Sensors and Actuators A 166 (2011) 193–200

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Sensors and Actuators A: Physicaljournal homepage: www.elsevier.com/locate/sna

Parylene-based integrated wireless single-channel neurostimulator
Wen Li a,∗ , Damien C. Rodger b , Anderson Pinto b , Ellis Meng c , James D. Weiland b ,Mark S. Humayun b , Yu-Chong Tai aaDepartment of Electrical Engineering, California Institute of Technology, Pasadena, CA 91125, USA
Keck School of Medicine, University of Southern California, Los Angeles, CA 90033, USAc
Department of Biomedical Engineering, University Park Campus, University of Southern California, Los Angeles, CA 90089, USAba r t i c l e

i n f oArticle history:
Received 15 April 2009
Received in revised form
28 November 2009
Accepted 4 March 2010
Available online 16 March 2010
Keywords:
BIONs
Implantable MEMS
Integrated neurostimulator
Parylene C
Retinal prosthesis
Wirelessa b s t r a c t
This paper presents the design, fabrication, and functional testing of a fully implantable, ﬂexible, Paryleneenable neurostimulator that features single-channel wireless stimulation capability [1]. This systemcomprises a fold-and-bond microelectromechanical systems (MEMS) coil for wireless power and datatransmission, a BION 1-3 CMOS stimulator chip, discrete capacitors, as well as a carrier substrate withtwo platinum stimulating electrodes and interconnections for system assembly. The physical geometries of the devices are designed for use in retinal implantation with the speciﬁcations determined fromthe implantation results in canine eyes. The coil and carrier substrate are separately fabricated using a
Parylene–metal skin technology. The unique properties of Parylene C allow these MEMS devices to be
ﬂexible and biocompatible, facilitating surgical procedure. The system assembly is achieved by interconnecting individual components together on the carrier substrate with a biocompatible silver epoxy.
A 10 ␮m layer of Parylene C is selectively deposited on the integrated system to protect it from corrosive eye environment. The system functionality is veriﬁed using a telemetry link setup, and single-phase
pulses with amplitudes ranging from 7 to 8.5 V are detected.
© 2010 Elsevier B.V. All rights reserved.

1. Introduction
Nervous system damages (e.g. brain trauma, spinal cord injury,
and retinal diseases) result in permanent physical disabilities tomillions of people worldwide [2,3] and are presently incurable conditions. Potential treatments such as neural transplantation andstem cell therapy [4–6] are promising but their implementation
faces great challenges due to the lack of clinical practice protocols,
high costs, and political issues. Artiﬁcial neural prostheses, which
utilize electronic devices as interfaces to central and peripheral
neural systems, have been increasingly accepted to partially treat
neurological disorders and diseases. Many different approaches
have been investigated, using electrical signal to stimulate neurons
and muscles for controlling of limbs and organs, or to elicit sensory
sensation for feeling, hearing, or seeing. Numerous applications
include limb prostheses for spinal cord injury and stroke, sacral
nerve stimulations for bladder emptying, deep brain stimulation(DBS) for essential tremor and the tremor of Parkinson’s disease,cochlear prostheses for restoring hearing, and retinal prostheses∗ Corresponding author at: Michigan State University, Electrical & Computer Engineering, 2120 Engineering Building, East Lansing, MI 48824, USA.
Tel.: +1 517 353 7832; fax: +1 517 353 1980.
E-mail address: wenli@egr.msu.edu (W. Li).
0924-4247/$ – see front matter © 2010 Elsevier B.V. All rights reserved.
doi:10.1016/j.sna.2010.03.003

