IOP PUBLISHING JOURNAL OF MICROMECHANICS AND MICROENGINEERING

J. Micromech. Microeng. 18 (2008) 075020 (8pp) doi:10.1088/0960-1317/18/7/075020

A polymer V-shaped electrothermalactuator array for biological applicationsWenyue Zhang1, Markus Gnerlich1, Jonathan J Paly2, Yaohua Sun1,Gaoshan Jing1, Arkady Voloshin2,3 and Svetlana Tatic-Lucic1,2

1 Sherman Fairchild Center, Electrical & Computer Engineering Department, Lehigh University,Bethlehem, PA 18015, USA2 Bioengineering Program, Lehigh University, Bethlehem, PA 18015, USA3 Department of Mechanical Engineering & Mechanics, Lehigh University, Bethlehem, PA 18015, USA

E-mail: svt2@lehigh.edu

Received 27 March 2008, in final form 8 May 2008Published 9 June 2008Online at stacks.iop.org/JMM/18/075020

AbstractA polymer V-shaped electrothermal actuator (ETA) array that is capable of compressing a livebiological cell with a desired strain was designed, fabricated and characterized. This polymerelectrothermal array is the core of a microelectromechanical systems (MEMS) device tomeasure the mechanical compliance of a cell. A polymer electrothermal actuation mechanismwas selected because it is able to operate in an electrolytic solution (cell medium), which wasneeded to keep cells alive during testing. The MEMS-based device was optimized utilizingfinite element analysis and the devices were fabricated using surface micromachiningtechniques. Characterization of these devices was conducted in air, deionized water and cellmediums. Operating these devices in liquid environments was performed using direct currentvoltages less than 2.0 V or high-frequency (800 kHz) alternating current voltages. Theactuator displacement was up to 9 µm in air and 3 µm in liquids, i.e. it achieves 30%displacement of that in air when operating in liquids. Such remarkable performance is due tothe large coefficient of thermal expansion and low thermal conductivity of the structuralpolymer (SU-8). Finally, we demonstrated the suitability of this actuator for biologicalapplications by compressing a cultured NIH3T3 fibroblast in the cell medium.

(Some figures in this article are in colour only in the electronic version)

1. Introduction

A study of the mechanical compliance of biological cells iscritical to improve research to benefit public health becausemeasurements of the compliance contribute to the studyof the pathophysiology of various diseases and the searchfor effective treatments. Biologists hypothesize that thebiomechanical properties of osteoblasts (bone formation cells)change as a function of age, and this change could be acontributing factor to the pathogenesis of osteoporosis [1].This hypothesis of the osteoblasts’ mechanosensitivity has notbeen examined due to the limitations of current measurementtechniques.

Biomechanical research could be traced back to the 19thcentury where it started from a gross anatomical level ofinvestigation and has advanced to the cellular level today. In

early research, strain gages were attached to the midshaft oflive animal bone to record the variations of strains duringthe animal’s natural activities in vivo for 1–2 days [2].Later, accelerometers were employed to obtain quantitativevalues necessary to evaluate the shock-absorbing capacityof the human locomotion system [3]. Recently, mechanicalproperties of live cells have been investigated by atomic forcemicroscopy (AFM) [4], soft substrate stretching [5], magneticbeads attachment [6], cytoindentation [7] and modified tensiletesting [8], among other techniques.

AFM has broad applications in characterizing theelasticity of biological materials [9]. It has been effective,not only for imaging the morphology of developing neuronsand their processes three-dimensionally, but also for studyingthe elastic properties of a live osteoblast on a submicrometerscale [10]. However, AFM has several drawbacks when used

0960-1317/08/075020+08$30.00 1 © 2008 IOP Publishing Ltd Printed in the UK

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

Cell positioning system (Dielectrophoresic electrodes)

Cell position

Thermal sensor

V-shaped electrothermal actuator array

Force sensor

µ

Figure 1. The fabricated BioMEMS device for measuring themechanical compliance of a biological cell.

in biomechanics studies, such as small deformations (typicaldeformation is around 100 nm [10]) and the ability to performonly local mechanical stimulation of the cell.

Microelectromechanical systems (MEMS) technologycan integrate both microsensors and microactuators in a singlechip, which is beneficial to most applications, including themeasurement of cells’ mechanical properties. A MEMS-basedsilicon sensor has been used to monitor indenting forces [11].Also, a comb-drive electrostatic actuator has been built tostimulate cells mechanically [12], and this actuator was basedon a frequency-dependent electrostatic actuation mechanismwhich could operate in ionic cell media. By applyingMEMS techniques to biological applications, measurementscan be made less time consuming and costly, so that themeasurement of cells’ mechanical properties could be morebroadly available. We are utilizing the full advantages ofMEMS techniques in the device described below.

