Available online at www.sciencedirect.com

Sensors and Actuators A 143 (2008) 41–48

An electrochemical intraocular drug delivery device
Po-Ying Li a , Jason Shih b , Ronalee Lo a , Saloomeh Saati c , Rajat Agrawal c ,Mark S. Humayun c , Yu-Chong Tai b , Ellis Meng a,∗a Biomedical Microsystem Laboratory, University of Southern California, USA
b Caltech Micromachining Laboratory, California Institute of Technology, USAc Keck School of Medicine, University of Southern California, USA

Received 31 March 2007; received in revised form 22 June 2007; accepted 26 June 2007
Available online 4 July 2007

Abstract
A microelectromechanical systems (MEMS) drug delivery device is investigated for the treatment of incurable ocular diseases. Unlike conventional ocular drug delivery devices, this MEMS device is capable of being refilled, features electronic control of the drug regimen, and enables
targeted intraocular drug delivery. The refillable design permits long-term drug therapy and avoids repetitive surgeries. Electronic control of dosing
is achieved by using electrolysis-actuated pumping to deliver pharmaceuticals directly to the intraocular space. A flexible Parylene transscleral
cannula allows targeted delivery to tissues in both the anterior and posterior segments of the eye. This electrochemically driven drug delivery device
was demonstrated to provide flow rates suitable for ocular drug therapy (pL/min to ␮L/min). Both continuous and bolus drug delivery modes were
performed to achieve accurate delivery of a target volume of 250 nL. An encapsulation packaging technique was developed for acute surgicalstudies and preliminary ex vivo drug delivery experiments in porcine eyes were performed.
© 2007 Elsevier B.V. All rights reserved.
Keywords: Drug delivery device; Electrolysis pump; Refillable reservoir; Intraocular

1. Introduction
Intraocular diseases such as retinitis pigmentosa, age-related
macular degeneration, diabetic retinopathy, and glaucoma arepresently incurable and affect millions worldwide [1]. Diseaseprogression and lack of treatment may lead to blindness. Drugtherapy is currently the most effective treatment, however, conventional administration methods are limited in efficacy. To
successfully treat these diseases, pharmaceuticals must penetrate the protective physiological barriers of the eye such as thecornea, sclera, and the blood–retina barrier and target difficultto-reach intraocular tissues such as the ciliary body, retina, andangle [2]. The goal of precise and targeted delivery to thesetissues remains elusive with modern treatments.
Conventional ocular drug delivery therapies include oral
drugs, eye drops, and intraocular injections. Oral and topicalmethods require large overdoses to reach therapeutic levels in
the intraocular space. It has been shown that less than 5% of the∗ Corresponding author.

E-mail address: ellismen@usc.edu (E. Meng).
0924-4247/$ – see front matter © 2007 Elsevier B.V. All rights reserved.
doi:10.1016/j.sna.2007.06.034

applied medication is able to penetrate physiological barriers [3].
Severe systemic side effects may also result. To overcome physical barriers to ocular drug delivery, periocular or intraocularapproaches are necessary. Intraocular injections directly deliver
pharmaceuticals to intraocular tissues but require repeated injections (as frequent as one to three per week) for treating chronic
diseases and may contribute to dose-related side effects [4].
Many of these methods exhibit low patient compliance.
Drug delivery devices provide advanced ocular drug
therapy. Sustained-release implants (biodegradable or nonbiodegradable) allow continuous release of drug over extended
periods of time [4,5]. These devices contain a fixed charge ofdrug specially formulated for release at a fixed rate; after implantation, it is not possible to change the drug, its concentration, or
its release rate. In addition, these devices cannot be refilled. Oncethe drug is expended, the implant must be surgically removed
and replaced. Sustained-release implants are limited by safetyconcerns and lack of precision in the erosion rate [6].
Implantable pump systems have not been present historicallyin ocular drug delivery due to large size and mediocre performance [5]. However, with MEMS technology, miniaturized
devices having precise delivery in either bolus or continuous