for outer retinal degenerative diseases such as age-related maculardegeneration (AMD) and retinitis pigmentosa (RP), to name just a
few [7–11].
In recent decades, integrated wireless microsystems emerge
and provide breakthrough opportunities for implantable neural
prosthesis. Microfabrication technologies can effectively increase
the degree of miniaturization and system complexity. Wirelesspower and data signal transmission through electromagnetic ﬁelds
avoids batteries or permanent wired connections, and thus reduces
surgical complexity as well as undesired physical damages topatients. Among the aforementioned examples of neural prostheses, an epiretinal prosthesis is one of the most prominent
prosthetic interfaces that will beneﬁt from wireless microsystemtechnologies. An epiretinal implant system typically comprises (1)radio-frequency (RF) coils for transferring power and data between
intraocular units and external data acquisition units, (2) integrated
circuitry for converting signal and providing stimulus to retina, and(3) a multielectrode array for stimulating neural cells on the innerretina (Fig. 1). For applications that involve completely intraocular implantation in human eyes, the following requirements on
the system and encapsulation method are critical: (1) sufﬁciently
high spatial selectivity, (2) biocompatibility and low cytotoxicity,(3) enough ﬂexibility to conform to anatomical curvatures of eyeballs, minimize tissue damage, and facilitate surgical procedure.
Current technologies using ﬂexible polymer-based microsystems

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results indicate that Parylene C is a biocompatible material suitable
for long-term intraocular retinal prostheses.
2. System design

Fig. 1. Conceptual schematic of an epiretinal prosthesis showing main components
and their placement.

for epiretinal implants are still in developmental stage [12–14] andfew have actually been transferred to the clinical practice due tochallenges with respect to material selection, device miniaturization, as well as fabrication and integration methods.
Our goal is to address these existing issues, by integrating
Parylene-based microelectromechanical systems (MEMS) devices(microcoils and electrode arrays) with other discrete components(application speciﬁc integrated circuits (ASICs), capacitors, etc.) todevelop ﬂexible and biocompatible wireless neural implants. In
this work, we present the ﬁrst completely Parylene-based functional neural stimulator, which consists of a MEMS coil and anelectrode array integrated with a single-channel stimulator chip.
For the ﬁrst prototype, MEMS devices are made separately, andassembled with other system components using biocompatible silver epoxy. A Parylene–metal skin technology has been developed,
which allows us to microfabricate RF coils and high-density multielectrode arrays in a process compatible way [15,16]. Parylene C ischosen as the main substrate and packaging material because of its
favorable properties, such as ﬂexibility (Young’s modulus ∼4 GPa),
optical transparency, chemical inertness, and biocompatibility [17].
In fact, Parylene C is an ISO 10993 and United States Pharmacopeia(USP) Class VI biocompatible material, the highest biocompatibility
and implant clearance among plastics. Upon intraocular implantation in two rabbits for 6 months, Parylene C is shown to generate
no detectable immune response affecting their retinas [18]. These

Fig. 2 depicts the system schematic of the singlechannel simulator, which consists of a complementary
metal–oxide–semiconductor (CMOS) chip, two capacitors, a
specially designed RF MEMS coil, and a carrier substrate. The
chip is a single-channel stimulator which was initially developed
by Loeb and co-worker to mimic muscle spindle function forpatients with muscle paralysis [8]. When implanted into paralyzed
muscles, power and command signals can be delivered to the chip
through inductive coupling over a 480–500 kHz power carrier
generated in an external personal trainer. Then the chip canemit precisely timed stimulation pulses with a highly regulated
amplitude and pulse-width. Fig. 3 shows a BION 1–3 chips usedin our experiments, and the pad functionalities are described in
the right hand table. This chip has physical dimensions of ∼1 mm
in width, ∼2.33 mm in length, and ∼257 ␮m in thickness. There
are ﬁve pads on the BION chip approximately 120 ␮m by 74 ␮m insize, and the distance between adjacent pads is about 110 ␮m.
A dual-metal-layer coil is designed as the main component for
wireless power and data transmission. The self-inductance of such a
MEMS coil is limited by its small dimensions, and can be calculatedusing the following equation [19]:Ls = 2dN 2 × 10−9



ln

4d
t





1+

t2
· ··
24d2



−

43t 2
1
+
· ·· (Henry),
2
288d2

(1)

where N is the total number of turns, d (in cm) is the mean diameter of the coil, and t (in cm) is the coil width. Two ceramic chipcapacitors (AVX Corporation, Myrtle Beach, SC, USA) are also incorporated into the system for circuit function, as shown in Fig. 2. C1
is a frequency tuning capacitor in parallel with the receiving coil toachieve a resonant frequency of ∼500 kHz, and C2 is a charge storage capacitor which has a capacitance of ∼22 nF. The size of thesecapacitors is approximately 1 mm in length, 0.5 mm in width, and0.56 mm in thickness. Finally, the carrier substrate is designed forsystem assembly, which comprises a ﬂexible cable, interconnection
leads and contact pads, and two stimulus electrodes of which one
serves as the signal terminal and the other as the ﬂoating ground.

