing acting as a transition tube. The
other end of the PEEK tubing was connected to a precision glass
syringe (Gastight Syringes, Hamilton Company, Reno, NV) driven
by a syringe pump (PHD 2000, Harvard Apparatus Inc., Holliston,MA).
6. Experimental methods
6.1. Benchtop calibration of flow sensors
First, the zero power resistance of each flow sensor wasdetermined by linear extrapolation from resistances measured
at two different applied currents (2400 SourceMeter, Keithley
Instruments Inc., Cleveland, OH). Next, the linear current–voltage
relationship of the flow sensors was verified. The temperature coefficient of resistance of a flow sensor was calculated from device
resistance measurements (2700 MultiMeter, Keithley Instruments
Inc., Cleveland, OH) obtained at known temperatures applied bya programmable oven (Environmental Chamber, Sun Electronic
Systems Inc., Titusville, FL). This process was controlled and automated with a custom LabVIEW program. The ratio of the change inresistance to nominal resistance, or resistance overheat ratio (aR ),was calculated and used to determine the temperature rise of theresistive sensor, or overheat temperature (OHT), from the appliedcurrent.
6.2. Flow sensor calibration in hot-film mode
The hot-film method tracks fluid flow by measuring convective
heat loss from the heated flow sensor to the fluid. Two different
hot-film modes (constant current and constant temperature) were
investigated with devices that had the pore plugged by a drop of
epoxy as shown in Fig. 5. The packaged device was placed in an ovenfor thermal isolation to minimize interference from fluctuations

Fig. 5. Illustration of thermal flow sensing principle under constant current hot-filmmode. The pore is plugged and forced convection of heat by fluid flow is detectedby the change in voltage in the resistive sensing element.

in the ambient temperature. A small opening in the oven allowed
for fluidic and electric access to the device. Deionized (DI) water
was chosen to demonstrate relevant sensor performance since thethermal conductivity of chemical solutions to be delivered in ourapplication are close to that of water.
6.2.1. Constant current
Sensor calibration was performed at three different overheat
ratios corresponding to overheat temperatures of 7.4 ◦ C, 10.1 ◦ C,
and 13.3 ◦ C. A single flow sensor was connected in a Wheatstonebridge under quarter-bridge configuration. The 2400 SourceMeter provided a constant current source and the 2700 MultiMeter
recorded the voltage output of the bridge under the control of a custom LabVIEW program. Filtered DI water was pumped through themicrochannel at predetermined flow rates using a syringe pump.
Data for each flow rate was acquired after sensor response stabilized and then averaged (n = 83, 0.5 samples per second).
6.2.2. Constant temperature
Sensor calibration under constant temperature operation was
performed at three different overheat ratios (the same ones usedin constant current mode). Constant overheat temperature wasmaintained by using a feedback circuit as previously described[17]. Briefly, the output of a Wheatstone bridge containing the sensor was fed into a negative feedback operational amplifier circuit.
Imbalance in the Wheatstone bridge due to cooling of the flow
sensor was automatically corrected by the circuit with increased
current in order to compensate for resistance and thus temperaturechange. The output voltage was further amplified to facilitate themeasurement. Data was acquired using a custom LabVIEW programinterfaced to the 2700 MultiMeter while a syringe pump controlled
the flow rate in the microchannels. The output voltage for each flowrate setting was recorded and averaged (n = 60).
6.3. Thermal flow sensing of fluid delivery through pore
A microchannel was investigated for electrical sensing of fluiddelivery through a pore with a diameter of 20 !m. Two flow sensors,
one before the pore and one after, were biased at 4 mA in hotfilm constant current mode using two separate 2400 SourceMeters
as a CC source for each (sensors equidistant from pore, distance
between sensors = 4.5 mm). Both sensors were placed in two separate and identical Wheatstone quarter-bridges (Fig. 6). The bridge
outputs were simultaneously recorded (n = 58) for each flow ratesetting (filtered DI water) using two separate channels of the
LabView controlled 2700 MultiMeter (Keithley Instruments Inc.,
Cleveland, OH). The data recorded at each flow rate were thenaveraged, and the resulting curves were compared to one another.

Fig. 6. Wheatstone quarter-bridge schematic used for thermal flow sensing of fluid
through pore. Two separate and identical bridges were used: one for the sensor
before the pore and one after.