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mode are possible. The advantages of MEMS fabrication for
producing miniaturized and efficient drug delivery systems havealready been realized for insulin delivery and delivery of bioactive compounds to neural tissue [7–9]. Due to the limitations of
traditional methods of treatment for ocular diseases, there is aneed for MEMS ocular drug delivery systems that are capableof targeted delivery to interior tissues, refillable for long-term
use, and automated to address patient compliance.
We previously demonstrated the feasibility of MEMS for
implantable intraocular drug delivery in a passive, or unpowered, device [10]. This device contained a drug reservoir
attached to a flexible check-valved cannula that was directed
through the eye wall by means of a small surgical incision. Thisconfiguration allowed drug to be delivered directly into the eye
and reach intraocular tissues in the vicinity of the cannula outlet.
To deliver drug into the eye, the device was actuated by manuallydepressing the drug reservoir. This action generated an overpressure in the reservoir which in turn caused a check valve in thecannula to open and allow drug to enter the intraocular space.
The reservoir was emptied over time by repeated dosing by manual actuation. Upon depletion of the reservoir, the drug reservoir
was refilled by puncturing the reservoir wall with a syringe needle and emptying the syringe into the reservoir. This was the firstrefillable ocular drug delivery device which extends device lifetime in the patient and provides long-term therapy. To achieve
variable delivery rates and the ability to select either bolus orcontinuous delivery, an active device having electrochemicallydriven drug delivery was investigated [11].

2. Theory
The electrolysis of water results in the phase transformation of liquid to gas. This reaction is utilized here to replace
the manual actuation in the passive device. The electrochemical
production of hydrogen and oxygen gases generates sufficientpressure within the drug reservoir to deliver drug into the eye(Fig. 1). Electrolysis actuation has attracted attention among
other forms of microactuation for its low power consumption,
simple construction, and ability to generate large displacements[12]. Electrolysis actuation as an effective means for pumping
fluids is well established in microfluidics [13–15].
The electrolysis of water can be summarized by the following
reactions:−cathode : 2H2 O(l) ↔ O2(g) + 4H+(aq) + 4e−
anode : 4H+(aq) + 4e ↔ 2H2(g)(1)

net : 2H2 O(l) ↔ O2(g) + 2H2(g)The net result is the production of oxygen and hydrogengas which contributes to a volume expansion of about a thousand times greater than the starting volume of water. This gas
evolution process proceeds even in a pressurized environment;
it has been reported that the maximum pressure at which thereaction can occur is 200 MPa [16]. Electrolysis requires fourelectrons in order to convert two water molecules into three gas
molecules (two hydrogen and one oxygen). To drive gas generation and thus pumping, current control is the preferred biasing
scheme as there is a direct correlation of the applied current to
the generated gas volume and the fluid flow rate. The theoretical pump rate (qtheoretical in m3 /s) at atmospheric pressure forcurrent-controlled electrolysis pumping is given by
qtheoretical =

3 i
Vm
4F

(2)

where i is current (in A), F the Faraday’s constant(96.49 × 10−3 C/mol), and Vm is the molar gas volume at 25 ◦ C
and atmospheric pressure (24.7 × 10−3 m3 /mol). The theoretical generated or dosed gas volume (Vtheoretical in m3 ) can be
determined by
Vtheoretical = qtheoretical t(3)where t is duration (in s) for which the current is applied.
The efficiency (␩) of an electrolysis actuator as a pump canbe defined as [15]η=

Vexperimental
Vtheoretical(4)

where Vexperimental is the actual volume of the generatedhydrogen and oxygen gases. Efficiency in electrochemical systems is a complex function of several parameters that havebeen categorized by Bard and Faulkner. The variables include
electrode design (material, surface area, geometry, and surface conditions), mass transfer conditions (transport mode,surface concentration, and adsorption), external parameters(temperature, pressure, and time), electrolysis solution (bulkconcentration of electroactive species, concentration of otherspecies, and solvent), and electrical parameters (potential, current, and quantity of electricity) [17].
3. Design

Fig. 1. Cross-section of the ocular drug delivery device depicting electrochemical pumping of drug into the eye.