Fig. 2. System schematic of the single-channel stimulator and the corresponding circuit layout, showing the connections between individual components.

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195

Fig. 3. (a) Microscope image of a BION 1-3 chip. (b) Pad function of the chip.

Overall physical dimensions of the MEMS coil and the carrier
substrate need to meet speciﬁcations used in retinal implantation.
Due to the small volume and complex anatomy of eyes, it is necessary to evaluate possible geometries in eyes prior to ﬁnalizing thedesigns of the devices. It is also important to verify the robustnessof the devices and materials under surgical manipulation. Thus,
mechanical models with various dimensions are ﬁrst fabricated andimplanted in canine eyes to study the system geometrical feasibility inside the eye, as shown in Fig. 4. In this experiment, carrier
substrates are made of Parylene C thin ﬁlms (PDS 2120 system,
Special Coating Systems, Indianapolis, IN, USA) using a photoresist
mask and oxygen plasma etching in a reactive ion etching (RIE)
system (Semi Group Inc. T1000 TP/CC). Plastic dummy coils andchips are then glued onto the carrier substrates using biocompatible epoxy EPO-TEK 301-2 (Epoxy Technology, Billerica, MA, USA),simulating the system integration of various key components. The
animal surgery is operated at the Keck School of Medicine of the
University of Southern California. During the implantation, the coilsite is placed right beneath the cornea, in the anterior chamber ofcanine eyes, less than 1 mm from external air. After lens extraction, the ﬂexible cable is passed through the anterior and posteriorlens capsule into the vitreous cavity, leaving the chip site ﬂoatingin the vitreous cavity. Vitrectomy is then performed and the electrode array site is attached on top of the inner retina using a retinal
tack modiﬁed by the addition of a PDMS washer. Based on preliminary surgical results, optimal device dimensions for canine eyes
are determined as indicated in Fig. 2. Devices for human implantations can be simply modiﬁed by reducing the cable length in orderto accommodate the smaller space of the vitreous cavity in humaneyes.
3. Fabrication
In order to reduce fabrication complexity, the coil and the carrier substrate are fabricated and calibrated respectively. Then thesystem is assembled by interconnecting individual components on
the carrier substrate with a conductive silver paste. Detailed fab-Fig. 4. A dummy structure for canine implantation.

rication is divided into three steps and described in the following
sections.
3.1. Carrier substrate fabrication
To build the ﬂexible substrate, a 200 nm layer of metal is e-beam(SE600 RAP, CHA Industries, Fremont, CA, USA) evaporated on a
Parylene C coated silicon substrate. Then the metal is patternedusing a lift-off method to form connection pads, interconnection
leads, and electrode array [16]. Platinum is selected as the primaryelectrode material due to its optimal stimulation capability andbiocompatibility. Parylene C-based arrays of thin-ﬁlm platinum
electrodes have shown excellent biostability when chronicallyimplanted in contact with canine retinas for 6 months [16]. Furthermore, platinum and Parylene are known to have good adhesion,requiring no additional adhesion metal layer during fabrication.
After metal patterning, another layer of Parylene C is deposited toseal the entire structure, followed by oxygen plasma etching witha photoresist mask to deﬁne the contour of the carrier substrate, as
well as to open the electrode sites and the contact vias. Finally, thedevice is peeled off from the silicon substrate in a water bath anddried in air. Fig. 5 illustrates the detailed process for making the carrier substrate, where steps (b)–(d) describe the lift-off technology
for platinum patterning.
Fig. 6 presents a fabricated carrier substrate. This device has atotal thickness of approximately 10 ␮m. Three extra pads are addedon the chip site for testing purpose and can be removed after system
integration. The microscope image in Fig. 6 shows a special chip site
design where Parylene ribbons are etched in a way such that thechip can be held in place and self-aligned to the contact vias on thesubstrate during system assembly. The array site contains a 450 ␮mdiameter tack hole so that it can be attached to the retina with theretinal tack during implantation.
3.2. RF MEMS coil fabrication
A fold-and-bond technology is involved in the MEMS coil fabrication [20]. First, a planar Parylene–metal–Parylene sandwich
structure is built with two coils arranged in series (Fig. 8). The
fabrication process is show in Fig. 7. Brieﬂy, a 3 ␮m layer of gold
is deposited on a Parylene coated substrate and patterned withmetal etching process. Then Parylene deposition is performed on
top of metal, followed by oxygen plasma etching to open contact
vias. Photoresist serves as a sacriﬁcial layer so the structures can bereleased from the substrate by dissolving photoresist in acetone.
After the planar structure is microfabricated, it is folded into
two layers and stacked together with the assistance of two glass
slides. Aluminum sheets are inserted between the Parylene surface and the glass slides to prevent Parylene from sticking to glass.
External pressure is applied by stacking a ∼2 lbs metal plate on
top of a 75 mm × 50 mm glass slide, allowing Parylene-to-Parylene
bonding at modest temperatures. After that, the whole unit isplaced in a vacuum oven (T-M Vacuum Products, Inc., Cinnamin-