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6.4. Optical monitoring of fluid delivery through pore
6.4.1. Fluid delivery rate in air
Fluid delivery rates generated through a 10 !m pore was investigated. Rhodamine B (0.11 mM) was used as a dye indicator andflowed through the device. This setup was placed underneath a
microscope (Motic, Xiamen, China) in order to image the dye bolus
ejected from the pore. Different flow rates were applied to themicrochannel via syringe pump (50–750 nL/min). The dye bolus
was estimated to be a hemisphere; the radius was analyzed using
NIH ImageJ v1.34 to calculate the delivered volume and hence theflow rate out of the pore. The evaporation rate was measured andthe data were corrected to account for evaporative fluid loss overthe course of the experiments.
6.4.2. Fluid delivery rate in aqueous solution
Fluid delivery rate into aqueous was performed in two modes:
diffusion and active ejection. The microchannel was first primed
with Rhodamine B. Next, filtered DI water was added to the SU8 reservoir immersing the pore (10 !m diameter). Fluid deliveryof Rhodamine B by diffusion through the pore into the water
was time-lapse imaged by a fluorescent microscope (Nikon Eclipse
LV100, Tokyo, Japan). The optical intensity of the dye was analyzed
using NIH ImageJ and normalized for different distances from theaperture or pore. This data was then compared to the predicted
results from point source diffusion theory with an initial concentration C1 set to 1 and the diffusion coefficient D set to 2.8 × 10−6 cm2 /s[36].
Fluid delivery of Rhodamine B by active ejection through thepore into the water was conducted for the three different poresizes at three different input flow rates to the microchannel(50, 100, 200 nL/min) imposed by a syringe pump. A separate
set of experiments were also performed in which a glass coverslip was placed on top of the final SU-8 layer applied to thedevice (75 !m above the pore). In the former case, volume andthus flow rate were estimated by assuming a hemispherical Rhodamine B plume, and for the latter, the volume was assumed to beapproximately cylindrical (limited to 75 !m in height by the glass
coverslip).
6.5. Tissue study with diffusion and ejection
In vitro focal delivery from the pore to dissected wildtype
Sprague–Dawley rat retinal tissue was performed. The retinal tissue
was placed over the microfluidic device and gently held flat against
the pore using a ring of Whatman filter paper in combination witha metallic ring clamp (Fig. 7). Rhodamine B was delivered to thetissue through the pore (20 !m). Time-lapsed photography (0.25 sintervals) under a fluorescence microscope was used to record diffusion and ejection of Rhodamine B from the pore into the tissue.
Diffusion-based delivery was first performed over a period of 25 s.
Ejection mode (dye from the pore ejects directly into air as opposed
to water) was then initiated by priming the device for 1 minuteand then delivery was recorded over a 25 s period on the same
tissue. The images were then processed using NIH ImageJ by subtracting the image at the start when no focal delivery had takenplace from subsequent images in order to show the difference due
to focal delivery of dye. The tissue was then removed for visualinspection.

Fig. 7. (a) Experimental test setup for Rhodamine B focal delivery to rat retina and(b) zoomed in brightfield image of retinal tissue mounted over a microchannel andpore.
Reprinted from Chang, L.Y. et al., Integrated Flow Sensing for Focal Biochemical
Stimulation, NEMS 2008, © 2008 IEEE.

bulk Pt but consistent with that of e-beam evaporated Pt thin films[30]. The zero power resistance, IV curve and OHT relationships fora representative sensor are shown in Figs. 9–11, respectively.
The OHT relationship was used to determine the maximum bias
current (4 mA) to keep the flow sensor’s overheat temperature
below 15 ◦ C. This is to ensure biocompatibility for application with

7. Results and discussion
7.1. Benchtop calibration of flow sensors
A representative TCR curve is displayed in Fig. 8. This sensor
exhibits a TCR of 1.9 × 10−3 / ◦ C, which is much less than that ofFig. 8. Typical Pt flow sensor TCR curve. This sensor exhibits a TCR of 1.9 × 10−3 /◦ C.