The drug delivery device consists of an electrolysis pump,drug reservoir, and transscleral cannual (Fig. 2). The electrolysis pump consists of two interdigitated platinum electrodes
immersed in an electrolyte. This electrode geometry improves
pumping efficiency by reducing the current path through thesolution which also serves to lower the heat generation [18].
When current or voltage is applied, with the drug as an electrolyte, electrolysis of water in the drug at the electrodes
produces oxygen and hydrogen gases. The gases generated result

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43

Fig. 2. Schematic diagram illustrating the components of the drug delivery
device. The arrows follow the drug flow path from the reservoir through the
Parylene cannula.

in an internal pressure increase in the sealed reservoir whichcauses drug to be delivered through the cannula and into theeye. Electrolysis is a reversible process and ceases when theapplied bias is turned off. This allows the gradual recombinationof hydrogen and oxygen to water.
Drug is stored in a reservoir integrated on top of the electrolysis pump. This reservoir was casted in silicone rubber (Sylgard
184, Dow Corning, Midland, MI) and assembled on top of thepump chip by using an encapsulation technique. Previously,
we demonstrated that silicone rubber is suitable as a refillable
reservoir material; thin silicone membranes were able to sustain
multiple punctures from a non-coring syringe needle required
to refill the reservoir. Silicone rubber membranes were able toreseal repeatedly without leakage after removal of the needle
and remained leak-free even after the application of differential
pressure across the membrane [10].
A flexible transscleral cannula is connected to the pumpchamber through a small port. This layout traps generated gases
inside the reservoir while allowing pumped drug to escape into
the eye (Figs. 2–3). This cannula was surgically inserted througha small incision in the eye wall and was directed into either theanterior or posterior segment of the eye. Targeted drug delivery
can be achieved by placing the cannula outlet in the vicinityof the desired treatment site. Parylene C (Specialty Coating
Systems Inc., Indianapolis, IN) was selected as the cannula material for its mechanical strength, biocompatibility, and ease of
integration. It is a USP Class VI material suitable for the construction of implants and is well established as a MEMS material[19].
The drug delivery system is assembled from two separately
fabricated components: (1) the pump chip with an integrated cannula and (2) the reservoir. The pump/cannula chip was fabricatedusing conventional silicon micromachining and the reservoir by
the casting of silicone rubber against a master mold.

Fig. 3. Conceptual illustration of the implanted ocular drug delivery device.
The device is implanted under conjunctiva with the cannula is directed througha limbal incision into the anterior segment of eye.

(Fig. 5A). LOR 3B (MicroChem Corp., Newton, MA) was spunon at 3 krpm followed by AZ 1518 (AZ Electronic Materials, Branchburg, NJ) at 3 krpm. Ti/Pt (200/2000 Å) was e-beam
evaporated and patterned by lift-off in ST-22 photoresist stripper (ATMI, Danbury, CT) to define the interdigitated electrodesof the electrolysis pump (Fig. 5B). A second lithography step
was performed (AZ 1518 at 3 krpm) to define the cannula footprint. The oxide layer was etched using buffered hydrofluoric
acid to expose the silicon below. The photoresist was stripped
then the exposed silicon was roughened by two cycles of XeF2
etching (Fig. 5C). The first sacrificial photoresist layer (AZ 4620
spun at 2.75 krpm and hard baked to yield a 5 ␮m thick layer)
was applied to facilitate release of the cannula from the substrate (Fig. 5D). The first Parylene C layer (7.5 ␮m) forming
the bottom of the cannula was deposited followed by thermal
evaporation of 2000 Å thick Cr etch mask. Following lithogra-

4. Fabrication
The fabrication process of the pump and cannula chip (Fig. 4)
starts with a thermally oxidized silicon substrate (5000 Å)Fig. 4. Fabricated pump and cannula chip (1 mm wide cannula with an embedded microchannel measuring 100 ␮m × 25 ␮m × 5 mm).