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W. Li et al. / Sensors and Actuators A 166 (2011) 193–200

Fig. 5. Fabrication process of the Parylene-based carrier substrate.

Fig. 6. A fabricated carrier substrate with electrodes, interconnection leads and contacts. (a) Parylene ribbons for holding the BION chip. (b) A 2-electrode array with a tackhole.

son, NJ, USA) with the chamber pressure of ∼10 Torr for bonding.
The oven temperature ramps from room temperature to bonding temperature, and then soaks at the bonding temperature for
2 days, followed by slowly cooling to the room temperature. Nitrogen backﬁll is introduced to the chamber in order to equalize
the chamber temperature. Experiments are performed at a setting temperature range of 180–250 ◦ C, and bonding results areevaluated by visual inspection. The results demonstrate that thevacuum pressure of 10 Torr can prevent Parylene from undesired oxidation at elevated temperatures so that the device canremain ﬂexible after bonding. It is also found that good Paryleneto-Parylene bonding occurs in samples annealed at temperatureshigher than 230 ◦ C. It should be noted that the bonding temperature has to be lower than the melting temperature of Parylene C(290 ◦ C), as strong recrystallization can happen beyond this point[21].
A fabricated coil is shown in Fig. 8, comprising two layers ofmetal with 10 turns in each layer. In-and-out leads are connectedto the carrier substrate from the center to facilitate the surgicalprocedure. Through vias are designed to overlap with the contact
pads so that the interconnections to the carrier substrate can beformed from either side. The electrical properties of this coil aremeasured with an Agilent 34401A multimeter and the HP 4192A LF
impedance analyzer (HP/Agilent Technologies Inc., Santa Clara, CA,USA), showing an inductance of approximately 2.24 ␮H and a DC
resistance of approximately 15.82 . A Q factor of 0.45 is calculated
at the operation frequency of 500 kHz. Theoretically, thicker metal
can lower the intrinsic resistance of the coil to achieve higher Q

Fig. 7. Fabrication process of Parylene–metal–Parylene sandwich structure.

W. Li et al. / Sensors and Actuators A 166 (2011) 193–200

Fig. 8. A fabricated fold-and-bond coil with two layers of metal. The electrical characteristics are measured and given in the table.

factor and better performance, but e-beam evaporated metals areoften limited in ﬁlm thickness due to process cost.
3.3. System assembly and packaging
For hybrid system assembly, discrete components (the BION
chip, the coil and the capacitors) are aligned to corresponding
interconnection vias on the carrier substrate. A small amount of
biocompatible silver epoxy EPO-TEK H20E is then applied on the197

