JOURNAL OF MICROELECTROMECHANICAL SYSTEMS,...

JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 15, NO. 2, APRIL 2006 277

Microoptical Characterization of PiezoelectricVibratory Microinjections in DrosophilaEmbryos for Genome-Wide RNAi Screen

Xiaojing Zhang, Member, IEEE, Matthew P. Scott, Calvin F. Quate, Life Fellow, IEEE, andOlav Solgaard, Member, IEEE

Abstract—In this paper, we study the effect of acoustic agitationon the penetration force for microinjections in Drosophila embryosfor genome-wide RNA interference (RNAi) screens, using an inte-grated optical MEMS force encoder for in vivo characterization ofthe dynamic penetration forces. Two modes of operation are inves-tigated. In the first mode of operation, the injector is brought intocontact and acts on the embryo with a fixed force, and the vibrationamplitude of the microinjector is increased till penetration occur.We observed a linear decrease in the penetration force of 1.6 Nwith every 0.1 m/s tip velocity increase. In the second mode of op-eration, the vibration amplitude is kept constant and the injector ispushed into the embryo until penetration. We simulate the opticalforce encoder eigenmodes and measure the injection force over thefrequency range from 0 to 16 kHz with actuation voltages up to150 V. Among the eight encoder eigenmodes with resonant fre-quency up to 16 kHz, the longitudinal vibration along the injector isshown to dominate the force reduction at 14 kHz. Two other modes,both involving significant out-of-plane injector motion, reduce thepenetration force by 52% around 3.1 kHz. The average penetra-tion force is calculated based on injections into multiple embryosfor each experimental condition. For each microinjection, the peak(or average) penetration force can be derived from the peak (oraverage) relative displacement of the two gratings upon penetra-tion. The achieved minimum peak penetration force was 15.6 N( 29 7% of the static penetration force), while the minimum av-erage penetration force was 2.7 N (5.1% of the static penetrationforce). [1547]

Index Terms—Drosophila embryo, force sensor, microsurgery,optical encoder, vibratory microinjection.

I. INTRODUCTION

THE knowledge created by recent genome sequencingprojects has brought unprecedented opportunities to

further study the genetic and molecular mechanisms of de-velopment and disease. The genome sequence of the fruitfly, Drosophila melanogaster (see Fig. 1), was published in2000 [1], and has enabled systematic studies of the functionsof the approximately 13 600 Drosophila genes. A powerful

Manuscript received March 11, 2005; revised October 3, 2005. This workwas funded by DARPA [Bio:Info:Micro] Program MDA972-00-1-0032.Subject Editor S. Shoji.

X. Zhang is with the Department of Biomedical Engineering, Texas Mate-rials Institute, and Microelectronics Research Center at J. J. Pickle ResearchCampus, The University of Texas at Austin, Austin, TX 78758 USA (e-mail:[email protected]).

M. P. Scott is with the School of Medicine, Stanford University, Stanford, CA94305 USA.

C. F. Quate and O. Solgaard are with E. L. Ginzton Laboratory, StanfordUniversity, Stanford, CA 94305 USA.

Digital Object Identifier 10.1109/JMEMS.2006.872242

technique for learning about gene functions is RNA-inter-ference (RNAi) [2], [3] through microinjections [4]. Specificgenes are silenced by the presence of dsRNA (double-strandedRNA). An observed change in phenotype indicates the functionof the silenced gene. RNAi-induced gene silencing controlsgene expression at all levels, including transcription, mRNAstability and translation. We are only beginning to understandthe physiological roles of the RNAi pathway and the functionof the many small noncoding RNA species, which are foundin eukaryotic genomes. In RNAi microinjection experiments,typically 100–200 fly embryos per assessed gene are injectedduring the first 60 min of their development, each with 60 plof dsRNA. However, common manual injection techniques in-volves injecting embryos and cells one at a time with individualglass micropipettes observed under a microscope [5], [6],which is extremely labor intensive, unreliable and inefficient.It is practical only when treating small numbers of embryosand cells, and therefore not well suited for high-throughputinvestigations. To improve accuracy, reliability, and efficiencyof RNAi experiments, we designed silicon nitride based mi-croinjectors aimed for high-throughput treatment of Drosophilaembryos [7], and demonstrated high injection efficiency ofthe silicon nitride injector tip [8] compared to conventionalglass-drawn needle tips.