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Fig. 9. Representative Pt thermal flow sensor zero power resistance result. For thissensor, the zero power resistance is 217.2 #.

cells and tissues and avoid denaturing soluble factors that might bedelivered by the microchannel.
7.2. Flow sensor calibration in hot-film mode
7.2.1. Constant current
A device with the pore sealed was used to calibrate the flow
sensors under different applied flow rates at three distinct overheat ratios in constant current injection mode. The resulting voltagechange in response to flow rates ranging from 0–1 !L/min wererecorded and averaged (Fig. 12). Kruskal–Wallis non-parametricstatistical analysis was performed on the sensor responses and astatistically significant difference (p < 0.001) was found between
the data taken at different flow rates. At equilibrium, measured
responses were stable and thus small changes in flow rate weredetectable and discernable despite the low voltage output. Altogether, this suggests a sensor resolution of at least 250 nL/min;lower flow rates were not tested here. Two sensitivity regimes wereobserved in the flow rate range likely due to thermal time constants
associated with the packaging materials. As expected, the CC results

Fig. 10. Representative IV curve for a typical sensor. A linear fit was applied to thedata (dots).

Fig. 11. Pt thermal flow sensor overheat behavior. The sensor resistance as a function of applied current was obtained and the overheat temperature was calculated
from the empirically determined TCR for the sensor element. In this case, a maximum
biasing current of 4 mA corresponding to ∼15 ◦ C increase in sensor temperature wasapplied to all experiments to maintain biocompatibility.

showed greater response to flow rates with higher constant current
injection.
7.2.2. Constant temperature
The flow sensors were calibrated under CT mode again with
controlled flow rates from 0–1 !L/min at different overheat ratios(Fig. 13). CT mode provided improved flow measurement resolution and greater signal amplitude due to amplification of the voltage
output from the feedback circuit. Once again, higher currents corresponding to higher overheat ratios exhibited greater response tochanges in flow rate. Despite the increase in complexity of the readout circuitry, constant temperature mode is further preferred for its
increased frequency response over constant current operation [18].
7.3. Thermal flow sensing of fluid delivery through pore
A device with an unobstructed pore was used to measure theflow delivered through the pore (DI water) by monitoring the out-

Fig. 12. Hot-film responses for three different overheat ratios in constant current
biasing (mean ± S.E. with n = 83). Statistical analysis of the data indicates a significant change in sensor response between all flow rate measurements, so the sensor
resolution is at least 250 nL/min.

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Fig. 13. Hot film responses acquired at three different overheat ratios in constanttemperature biasing mode (mean ± S.E. with n = 60) with amplification of output
signal by the feedback circuit.

put of two flow sensors, one located upstream of and the otherdownstream of the pore. Again, known flow rates from 0–1 !L/min
were used and the responses of two sensors were measured andrecorded in CC mode (Fig. 14). Fluid delivery through the pore wasconfirmed by optical inspection under a microscope. For higher
flow rates, the upstream sensor exhibited a greater voltage output than the downstream sensor indicating that fluid delivery
through the pore can be measured by using this differential measurement technique. However, increased resolution is required
for nL/min flows and may be achieved by incorporating thermal isolation structures underneath the sensor regions to prevent
undesirable heat transfer to the glass substrate. Statistical analysis(Kruskal–Wallis non-parametric test) was performed on the dataand statistically significant differences (p < 0.01) were found for all
but three points (400 and 1000 nL/min upstream of the pore and
400 nL/min downstream of the pore). One contributing factor is
likely the low signal-to-noise ratio.

Fig. 14. Hot film response for 4 mA constant current biasing for two flow sensors,
one upstream and one downstream of the pore in the channel (mean ± S.E. withn = 58). Both sensors were monitored simultaneously. Data points marked with anasterisk have non-significant changes due to change in flow rate based on statisticalanalysis. This is attributed to low signal-to-noise ratio.

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Fig. 15. Fluid delivery into air with and without evaporation compensation from a
10 !m pore. A linear fit was applied to the uncompensated data.
Reprinted from Chang, L.Y. et al., Integrated Flow Sensing for Focal Biochemical
Stimulation, NEMS 2008, © 2008 IEEE.