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were released manually by gently lifting them off the substrate.
Finally, individual dies were separated and the remaining siliconbeneath each cannula was removed by scribing and breaking it
off (Fig. 5L).
4.1. Device assembly and surgical packaging
Additional packaging and assembly were required for implementation of the device in surgical experiments. The pumpchip containing the electrolysis actuator and cannula was combined with the drug reservoir and electrical wiring for supplying
current to the electrolysis electrodes. This assembly is shownin Fig. 6. First, electrical wires were bonded to the electrolysis electrode contact pads with conductive epoxy (Ohmex-AG,
Transene Company Inc., Danvers, MA). The epoxy was cured
at 150 ◦ C for 15 h under vacuum. The pump chip and drug
reservoir were then assembled by using an encapsulation technique based on silicone soft lithography [20,21] (Fig. 7). A
silicone spacer (Sylgard 184, Dow Corning, Midland, MI) was
casted against a stainless steel sphere (17.5 mm in diameter)
to adapt the planar device to fit comfortably against the curvature of the eye (Fig. 7A). Silicone rubber was mixed using a
10:1 base-to-curing agent ratio (AR-250 Hybrid Mixer, Thinky
Corp., Tokyo, Japan) and partially cured (65 ◦ C for 20 min).
The sphere mold was removed and the resulting crater was
filled with wax (Parowax, Roswell, GA) (Fig. 7B and C). A
drug reservoir was prepared by casting silicone rubber against a
conventionally machined acrylic mold (partially cured at 65 ◦ C
for 20 min) (Fig. 7D and E). The reservoir has internal dimensions of 6 mm × 6 mm × 1.5 mm which provides a maximumof 54 ␮L of drug storage. The silicone reservoir was aligned

Fig. 5. Simplified fabrication process for the pump/cannula chip.

phy (AZ 4620 at 500 rpm), the Cr was etched in CR-7 (Cyantek,
Fremont, CA) then the photoresist was stripped. The Parylene
layer was then patterned in an oxygen plasma and the Cr etchmask was removed (CR-7) (Fig. 5E). A second photoresist sacrificial layer was deposited (AZ 4620 spun at 450 rpm andhard baked to yield a 25 ␮m thick layer) to define the channel height (Fig. 5F). A second Parylene layer (7.5 ␮m) was
deposited to complete the cannula (Fig. 5G). To define the cannula from the Parylene/photoresist/Parylene sandwich, Ti/Au(200/2000 Å) was deposited as an etch mask. The etch mask
was patterned (AZ 4620 spun at 425 rpm) and etched first with
Au Etchant TFA (Transene Company Inc., Danvers, MA) andthen 10% HF (Fig. 5H). Finally, the sandwich was etched in oxygen plasma and the masking layer was stripped (Au Etchant TFA
and 10% HF) (Fig. 5I). Following the etch, the entire wafer wascleaned in 5% HF dip and by exposure to oxygen plasma. SU8 50 (MicroChem Corp., Newton, MA) was spun at 2200 rpm
resulting in a 70 ␮m thick layer after post-baking (Fig. 5J). The
sacrificial photoresist was removed by dissolving in a 40 ◦ Cacetone solution for 1 day (Fig. 5K). The individual cannulaeFig. 6. Pump and cannula chip (upper left), silicone drug reservoir (lower left),and packaged drug delivery device with attached wires for electrolysis actuation(right).

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45

Fig. 8. Schematic diagram of the flow rate testing apparatus.

Fig. 7. Fabrication process detailing steps for device packaging.

to the pump/cannula chip and silicone spacer (Fig. 7F). Thenthe Parylene cannula was immersed in deionized water which
serves as a mask to prevent coating by silicone rubber during
the encapsulation step (Fig. 7G). This assembly process exploits
the hydrophobicity of silicone rubber. The stack was immersed
in silicone prepolymer and cured at room temperature for 24 h(Fig. 7H). Finally, extraneous silicone material was removed
from the device to complete the assembly process (Fig. 7I andK).
5. Experimental methods
5.1. Bench-top testing
To investigate the performance of the electrolysis pump,experiments examining continuous delivery, bolus delivery,
pump efficiency, gas recombination, and backpressure wereconducted. For these tests, a custom testing apparatus was lasermachined (Mini/Helix 8000, Epilog, Golden, CO) in acrylic.
The experimental setup consisted of a computer-controlled CCD
camera (PL-A662, PixeLINK, Ottawa, Canada) for collecting
flow data from a calibrated micro-pipette (Accu-Fill 90, Becton,
Dickinson and Company, NJ) connected to the output port of thetest fixture (Fig. 8). Testing was performed using dyed deionizedwater as the electrolyte. The electrolysis was sustained underconstant current conditions (50 ␮A to 1.25 mA) for continuous
delivery operation. The pump efficiency associated with theseresults was calculated from (4). The relationship between efficiency and recombination of hydrogen and oxygen to water wasdetermined.