Minimization of the penetration force is critical to the designof high throughput micro-manipulating instruments for biologyand genetics studies, such as RNAi for gene silencing. Vibra-tion is a well-known method for reducing the cutting force ofmacroscopic tools. In biomedical instrumentation, applying me-chanical cutting forces at kilohertz frequencies is referred to asvibratory surgery (or ultrasonic surgery for the operating fre-quency above the audible range). Arai et al. demonstrated a vi-bratory microknife using a commercial probe needle with as-sembled multilayer piezoelectric actuator for efficient cuttingof the onion epidermal cell [9]. The cutting force was measuredby touch-probe sensors that have a similar operating principleas tapping mode Atomic Force Microscope (AFM) [10], [11].For MEMS vibratory surgical devices, improved efficiency wasfirst observed in dissecting cataractous lenses using piezoelec-trically actuated, centimeter-scale silicon cutters [12]. The cut-ting force was later characterized using piezoresistors [13] andlongitudinal strain sensors [14].

Non-contact measurement of vibration of microsurgical toolsthat have micron peak-to-peak vibration amplitudes, is desir-able for system simplicity and compactness. However, com-

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278 JOURNAL OF MICROELECTROMECHANICAL SYSTEMS, VOL. 15, NO. 2, APRIL 2006

Fig. 1. Drosophila and Drosophila embryos. (a) Scanning electron micrograph (SEM) of Drosophila. (b) A typical Drosophila embryo after its outer membrane,the chorion, has being removed without harming the embryo.

Fig. 2. Vibratory microinjector with integrated optical MEMS force encoder. A piezoelectric stack (consisting of piezoelectric ceramic layers that are assemblesin series mechanically and in parallel electrically) is bonded onto the backside of the force sensor chip. The probe displacement at the maximum actuation voltageof 150 V is 17.4 �m �2:0 �m.

mercially available optical vibrometers are modified versionsof Twyman-Green interferometers and have complicated struc-tures leading to alignment difficulties [15]. Recently, a compactvibration measurement system was demonstrated in measure-ments of vibration amplitudes of several micrometers using aself-mixing laser diode [16], with the limitation that the vibra-tion amplitude must be less than half-wavelength of the illumi-nating light.

We demonstrated a microoptical, encoder-based force sensorfor characterization of the force required for penetration andinjection of Drosophila embryos [8]. The injections were per-formed both with a traditional, commonly-used drawn-glassneedle with a typical 75 m diameter tip and with the sil-icon-nitride force encoder probe with a 30- m tip. The MEMSprobe requires N to penetrate the newly hatched embryo.This is about one fourth the force needed for penetration withconventional glass needles. In addition, the nitride probe needsa shorter traveling distance to reach penetration. Dynamicoperation of the MEMS injector by off-chip piezoelectricactuation is expected to further improve injection speed andcause less damage to the embryo. In this paper, we report onpenetration force minimization through vibratory actuation of

silicon-nitride microinjectors with in vivo microoptical forcecharacterization. Actuated by piezoelectric actuators, the injec-tors are longitudinally vibrated with tip velocities controllableby actuation frequency and voltage. The forces are measuredusing MEMS optical-encoder force sensors integrated with thevibrating microinjectors.

II. THEORY

Vibratory microinjection is performed with silicon-nitrideprobes vibrated by piezoelectric actuators, as shown in Fig. 2.The probes have integrated force sensors consisting of two ver-tically separated microgratings. The two gratings are perfectlyaligned when no force or acceleration is applied. The upperindex grating is connected to the microinjector, and is designedto have eigenmodes with low kilohertz resonant frequency.The static diffraction characteristics of the force encoder undernormal plane-wave illumination can be analyzed by Fraunhoferdiffraction theory [17].