7.4. Optical monitoring of fluid delivery through pore
7.4.1. Fluid delivery rate in air
The flow rate of fluid ejected from the pore was investigated byflowing dye through the microchannel with different set flow rates
from 50 to 750 nL/min. The raw data and corrected data to accountfor evaporative loss are shown in Fig. 15. As expected, the rate of
fluid delivered from the pore has a positive correlation with theflow rate in the microchannel. Fluid delivered from the 10 !m pore
ranged from 16 to 106 nL/min. As applied flow rate increased, theproportion of the overall flow delivered through the pore decreased
although the fluid delivery rate through the pore increased. The
data suggest that localized delivery of pL/nL volumes immediately
adjacent to the pore is possible using such an approach.
7.4.2. Fluid delivery rate in aqueous solution
Diffusion from a 10 !m sized pore was investigated and compared to theoretical diffusion in a constant point source model(Fig. 16). The measured diffusion optical intensity profiles (onedimension) correlate well to those predicted by the model. Thus,

Fig. 16. Fluid delivery of Rhodamine B by diffusion (dots) into DI water comparedto point source diffusion theory (solid lines) from a 10 !m pore.
Reprinted from Chang, L.Y. et al., Integrated Flow Sensing for Focal Biochemical
Stimulation, NEMS 2008, © 2008 IEEE.

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cultured directly on the device will mimic the first case in which
there is no glass coverslip. Cells will also be culture on a coverslip
and inverted over the pore to mimic the latter case with a glass
coverslip.
7.5. Tissue study with diffusion and ejection
Delivery of Rhodamine B from the pore to retinal tissue wasdemonstrated in two fluid delivery modes: diffusion and ejection.
Time-lapse fluorescence microscopy images were captured in bothmodes showing dye accumulation over time in the tissue (Fig. 18).
The rate of accumulation of dye within the tissue in ejection modeis visibly more pronounced than in diffusion mode. This is expected
since diffusion is mass transport limited which effectively reduces
the dye delivered to the tissue compared to ejection.
8. Conclusion

Fig. 17. Flow rates from different sized pores at different flows in the microchannel(mean ± S.D. with n ≥ 3).

this model can then be used to predict and plan concentration
profiles of released soluble factors. Experimentally, it was determined that the diffusion plume exhibited an average diffusion rateof 4.9 !m/s from the 10 !m pore (data not shown).
Data for fluid delivery by syringe pump-driven ejection withand without a flow limiting glass coverslip were obtained (Fig. 17).
The influence of pore size on the flow rate is clearly observed at
200 nL/min but less obvious at lower flow rates when no glass coverslip is placed over the pore. The presence of the glass coverslip
reduces the flow rate compared to the case without in all instances.
Some evaporative fluid loss may occur without the coverslip, especially since the data is acquired under magnification in a compound
optical microscope. The presence of the glass coverslip eliminates
fluid loss but also limits fluid delivery. In future studies with cells
cultured near the pore, both situations will be investigated. Cells

A microfluidic platform for precise focal delivery of soluble
factors to cells and tissues was demonstrated. An array of three
Parylene microchannels each centrally perforated with a uniquely
sized small diameter pore were investigated. This arrangement
allows a one-to-one correlation between the flow delivery channel
and each pore. Each channel possesses Pt thermal flow sensors for
integrated flow sensing. Focal fluid delivery capabilities of the system were characterized by calibration and operation of the thermal
flow sensors, investigation of delivery in different modes throughdifferent sized pores, and preliminary demonstration of focal fluiddelivery to tissue. The thermal flow sensors operated under steadystate conditions were able to detect nL/min flow rates.
Future work entails feedback control of fluid flow with the flow
sensors for precise delivery of soluble factors through the pore.
Future work also includes culturing of cells on top of the microfluidic platform for in vitro studies of cells with precise nL/min flowrate focal delivery of soluble factors. High-density arrays of individually controlled pores will allow for more sophisticated interfaces
to tissues and cells and greatly enhance the ability of researchers
to address complex biological questions.

Fig. 18. (a) Fluorescence time-lapse images of focal delivery of Rhodamine B by diffusion to rat retina and (b) ejection mode time-lapse images focal delivery of Rhodamine
B on the same tissue after diffusion. The arrow indicates the vicinity above the 20 !m pore. The original image at time = 0 s was subtracted from the subsequent images after
focal delivery start for both diffusion and ejection to highlight the difference in dye accumulation, hence the “donut” shape for the dye plume in (b).