Bolus delivery was also examined. A constant current pulse(0.5, 1.0, and 1.5 mA) was applied for 1, 2, and 3 s. Repeatedtrials were performed (n = 4) to obtain the average dosed volume. Normal intraocular pressure (IOP) ranges from 5 to
22 mmHg (15.5 ± 2.6 mmHg (mean ± S.D.)) [22]. Values outside this range correspond to abnormal intraocular pressure
which is a characteristic of glaucoma (>22 mmHg). Thus, it isnecessary to characterize pump performance under these physiologically relevant conditions. The experimental setup in Fig. 8
was modified for this analysis by connecting a water column tothe outlet of the micro-pipette. Backpressure was applied to thedrug delivery device by adjusting the height of the water column.
Data was collected for backpressures corresponding to normal
IOP (20 mmHg) and abnormal IOP (0 and 70 mmHg).
5.2. Surgical testing
The prototype drug delivery devices were implanted in enucleated porcine eyes. Preliminary ex vivo surgical modelingin enucleated porcine eyes is necessary to prepare for device
demonstration in vivo. The operation of each surgical device
was tested prior to the surgical experiment to check for clogs in
the cannula and to verify the integrity of the electrical connections. The drug reservoir was filled with dyed deionized waterthen the reservoirs were manually depressed which generates
sufficient pressure to expel the fluid from the reservoir. A second test was conducted to verify operation of the electrolysis
pump by connecting to an external power supply and driving
fluid from the reservoir by electrolysis pumping. An enucleatedporcine eye was prepared for the surgical study and a limbal
incision was made (between the cornea and sclera). The cannula
was implanted through the incision into the anterior chamber ofthe eye (Fig. 9). The enucleated porcine eye was pressurized to
15 mmHg by using an infusion line. Constant current (0.5 mA)
was applied for 1 min. The device was surgically removed after
the experiment.
6. Results and discussion
The electrolysis pump was operated at flow rates in the
pL/min to ␮L/min range using driving currents from 5 ␮A to
1.25 mA (Fig. 10A–B). The highest flow rate was 7 ␮L/min for1.25 mA and the lowest was 438 pL/min at 5 ␮A. Both data sets
were corrected to compensate for the evaporation of fluid during

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Fig. 11. Bolus delivery of 250 nL doses using current pulses.

Fig. 9. Ex vivo testing of the device in porcine eye showing the electrolysis
driven delivery of dyed deionized water into anterior chamber.

testing. Flow rates below about 2 ␮L/min are preferred for ocular drug delivery. This is consistent with naturally occurring flow
rates in the eye; the ciliary body of the eye produces aqueous
humor at 2.4 ± 0.6 ␮L/min in adults [22]. As current decreases,
it was observed that pumping efficiency, which ranged from 24%
to 49%, also decreased (Fig. 10C). These efficiency values are
consistent with those reported for other electrochemical pumps[15]. Electrolysis-driven pumping efficiency was affected by
the competitive recombination of hydrogen and oxygen gases
to water. This effect was further accelerated by exposure to
the platinum electrolysis electrodes which serve to catalyze therecombination reaction. In Fig. 10D, a typical accumulated volume curve is shown that illustrates the effect of recombination
after the applied current is turned off. The measured recombination rate was 62 nL/min.
Bolus delivery mode was also evaluated (Fig. 11). If thedesired dosing regimen requires 250 nL per dose, this volume
can be obtained by driving the pump for a short duration that is
associated with the magnitude of the applied current. For example, a 1.0 mA driving current will dose 250 nL in 2.36 s and,for 1.5 mA current, the pulse time can be set as 1.75 s. Under
normal operation in the eye, the drug delivery device will experience a backpressure equivalent to the IOP of the eye. Bench-top
experiments indicated that the pump was able to supply sufficient drug flow (up to 1500 nL/min) over the range of normal