Under vibratory actuation, the index grating vibrates alongwith the microinjector and hence changes its position relative tothe fixed scale grating. The force on the injector is determined

ZHANG et al.: MICROOPTICAL CHARACTERIZATION OF PIEZOELECTRIC VIBRATORY MICROINJECTIONS 279

Fig. 3. The sensor output evolves from (a) a sinusoidal signal at small displacement of the injector to (b) a distorted output as the tip velocity increases. Afterpenetration, the encoder bias-point recovers to zero and the output signal doubles in frequency.

by the relative displacement of the two gratings, which isdetermined by the intensity distribution of the diffraction or-ders. The first diffraction mode intensity, , is a periodic func-tion of injector displacement, , as shown by the solid curve

in Fig. 3. Under vibratory actuation, the index grating vibratesalongwith the microinjector and hence changes its position rel-ative to the fixed scale grating. The force on the injector is de-termined by the relative displacement, , of the two gratings,


Fig. 4. SEM’s of force encoders for vibratory microinjections. (a) the force probe integrated with the movable index grating. (b) 3 �m uniform vertical gapbetween the dual gratings.

which is determined by the intensity distribution of the diffrac-tion orders. For the first diffraction mode, we have

(1)

where is the illuminating light intensity, is the number ofgrating periods under illumination,is the phase-delay over the thickness of one grating finger, isthe period of the grating,is phase-modulating term, is the quasistatic relative dis-placement of the two gratings caused by the linear motion of themicroinjector relative to the target, and is the sinusoidaldisplacement of the injector under piezoelectric actuation. Therelative displacement can be expressed

(2)

where

(3)

and m/V, are the piezoelectric constant, theactuation voltage, the driving frequency of the piezoactuator,respectively. The force is related to the displacement throughthe spring constant, , of the encoder

(4)

The penetration of the embryo membrane can be caused by ei-ther linear or vibration motion of the microinjector. We assumethe microinjector interacts with the membrane during the fullvibration cycles till penetration. In our studies, the linear trans-lation velocity of the microinjector is much smaller than that ofthe vibration. The average (or peak) penetration force can be de-rived from the average (or peak) relative displacement of the twogratings, using the spring constant of encoder N/m.

The dotted lines in Fig. 3 show how the force sensor outputevolves from a sinusoidal signal at small injector displacement

(case I in 3a), to a distorted signal (case II in 3b) as the injectorvibration amplitude increases. Penetration, caused by eitherincreased vibration or translation, significantly reduces the forceon the injector. Consequently, the average relative grating dis-placement goes to zero, and the output signal abruptly changesto a periodic signal with doubled frequency (case III in 3b).

III. DESIGN AND FABRICATION

The force encoders were fabricated by LPCVD depositionand patterning of two silicon nitride layers for the dual grat-ings, separated by low temperature sacrificial oxide, on the sil-icon substrates. The fabrication process is described in detail in[8]. The scanning electron micrographs (SEM’s) of the micro-machined optical encoder force sensor, with large vertical gapssuitable for vibratory actuation, are shown in Fig. 4. The SEM’sshow a force probe integrated with the movable index grating,the index and scale gratings with 20 m pitch, and a close-viewof the 3 m vertical gap and junction between the gratings andthe supporting beams.

The silicon nitride grating layers, deposited under NH -richconditions, have excellent optical surface quality. Fig. 5 showsthe surface topography characterization results of the encodergratings using optical phase-shifting interferometry (Zygo™white-light 3-D surface profiler, Middlefield, CT). The indexgrating surfaces show less than 50 nm root-mean-square rough-ness. Note that the overhanging probe appears to be bendingupwards. This is caused by the shift of optical reference planedue to the removal of the silicon substrate under the probe.

The sensor chip is completed by bonding a piezoelectricstack, consisting of many piezoelectric ceramic layers that areassembled in series mechanically and in parallel electrically,onto the back side of the encoder (see Fig. 2). The piezoelectricstack is used to vibrate the microinjector longitudinally in the

-direction. The piezoactuator has a resonance frequency of69 kHz, and a displacement at the maximum drive voltage of150 V is 17.4 m m.