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Acknowledgement
This work was funded in part by an NSF CAREER grant underAward Number EEC-0547544. The authors would like to thank Mr.
Merrill Roragen and Dr. Donghai Zhu for fabrication assistance,
Dr. David Merwine for providing the retinal tissue, and membersof the Biomedical Microsystems Lab at the University of Southern California for their contributions to this work (especially Dr.
Gabriela Mallén-Ornelas, Mr. Lingyun Zhao, Mr. Michael Wang, andMr. Lawrence Yu).
Appendix A. Supplementary data
Supplementary data associated with this article can be found, inthe online version, at doi:10.1016/j.snb.2010.12.019.
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Biographies
Jonathan T.W. Kuo received the B.S. degree in bioengineering from the Universityof California, San Diego, in 2007. He is currently working towards the Ph.D. degree
in biomedical engineering at the University of Southern California, Los Angeles. His
research interest lies in using bioMEMS to create biosensors.
Li-Yuan Chang received the B.S. degree in biomedical engineering from the University of California, San Diego, in 2003, and the M.S. degree in Medical Device
and Diagnositic Engineering from the University of Southern California, Los Angeles, in 2007. Since 2008, he has joined Citibank Commerical Banking Division in
Taiwan specializing in company valuations and loan originations for small medium
enterprises in pharmaceutical and medical space.
Po-Ying (Brian) Li received the B.S. and M.S. degree in mechanical engineering fromthe Tatung Institute of Technology, Taipei, Taiwan, in 1996 and the National Tsing
Hua University, Hsinchu, Taiwan, in 2001. He also received a M.S. degree in materials
science in 2004 and a Ph.D. degree in electrical engineering in 2009 from the University of Southern California, Los Angeles. Currently, he is a Micro Device Engineer
at the Replenish, Inc. He was a research assistant member with the National Science Foundation Biomimetic MicroElectronic Systems Engineering Research Center(BMES ERC) in the University of Southern California from 2004 to 2009. His research
focused on the fabrication and characterization of the first implantable bioMEMS
intraocular drug delivery device for the treatment of glaucoma and age-related
macular degeneration. Dr. Li is a member of Eta Kappa Nu (HKN), Sigma Xi, Institute of Electrical and Electronics Engineers (IEEE), American Society of MechanicalEngineers (ASME), American Chemical Society (ACS), and American Vacuum Society(AVS).
Tuan Hoang received the B.S. degree in biological sciences from the California
Institute of Technology, Pasadena in 1995. He was a graduate student in the Interdepartmental Ph.D. Program for Neuroscience at UCLA from 1997 to 2004. Since 2006,
he is a faculty member in the Department of Biomedical Engineering, University ofSouthern California, Los Angeles. His research interest is to apply neuroengineering
principles for the study of neurodegeneration and neurorepair. He was a recipient of the Caltech-Ahmanson Foundation scholarship from 1992 to 1995 and a NIH
predoctoral training grant for molecular and cellular biology from 1999 to 2001.
He was a member of the International Society of Magnetic Resonance in Medicine
from 1996 to 1998 and a member of the Society of Neuroscience from 1998 to
2004.
Ellis Meng received the B.S. degree in engineering and applied science and the M.S.
and Ph.D. degrees in electrical engineering from the California Institute of Technology (Caltech), Pasadena, in 1997, 1998, and 2003, respectively. She is an associate
professor in the Department of Biomedical Engineering, University of Southern California, Los Angeles, where she has been since 2004. She currently holds a joint
appointment in the Ming Hsieh Department of Electrical Engineering. Her researchinterests include bioMEMS, implantable biomedical microdevices, microfluidics,

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multimodality integrated microsystems, and packaging. She was the Thrust Leader
for Interface Technology and the Associate Director of Education and Student Diversity Researcher at the National Science Foundation’s Biomimetic MicroElectronic
Systems Engineering Research Center. She held the Viterbi Early Career Chair in theViterbi School of Engineering. Dr. Meng is a member of Tau Beta Pi, the Biomedical Engineering Society, the Institute of Electrical and Electronics Engineers, theAmerican Society of Mechanical Engineers, the Society of Women Engineers, and theAmerican Society for Engineering Education. She was a recipient of the Intel Womenin Science and Engineering Scholarship, the Caltech Alumni Association Donald S.
Clark Award, and the Caltech Special Institute Fellowship. She has also received the
NSF CAREER and Wallace H. Coulter Foundation Early Career Translational ResearchAwards.