contacts and cured at 80 ◦ C for 3 h in a convection oven. The conductive epoxy serves two purposes: to form the interconnections
between the components as well as to bond the components onto
the substrate. Fig. 9 shows an assembled single-channel stimulator system, and close-up views of the interconnections for eachcomponent.
Because the BION chip used in our experiment has only ﬁve pads,
hand assembly can be used as a temporary solution for fast integration. However, a major problem of manual assembly is the inability
to control epoxy dosage when applied by hand. Short circuits canbe created if too much epoxy is applied, and epoxy reﬂow can also
happen during high temperature curing, resulting in short circuits
between adjacent pads. For circuitry with high-density pad layouts, this hand assembly is no longer feasible, and thus a wafer
level integration technology becomes necessary [22].
Another big concern is the isolation of the silver epoxy from theeye environment. Human eye is known as a delicate and sensitive
organ. Any toxin released from the epoxy can cause inﬂammationand infection. In addition, eye ﬂuids are corrosive to exposed metals [23], which can reduce the durability of silver epoxy bonding,and ﬁnally result in open circuits. To overcome these challenges,∼10 ␮m Parylene C is coated over the entire system just leaving
the electrode sites open to protect the circuitry from corrosive eye
ﬂuids. The packaging performance of Parylene C has been studied
using accelerated-lifetime soak tests in related work [24]. Preliminary results indicate that Parylene C protected metal can remain

Fig. 9. (a) An assembled BION system. (b) Interconnects of coil contacts formed with biocompatible conductive silver epoxy. (c) Interconnects of chip and capacitor contacts
formed with biocompatible conductive silver epoxy.

Fig. 10. (a) A telemetry setup for functionality test of the assembled BION system. (b) A close-up image showing the tested system placed between the primary coil.

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Fig. 11. Recorded stimulation pulses at different separation distances between two
coils.

intact in saline (0.9% NaCl) at body temperature (37 ◦ C) for over 20
years, which is very promising for our application.
4. Test results and discussions
The system functionality is veriﬁed using a telemetry link setup(Fig. 10). The primary stage, which comprises a personal trainer
unit, a class-E coil driver, and a hand-wound transmitting coil, cangenerate a power carrier of approximately 500 kHz. The personaltrainer stores command programs personalized for individual subjects, records the time and duration of treatment, and transfers
this information to an external computer for real time monitoring. Upto three programs can be preloaded into the memory of
the personal trainer [25]. The coil driver is connected to the personal trainer through a custom made adapter. The transmitting coilhas an inductance of ∼46.4 ␮H and a Q factor of ∼118 at 500 kHz.
Litz wires 1025-44 SPN are used for winding the primary coil inorder to reduce the skin effect and proximity effect losses. The
transmitting coil is built in a solenoid shape to establish a more
uniform electromagnetic ﬁeld inside the coil. It is also known thatthe induced voltage at the receiving coil increases proportionallywith the increase of magnetic ﬂux going through the inductive
link. Therefore, ferrite cores are incorporated in the transmitting
coil to magnify electromagnetic ﬁeld and improve voltage transfer
efﬁciency,
In vitro measurements have been conducted in air using thissetup, and output signals of the integrated stimulator are monitored by connecting two electrodes directly to an HP 54645A
oscilloscope. The distance between the two coils is varied during
testing, and a maximum detectable range of ∼4 mm is found. The
operation range is relatively short, suffering from the low Q factorof the receiving coil as expected. The recorded stimulating pulses
at the different separation distances are given in Fig. 11, showing apulse width of approximately 500 ␮s, and amplitudes varying from
7 to 8.5 V.
In order to estimate the power transfer capability of the MEMS
coil, voltage across the receiving coil terminals and current delivered to the chip are measured. Fig. 12 shows typical waveforms ofthe transferred voltage and current, indicating a resonant frequencyof ∼505 kHz and a ∼25◦ phase drift between voltage and current. As
mentioned earlier, the tuning capacitor is a commercially available
chip capacitor, which has limited options for capacitance values.
Therefore, it is difﬁcult to ﬁne tune the resonant circuit to achieveprecise synchronization. Delivered power at different separation
distances is also investigated, as shown in Fig. 13. It can be seenthat our MEMS coil can transfer a maximum power of ∼43 mW
through this inductive link at a separation distance of 1 mm. As theseparation distance increases to 2 mm, the power drops dramati-

Fig. 12. Typical waveforms of transferred voltage and current to the chip.