IV. EXPERIMENTAL RESULTS AND DISCUSSION

The measurement setup for vibratory microinjection and invivo penetration force characterization is shown in Fig. 6. We


Fig. 5. Surface topography of encoder gratings measured by Zygo™ optical phase-shifting interferometry using scanning white light. (a) three dimensionalencoder surface profile. Note that the overhang probe appears to be bending upwards. This is due to the shift of optical reference plane caused by the removal ofthe silicon substrate under the probe. (b) the cross-sectional (A-A’) view of the step heights of the movable index grating (upper trace) and part of the scale grating(lower trace) showing less than 50 nm root-mean-square surface roughness. Positions of the grating elements are illustrated in the graph as ideal rectangular blocks.

position newly hatched Drosophila embryos1 on 2-D microflu-idic self-assembly chip [18] ready for vertical injection. The en-coder is operated in transmission through an on-chip backsideoptical window opened by a Deep Reactive Ion Etching (DRIE)process [8], and illuminated by a HeNe laser (633 nm/4 mW)with spot sizes ranging from 60 to 160 m to achieve tunablesensitivity. The output intensity signals are detected using a pho-toreceiver connected to an oscilloscope. Spatial filtering is pre-formed to minimize the cross-talk between diffraction ordersand therefore optimize the contrast. In earlier static penetrationmeasurements with switch open [8], we found an average pen-

1Drosophila embryos 50 min after hatching are dechlorinated in 60% bleachfor 1.5 min and then rinsed thoroughly with water (20 C) . Properly stagedembryos are selected and desiccated for 15 min in a sealed glass jar containingcalcium sulfate (CaSO ) desiccant. Finally, embryos are covered in Halocarbon700 oil (Aqua-Air Industries Inc., Harvey, LA) and ready for microinjection

etration force of % N and an embryo deformationof % m.

During vibratory microinjection experiments ( closed),the injector is translated toward the target while vibrating.The index-grating vibrates along with the injector and hencechanges its position relative to the fixed scale grating. The forceacting on the injector is determined by the relative displace-ment of the two gratings, which is determined by the intensitydistribution of the first diffraction order under observation.

The spring constant, , of a force encoder (beam width is8 m, beam thickness is 0.8 m, beam length is 1850 m,grating period is 10 m) was measured using the same tip cal-ibration setup as in Fig. 6. The measured value was 1.85 N/m

% , in reasonable agreement with the simulation result( N/m) of static displacement versus force shown inFig. 7. A three-dimensional meshed solid model (embedded)was generated using CoventorWare™ for the encoder. By fixing


Fig. 6. Experimental setup for operating the optical encoder force sensor in transmission mode on Drosophila embryos positioned on a 2-D self-assembly chip(shown in the embedded optical micrograph). The power in the first-order diffracted mode is measured with a photodetector.

Fig. 7. Simulated displacement versus injection force of an optical force encoder using CoventorWare™. The sensor behaves like a linear spring responding tothe force in the Y-direction only, and it is insensitive to lateral (X-direction) and out-of-plane (Z-direction) displacement. The encoder design parameters: indexgrating period 20 �m, grating thickness 0.8 �m, tip side-wall area 10 �m , supporting beam thickness 0.8 �m, beam length 1850 �m, and beam width 8 �m.Embedded graph shows the finite element 3-D meshed solid model used for the simulations. F and F are symmetrically applied to the two tip sidewalls. Theinjector displacement is mainly in the negative Y direction. The simulated spring constant in the Y-direction is 2.2 N/m.

the ends of the supporting springs, the microinjector displace-ment are calculated by simultaneously applying two increasingforce and orthogonal to each of the tip side walls inthe solid model. The total displacement, corresponding to

, is the summation of the individual displacement underand . The discrepancy between the simulated and the mea-

sured spring constant is assumed to be due to over-etching of thesprings supporting the movable upper grating.


Fig. 8. Modal analyzes of the optical force encoder. There are eight modes within the encoder operational frequency up to 16 kHz. Three low-order modes areshown: (a) the Z-direction vibration with resonant frequency of 1.4 kHz, (b) the rotation around X-axis at 4.6 kHz, and (c) Y-direction vibration at 13.9 kHz.These modes are critical for penetration force reduction, by providing significant injector tip translation while maintain the flatness of the gratings for readout. TheY-direction vibration is most desirable due to its maximum penetration into the object to be injected. In addition, the Y-direction motion is easily detected since theencoder is designed to be sensitive to translation in the Y-direction. Both the X-rotational mode and the Z-direction vibration will facilitate the penetration due tothe tip-tilt motion of the tip. The Z-direction vibration is the fundamental mode with resonant frequency 10 times lower than that of the Y-direction. This differenceis due to the 10:1 height-to-width ratio of the supporting beam cross-sections (beam width 8 �m, thickness 0.8 �m). However, the encoder sensitivity to z-directionlevitation is minimized through destructive interference by using gratings of proper thickness. So the Y-direction vibration is expected to be the dominant modefor force reduction.