Fig. 10. (a) Current-controlled flow delivery after evaporation compensation (mean ± S.E., n = 4). The calibrated water evaporation rate in the micro-pipette is∼30 nL/min. (b) Low flow rate operation of the drug delivery device. (c) Pump efficiency calculated from flow delivery data. (d) Typical gas recombination observedin the drug delivery device. Current (50 ␮A) was applied for 10 min and then turned off.

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47

intraocular fluids into the device through the cannula. Operation with ophthalmic drugs and in vivo demonstration are alsoplanned.
Acknowledgements
This work was funded in part by Engineering Research Centers Program of the NSF under Award Number EEC-0310723
and Bausch & Lomb. We would like to thank Mr. Trevor Roper,
Dr. Tuan Hoang, and members of the Biomedical Microsystems
Lab at the University of Southern California.
Fig. 12. Flow performance under physiological back pressures (mean ± S.E.,
n = 4).

and abnormal IOP equivalent backpressures (Fig. 12). The flow
rates subjected to abnormal IOPs varied by approximately 30%
compared to normal IOP values.
Our device has demonstrated the flexibility to achieve pL
to ␮L/min flow ranges suitable for a wide range of intraocularpharmacological therapies. For example, drugs used in current
sustained-release implants for treating conditions that affect ocular tissues in the difficult-to-reach posterior segment of the eye
are fluocinolone acetonide and ganciclovir. A fluocinolone acetonide implant for treating posterior uveitis was demonstrated
to release drug at 0.6 ␮g/day for over 1000 days [23]. A ganciclovir implant for treating cytomegalovirus retinitis released at
48 ␮g/day for over 190 days [24]. In our electrolysis-based delivery system, ∼10% by weight liquid drug formulation with anestimated solution density of 1.0 g/cm3 will be used. For release
rates corresponding to 0.6–48 ␮g/day, the respective drug delivery rates of 6.5–520 nL/day are required and fall within theperformance range of our device. Extensive animal testing isrequired to fully assess the pharmacological efficacy and safetyof these flow ranges and is planned.
Initial surgical results yielded promising results in enucleatedporcine eyes. Following removal of the device after the surgical experiment, post-surgical examination of the eye throughan ophthalmic operating microscope revealed a small spot ofdye located above the iris near the position of the cannula tip
indicating that dye was delivered into the eye.
7. Conclusion
An electrochemically driven implantable MEMS drug delivery device featuring integration of a pump, drug reservoir,
and transscleral cannula was prototyped and tested. Bench-top
testing results demonstrated that this electrochemically driven
micropump can provide a wide operation range (pL/min to␮L/min) suited for ocular drug delivery treatment of chronic
diseases. Drug delivery against opposing backpressures equivalent to normal and elevated intraocular pressures was also
demonstrated. Preliminary ex vivo testing was performed
demonstrating the feasibility of this device for ocular drug delivery. Future work will focus on optimizing the electrolysis pumpand the integration of a check valve to prevent back flow ofReferences[1] D.H. Geroski, H.F. Edelhauser, Drug delivery for posterior segment eye
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Biographies
Po-Ying “Brian” Li received his BS and MS degree in mechanical engineering from Tatung Institute of Technology, Taipei, Taiwan, in 1996 and National
Tsing Hua University, Hsinchu, Taiwan, in 2001. He also received an MS degree
in materials science from University of Southern California, Los Angeles, in
2004. Currently, he is a PhD candidate in electrical engineering in University ofSouthern California and member of the Biomedical Microsystems Laboratory.
His research activities include drug delivery systems, thermal flow sensors, andelectrothermal valves. His broad research interests include BioMEMS, NEMS,
solid state physics, microfluidics, microactuators and microsensors, crystallography, solid state physics, materials characterization, solid mechanics, finite
element analysis, and bio-electronic packaging.
Jason Shih received his BS from the California Institute of Technology in 2003.
He is currently pursuing his PhD in electrical engineering at the same institution.
He is a member of the Caltech Micromachining Laboratory and his researchinterests are in the field of microfluidics. Current projects include chip based
liquid chromatography systems, miniature devices for drug delivery, and early
detection systems for biowarfare agents.
Ronalee Lo received her BS from Harvey Mudd College in 2001. She worked
at a Washington, D.C. based biotechnology consulting firm, BioStar, for one
year prior to enrolling at the University of Southern California to pursue a PhD
She received a MS in biomedical engineering medical devices and diagnostics
from the University of Southern California in 2004. She is currently a memberof the Biomedical Microsystems Laboratory and is working toward a PhD inbiomedical engineering. Her current research focuses on MEMS drug delivery
devices and microfluidic interconnects.