cally by 62%, mainly due to the divergence of electromagnetic ﬁeld.
When delivered power is less than 10 mW, no stimulating pulse
can be detected from the output.
Similar power transfer measurements have also been performedby immersing the integrated stimulator in regular saline solution
to simulate in vivo conditions. Preliminary experimental results
have shown that the maximum detectable range of this telemetrylink is decreased to approximately 2 mm. This signiﬁcant reduction(∼50%) is attributed to a substantial environmental permittivity
change in the liquid medium as compared with air, which resultsin the increase of magnetic energy losses into the saline and a low
Q factor of the MEMS coil [26].
Typical detection distance for retinal prostheses in current clinical applications is estimated to be approximately 15 mm, which isdeﬁned as the coaxial separation distance between the implanted
coil and the external coil. In our system, the internal coil is placed
right beneath the cornea, and the limiting factor in reducing thisdistance is the large separation of external coil from eye surface forsafety reason. To accommodate such requirements and increase thepracticality of our stimulator, further improvement can be made toincrease the detectable range by optimizing coil design, such asincreasing metal thickness and/or the number of metal layers, toenhance the coil Q factor and the power transfer efﬁciency. Particularly, for a fold-and-bond coil used in this integrated system,
the coil Q factor is proportional to the number of folds and can beenhanced by n-times when n single layers of Parylene–metal thinﬁlms are stacked and bonded together [20].
The functionality of the integrated system has been successfully
demonstrated in both air and saline. Some efforts have been carried out to test the mechanical reliability of the integrated system
in rabbit eyes, as well as to optimize the surgical procedure. Fig. 14
shows an example of the acute implantation, in which an integrated system is implanted in a rabbit eye. After implantation, theFig. 13. Transferred power at different separation distances between two coils.

W. Li et al. / Sensors and Actuators A 166 (2011) 193–200

199

Fig. 14. Acute implantation images of an integrated system in a rabbit eye, showing: (a) the coil sitting in front of the iris; (b) the ﬂexible cable and array implanted in thevitreous cavity.

subject is sacriﬁced and its eyeball is soaked in forming solution
for 1–2 weeks to maintain the eyeball shape. Then the eyeball isremoved from the solution and cut into halves in order to examine
the implanted device. Preliminary outcomes suggest the implanted
device can mechanically withstand the surgical procedure ﬁne.
Future work will aim to improve the power transfer performanceof the telemetry link as well as to verify the system functionality
and biostability using ex-vivo testing and in vivo implants in rabbit
subjects.
5. Conclusion
A Parylene-based single-channel neural stimulator has been
designed, and the ﬁrst prototype has been successfully microfabricated and assembled. In vitro active measurements are performed
to verify the functionality of the integrated system. The test results
demonstrate that the BION chip can be driven by the MEMS coilwithin a 4 mm separation distance in air. Output pulses with∼500 ␮s pulse width and greater than 7 V amplitude are detectedfrom the stimulating electrode. Preliminary animal implantation
result indicates that the integrated system is mechanically reliable and surgically compatible. We are expecting to implant thedevice in animal subjects to further characterize its system performance in vivo. Although speciﬁcally tailored to the needs of retinal
prostheses, this system integrates Parylene-based MEMS devices
with discrete electrical components and ASICs, and thus enables a
new range of true system solutions for both biomedical and nonbiomedical applications.
Acknowledgments
This work is supported in part by the Engineering ResearchCenter Program of the National Science Foundation under Award
Number EEC-0310723 and by a fellowship from the Whitaker Foundation (D.R.). The authors would like to thank Dr. Gerald Loeb andhis group members for providing BION chips and testing units. My
gratitude also goes to Mr. Trevor Roper and all other members of theCaltech Micromachining Laboratory for assistance with equipment
and fabrication.
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Biographies
Wen Li received her PhD in 2008 and MS in 2004 in electrical engineering from
California Institute of Technology. Prior to that, she studied at Tsinghua University
and received an MS degree in microelectronics in 2003 and a BS degree in material science and engineering in 2001. Dr. Li is currently an assistant professor ofelectrical and computer engineering at the Michigan State University. Her researchinterests include MEMS/NEMS technologies and systems, micro sensors and actuators, biomimetic devices and systems, microﬂuidic and lab-on-chip systems, andmicrosystem integration and packaging technologies. She is a member of the Institute of Electrical and Electronics Engineers, the Eta Kappa Nu, and the American
Chemical Society.
Damien C. Rodger received his BS in electrical engineering (magna cum laude withhonors) from Cornell University in 2000, and his PhD in bioengineering from the
California Institute of Technology in 2008. He earned his MD at the Keck School of
Medicine of the University of Southern California in 2009 and conducts researchon microtechnologies for retinal and spinal cord prostheses and on other novel
bioMEMS for ophthalmic use. He will be pursuing a residency in ophthalmology
at the Doheny Eye Institute in Los Angeles, CA. He held a Whitaker Foundation
Graduate Fellowship from 2003 to 2006 and is a member of the IEEE EMBS and
the Association for Research in Vision and Ophthalmology (ARVO).
Anderson Pinto received his MD from University Severino Sombra, Rio de Janeiro,
Brazil. He completed his residency in Ophthalmology at University of Mogi das
Cruzes, São Paulo, Brazil. After that, he went to Federal University of São Paulo,
Brazil, for a Fellowship in retina and vitreous. He also earned a PhD degree from thesame university. Currently, he is a post-doctoral student at the Doheny Eye Institutein Los Angeles, CA.
Ellis Meng received the BS degree in engineering and applied science and the MS
and PhD degrees in electrical engineering from the California Institute of Technology (Caltech), Pasadena, in 1997, 1998, and 2003, respectively. Since 2004, she
has been an Assistant Professor in the Department of Biomedical Engineering, University of Southern California, Los Angeles. She holds a joint appointment in the
Ming Hsieh Department of Electrical Engineering. Her research interests include
bioMEMS, implantable biomedical microdevices, microﬂuidics, multimodality inte-