The microoptical encoder can be treated as a linear spring-mass-dashpot system [19]–[21] . Therefore the dynamic prop-erties of the encoder can be studied through eigenfunction anal-ysis. We calculate the distribution of mass and stiffness from theencoder solid model, and solve the eigen-problem that yieldsfrequencies and mode shapes satisfying the force balance equa-tion. The modes of vibration are simulated using the Coven-torWare™ MemMech module. There are eight modes withinthe encoder operational frequency up to 16 kHz. Three loworder modes, the Z-direction vibration with resonant frequencyof 1.4 kHz, the rotation around X-axis at 4.6 kHz, and Y-di-rection vibration at 13.9 kHz, are critical for penetration forcereduction. These modes provide significant injector tip transla-tion while maintain the flatness of the gratings for readout. TheZ-direction vibration is the fundamental mode with a resonantfrequency 10 times lower than that of the Y-direction vibration.

This difference is due to the 10:1 height-to-width ratio of thesupporting beam cross-sections (beam width 8 m, thickness0.8 m). However, the encoder sensitivity to Z-direction lev-itation is minimized through destructive interference by usinggratings of proper thickness [8]. Likewise, the optical encoderis insensitive to X-axis rotation, so only Y-direction vibrationsare measured during penetration and vibration experiments.

The Y-direction resonant frequency is measured using thesame interferometric optical encoder used for the force calibra-tion. The incident laser beam is focused on the gratings, whilethe chip with the encoder is driven laterally by a thin slab ofpiezoelectric actuator. The output of the photo sensitive detectoris connected to an rms-to-dc converter and is observed versusdrive frequency to the piezoelectric actuator. The average mea-sured resonant frequency is 14 kHz, in good agreement with thesimulated 13.9 kHz Y-direction resonance.


Fig. 9. Experimental results for vibratory microinjection into Drosophila embryos. The injector was pushed into the embryo till a fixed bias point. For each biaspoint, the actuation voltage, and therefore the injector tip velocity, was increased (f = 14 kHz) till penetration occurred. A linear relationship is found betweenprobe average tip velocity at penetration and the probe bias position, showing the peak penetration force decreasing by �2:5 �N with every 0.1 m/s tip peakvelocity increase.

TABLE IPEAK PENETRATION FORCE, F , VERSUS BIAS POSITION d , ACTUATION

VOLTAGE, V , VIBRATION AMPLITUDE, d , AND TIP PEAK VELOCITY, v

We perform two sets of vibratory microinjection experiments.First the injector is brought into contact with the embryo to givea fixed bias offset of the two gratings, i.e., there is a fixed av-erage force acting on the injector. At each bias point, the actua-tion voltage, and therefore the injector tip velocity, is increasedtill penetration ( kHz). The critical peak-to-peak actu-ation voltage, , for penetration is used to calculate the probevibration amplitude, . During a full cycle of encodersinusoidal vibration , the tip would travel a total distance of

with the peak tip velocity .As shown in Fig. 9, with zero static force bias, the minimum

tip peak velocity needed for penetration is m/s (ac-tuated at kHz and V). This corresponds to apeak penetration force of 27.1 N, or a 48.5% reduction com-pared to the static penetration without vibration. Table I sum-marizes the calculated tip velocity and the associated peak pen-etration force at each fixed bias offset of the microinjector. Inover 200 vibratory microinjection experiments, we observed alinear decrease in the penetration force of N with every0.1 m/s tip peak velocity increase.