Saloomeh Saati received her MD from IRAN University of Medical Sciences in 1999. She finished her training by completing an ophthalmology
residency at Tehran University of Medical Sciences in 2005. Currently, she is a
research fellow of ophthalmology at the University of Southern California. She
is working on Artificial Retina Project and bioMEMS drug delivery devices.
She is a member of Association for Research in Vision and Ophthalmology(ARVO).
Rajat Agrawal received his MD from Nagpur University, India, in 1990. He
finished an ophthalmology residency in 1993 from Mahatma Gandhi Memorial
Hospital in Bombay, India, and fellowship training in 1995 in Vitreo-Retinal
Surgery and Diseases from the Medical Research Foundation in Chennai, India.
He is a Diplomate of the National Board in Ophthalmology. Dr Agrawal is currently an Assistant Professor of Ophthalmology at the University of Southern
California and Co-Director of Intraocular Implants at the Doheny Eye Institute,Los Angeles. He currently directs the development of Artificial Retinal Prosthesis program. He is also involved in development of new ophthalmic procedures
and treatments as well as development of new ophthalmic instruments and drugdelivery devices.
Mark S. Humayun received his BS from Georgetown University in 1984, his
MD from Duke University in 1989, and his PhD from the University of North
Carolina, Chapel Hill in 1994. He finished his training by completing an Ophthalmology residency at Duke and a Fellowship in Vitreoretinal diseases at
Johns Hopkins Hospital. Currently, Dr. Humayun is a professor of Ophthalmology, Biomedical Engineering, and Cell and Neurobiology at the Universityof Southern California. Dr. Humayun is the Director of the National Science
Foundation Biomimetic MicroElectronic Systems Engineering Research Center (BMES ERC). He is also the Director for the Department of Energy Artificial
Retina Project that is a consortium of five Department of Energy labs and four
universities, as well as industry.
Yu-Chong Tai received his BS degree from National Taiwan University, andthe MS and PhD degrees in electrical engineering from the University of California at Berkeley. He is currently a professor of Electrical Engineering andBioengineering at the California Institute of Technology, and director of theCaltech Micromachining Laboratory. His current research interests include flexible MEMS, bioMEMS, MEMS for retinal implants, parylene-based integrated
microfluidics, neuroprobes/neurochips, and HPLC based labs-on-a-chip.
Ellis Meng received her BS degree in engineering and applied science from the
California Institute of Technology in 1997. She pursued her graduate studies
in electrical engineering and received her MS in 1998 and PhD in 2003 at thesame institution. She is now an assistant professor of Biomedical Engineering
at the University of Southern California. In the National Science Foundation
Biomimetic MicroElectronic Systems Engineering Research Center (BMES
ERC) she is a Thrust Leader for Interface Technology and the Associate Directorof Education and Student Diversity. She is a member of Tau Beta Pi, the American Society of Mechanical Engineers (ASME), the Society of Women Engineers(SWE), and the Association for Research in Vision and Ophthalmology (ARVO).