grated microsystems, and packaging. Currently, she is the Thrust Leader for Interface
Technology in the National Science Foundation Biomimetic MicroElectronic Systems Engineering Research Center (BMES ERC). She also holds the Viterbi Early
Career Chair in the Viterbi School of Engineering. Dr. Meng is a member of Tau
Beta Pi, Biomedical Engineering Society, Society of Women Engineers, and American Society for Engineering Education. She was a recipient of the Intel Women in
Science and Engineering Scholarship, Caltech Alumni Association Donald S. Clark
Award, and Caltech Special Institute Fellowship. She has also received the NSF
CAREER and Wallace H. Coulter Foundation Early Career Translational Researchawards.
James D. Weiland received his BS from the University of Michigan in 1988. After
4 years in industry with Pratt & Whitney Aircraft Engines, he returned to Michigan
for graduate school, earning degrees in biomedical engineering (MS in 1993, PhD in1997) and electrical engineering (MS in 1995). In 1999 he was appointed an assistant professor of ophthalmology at Johns Hopkins. Dr. Weiland is now the Directorof the Intraocular Retinal Prosthesis Lab at the Doheny Retina Institute, and is anassociate professor of ophthalmology and biomedical engineering at the Universityof Southern California. He is a member of the IEEE EMBS, the Biomedical Engineering
Society, and the Association for Research in Vision and Ophthalmology.
Mark S. Humayun received his BS from Georgetown University in 1984, his MD fromDuke 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 andneurobiology at the University of Southern California. Dr. Humayun is the Director of
the National Science Foundation Biomimetic MicroElectronic Systems Engineering
Research Center (BMES ERC) and 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. He is also a member of the National Academies
Institute of Medicine and the Institute of Electrical and Electronics Engineers(IEEE).
Yu-Chong Tai is a Professor of Electrical Engineering, Mechanical Engineering andBioengineering at the California Institute of Technology. He developed the ﬁrst electrically spun polysilicon micromotor at UC Berkeley. At Caltech, his main researchinterest is MEMS for biomedical applications including lab-on-a-chip and neural
implants. He has published more than 300 technical articles in the ﬁeld of MEMS.
He is the recipient of the IBM Fellowship, Ross Tucker Award, Best Thesis Award(at Berkeley), Presidential Young Investigator (PYI) Award, Packard Award and ALA
Achievement Award. He is a fellow of IEEE.