In the second set of experiments, the vibration amplitudewas kept constant and the injector was pushed into the embryountil penetration. Both peak and average penetration force were

measured. Fig. 10 shows the measured peak penetration forceas a function of actuation voltage and frequency. Each of the170 points on the surface represents an averaged peak penetra-tion force measured on three to five embryos. The read-out ofthe total linear injection-stage displacement upon penetrationis scaled with respect to the encoder pitch period to get thelinear displacement, , between the two gratings, while theamplitude of microinjector vibration is calculated based on theactuation frequency and voltage. The peak force can then bederived from the total relative displacement of the two gratings.

Both the average and the peak penetration force decrease asthe driving frequency and/or actuation voltage increases. At afixed actuation voltage, the penetration force reached a min-imum at the encoder resonant frequency of 14 kHz. The localminimum around 3.1 kHz is likely caused by the combination ofz-direction vibration and x-direction rotation (see Fig. 8). Bothmodes involve significant motion of the injector tip that can fa-cilitate penetration. At the maximum actuation voltage of 150 V,the minimum tip peak velocity for penetration was 1.5 m/s. Theminimum peak penetration force was 15.6 N, % of thestatic penetration force, while the minimum average penetrationforce was 2.7 N, % of the static penetration force.

We have recently reported an automated system forDrosophila injection for high-throughput RNAi screens [4],[7]. The system uses two CCD cameras to position an injectionneedle relative to the embryos. The microfabricated injectionneedle is integrated with piezoresistive pressure sensor en-abling control of the injected volume. With this set-up 100embryos were injected within 7 min. The first experimentswith double-stranded RNA injection proved successful, and theexpected genetic modification of the embryos was observed.Further reduction of penetration force through vibratory actua-tion will likely to enhance the throughput the RNAi screen.


Fig. 10. Peak penetration force measured over the frequency range f = 0 � 16 kHz with the actuation voltage, V , varying from 0 to 150 volts. The averageforce was calculated based on injections on multiple embryos under each of the 170 experimental conditions (f ;V ). The penetration force reached its minimumat the encoder resonant frequency of 14 kHz. The second dip around 3.1 kHz is likely to be caused by the combination of mode (a) and mode (b) as shown in Fig. 8through tip levitation.

V. CONCLUSION

Membrane-impermeable macromolecules such as peptides,proteins, oligonucleotides, DNA, RNA, and a variety of otherprobes can alter or assay cell function. A complementarymethod is to modulate the external cellular environment atwell-controlled spatial–temporal resolution [22], [23]. Amongavailable methods for introducing molecules into embryosor cells, such as chemical (ATP, EDTA), vehicular (erythro-cyte fusion, vesicle fusion), electrical (electroporation), andmechanical (microinjection, hyposmotic shock, sonication ormicroprojectiles), microinjection is the standard method forloading embryos and cells. It can reproducibly deliver largenumbers of macromolecules to most embryo and cell types withhigh viability and function. Minimally invasive micro injectionand surgical tools with integrated sensors are critical for a widerange of studies in biology and medicine, including calibratedtransmembrane delivery of genetic material into biologicalmodel systems, such as Drosophila embryos, by enabling highthroughput screening of gene functions.

We experimentally demonstrate an integrated optical-en-coder force sensor with configurable sensitivity and sufficientdynamic range for monitoring the linear penetration and injec-tion force in Drosophila embryos, and demonstrate significantforce reduction through dynamic tip-membrane interaction.For vibration-only operation, the force is reduced by 4.8% forevery 0.1 m/s velocity increase. When the motion of the forceprobe is a combination of vibration at the encoder resonanceand linear translation, the peak force is reduced by 70.3%,while the average force is reduced by 94.9% with respect to theforce required for linear translation alone. The integration ofsilicon optical sensors with the microinjectors allow for closedloop control of the vibratory actuation and characterization ofthe injection process.

ACKNOWLEDGMENT

The device fabrication and characterization were performedusing the National Nanofabrication Users Network (NNUN)facilities at the Center for Integrated systems and EdwardL. Ginzton Lab at Stanford University. The authors wish tothank I.-W. June for the help on the modal analysis, Dr. S.Zappe for providing the scanning electron micrograph (SEM)of Drosophila embryos, and M. Fish for preparing Drosophilaembryos for microinjection experiments.

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Xiaojing Zhang (M’04) received the Ph.D. degreein electrical engineering from Stanford University,Stanford, CA, in 2004.

His industrial experience includes working atHewlett-Packard on the design of parallel opticalinterconnects, and at Cisco Systems on the designand evaluation of microphotonic devices and opto-electronic subsystems. From 2004 to 2005, he wasa Research Scientist at Massachusetts Institute ofTechnology (MIT), Cambridge, and later, a VisitingResearch Associate at the University of Chicago, IL.

In 2005, he joined the faculty at The University of Texas at Austin. Currently, heis an Assistant Professor in the Department of Biomedical Engineering, and theUniversity of Texas Microelectronics Research Center. His research interestsare integration of photonics with microelectromechanical systems (MEMS)and microfluidic devices for in vivo imaging, biomanipulation, and nanoscalesensing. The actively pursued areas are: 1) Miniaturized silicon instrumentsfor in vivo cell and embryo manipulation and culturing; in particular, microin-jections, ultrasonic cellular-scale surgical tools, self-assembly and high-speedparticle sorting for studying cellular interactions and embryo developmentnetwork. 2) Nano-micro fabricated photonic sensors for characterizing the cellmechanics, and microscanners for molecular imaging and microscopy towardsminiaturized endoscopic precancer detection and diagnosis. 3) Multiscale sim-ulation of fundamental force, flow, energy processes involved in cell-substrateinteractions. Recent work includes theoretical and experimental studies on theenergy dissipation process associated with fluidic self-assembly for both softbiological samples and silicon chips.

Matthew P. Scott received the B.S. and Ph.D. de-grees in biology from the Massachusetts Institute ofTechnology (MIT), Cambridge.

He did postdoctoral research at Indiana Universityand then joined the faculty at the University ofColorado at Boulder. In 1983, he moved to StanfordUniversity School of Medicine, where he is nowProfessor of Developmental Biology and of Ge-netics. He has published more than 130 papers andthree patents. His research areas are developmentalgenetics and cancer research, particularly the roles

of signaling systems and transcriptional regulation in embryonic development.His research employs genetics, genomics, cell biology, and molecular biologyin exploring how cells acquire their fates and are patterned. He is an editorof Current Opinion in Genetics and Development and of the Proceedings ofthe National Academy of Sciences. He is a past president of the Society forDevelopmental Biology, a member of the American Academy of Arts andSciences, and a member of the National Academy of Sciences. He is presentlychairing Stanford’s Bio-X program, which is designed to accelerate the comingtogether of engineering, physics, and chemistry with biology and medicine.

Calvin F. Quate (S’43–A’50–M’55–F’65–LF’89)received the B.S. degree from University of Utah in1943 and the Ph.D. degree from Stanford University,Stanford, CA, in 1950.

He held positions at Bell Laboratories and SandiaCorporation before joining the faculty at StanfordUniversity, where he has been since 1961. He iscurrently the Leland T. Edwards Professor of elec-trical engineering and applied physics at StanfordUniversity. His research focuses on the developmentand application of scanning probe microscopes,

MEMS and Nanotechnologies.Dr. Quate is a Member of the National Academy of Engineering and the Na-

tional Academy of Science, an Hororary Fellow of the Royal Microscopical So-ciety, and a Foreign Member of the Royal Society London. He has received theIEEE Morris N. Liebmann Award in 1981, the Rank Prize for Opto-electronicsin 1982, the IEEE Medal of Honor in 1988, the President’s National Medal ofScience in 1992, and the American Physical Society Keithley Award in 2000.

Olav Solgaard (M’90) received the B.S. degree inelectrical engineering from the Norwegian Instituteof Technology and the M.S. and Ph.D. degrees inelectrical engineering from Stanford University, CA.

He held a postdoctoral position at the Universityof California at Berkeley, and was an Assistant Pro-fessor at the University of California at Davis, beforejoining the faculty of the Department of ElectricalEngineering at Stanford University in 1999. His re-search interests are optical communication and mea-surements with an emphasis on semiconductor fabri-

cation and MEMS technology applied to optical devices and systems. He hasauthored more than 150 technical publications and holds 18 patents. He is a co-founder of Silicon Light Machines, Sunnyvale, CA, and an active consultant inthe MEMS industry.

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