2. Overview of the BioMEMS device for measuringbiological cells’ compliance

A novel technique for measuring the mechanical complianceof biological cells utilizing MEMS technology as aplatform has been developed [13]. Each BioMEMS device(figure 1) has four functional modules: (1) a polymer V-shapedelectrothermal actuator (ETA) array that provides a desiredstrain to a cell; (2) a force sensor that measures the magnitudeof the force applied to a cell; (3) a cell-positioning systemthat traps a cell precisely between the plunger and the forcesensor using dielectrophoresic quadrupole electrodes and (4) athermal sensor that monitors the temperature elevation on-chipduring experiments. The core of this BioMEMS device is apolymer V-shaped in-plane ETA array. Each side of the flat V-shape expands along its length when heated, and the expansionproduces a forward movement at the tip. The design intentionof the V-shaped electrothermal array is to combine movementof many actuators into a stable forward motion with higherforce than just one actuator could achieve.

3. Design and modeling of the BioMEMS actuatorarray

Operation in a conductive liquid was one of the mainchallenges in this BioMEMS actuator design. Several MEMSactuators have demonstrated actuation in liquid environments,such as electrostatic actuators (frequency dependent) [10],electrochemical actuators [14] and ETAs [15]. Among them, apolymer-based ETA with hot and cold arms had been applied ina biological application successfully [16]. During underwateroperation, ETAs need to overcome difficulties of electrolysis,thermal losses, electrical losses and surface tension [17].Among them, electrolysis and thermal loss have the biggestimpacts. Electrolysis stops high electrical dc power frombeing applied to the actuator because it can produce largeamounts of gases in liquid due to the breakdown of water intohydrogen and oxygen when the dc voltage exceeds 1.23 V[18]. Meanwhile, thermal loss is due to the higher thermalconductivity in water (0.598 W mK−1), which is approximately20 times larger than that of air (0.0263 W mK−1) [17]. Thismeans a great amount of heat is transferred through thesurrounding liquid, instead of heating up the thermal actuator.In summary, thermal losses require that more electrical poweris needed to operate the thermal actuator underwater, butelectrolysis prevents higher power from being input to theactuator.

A V-shaped ETA structure was selected because it hasbetter underwater performance [15] and provides larger forcesthan hot/cold arm structures can [19]. A polymer (SU-8,Microchem Inc., MA) was chosen to be the structural materialprimarily due to its large coefficient of thermal expansion(CTE) (52 ppm ◦C−1), low thermal conductivity (0.3 WmK−1) [20] and good biocompatibility [16]. The high CTEand low thermal conductivity reduce the necessary actuationtemperature and help isolate the hot areas of the actuator fromthe biological cell. These thermal properties allow a MEMSdevice to be built which has a low risk of thermally damagingdelicate biological cells.

The basic unit structure of a single V-shaped ETA wasa suspended polymer beam with two ends anchored on asubstrate (figure 2). Because the polymer layer is notelectrically conductive, a metal heater was deposited on topto create a current loop when a voltage was applied throughthe Vin and Vout pads. If the polymer beam were oriented atan angle (θ ) to the array yoke, it would move in the directionof the tip after resistive heating [21]. An array consisting offour single V-shaped ETAs connected by a yoke was designedto achieve larger force output by coupling forces from singleETAs [22]. At the end of the yoke, a plunger was used tocontact a cell directly.

To optimize the actuator’s geometry, finite elementanalysis (FEA) was performed using the CoventorwareTM

MemMech simulator. The device was a four-beam V-shapedarray made of SU-8, whose Young’s modulus was 4.95 GPa[23]. The design specification, which was to achieve upto 2–4 µm in-plane displacement, was confirmed by theFEA. Out-of-plane displacement was clearly predicted dueto the mismatches in the CTE between the structural polymer

2

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

W

Polymer Beam

Vout

Vin

Metal Heater

L

Anchor

Plunger Array yoke

Single V-shaped ETA Array

θ

Figure 2. Schematic of the meshed model of an array of V-shaped electrothermal actuators (vertical dimension magnified 10× for bettervisibility).

(52 ppm ◦C−1) and gold (14.3 ppm ◦C−1). The FEA showedthat this negative impact can be ignored when the device isoperated at low power (less than 100 mW). The resolutions offorce and displacement sensors were designed to be 200 nNand 500 µm, respectively.

4. Fabrication of the BioMEMS device

Using surface micromachining techniques, the BioMEMSdevice was fabricated starting with 3 in diameter siliconwafers. The first composite metal layer (20 nm titaniumand 80 nm platinum) was deposited on top of the oxidizedsilicon, where the titanium was used to increase adhesion ofplatinum to the substrate (figure 3(a)). Afterward, a sacrificiallayer consisting of 5 µm thick AZ P4620 positive photoresist(Clariant Inc., NJ) was spin-coated and patterned (figure3(b)). Before further processing, the photoresist residue onthe wafer surface was removed in a UV/ozone oven (modelT10 × 10/OES, UVOCS Inc., Montgomeryville, PA). Anotherpurpose of this step was to harden the photoresist patterns inorder to resist physical and chemical damages introduced inthe following steps [24]. Next, the structural polymer (15 µmthick SU-8 negative photoresist, MicroChem Inc., MA) wasspin-coated and exposed with a conventional UV aligner (KarlSuss MJB3, figure 3(c)). Finally, lift-off photolithography wasused to pattern the second metal composite layer (20 nm oftitanium and 80 nm of gold), which adhered to the top of thestructural polymer layer (figure 3(d)). Finally, the sacrificiallayer was dissolved in an AZ 400 T photoresist stripper(Clariant Inc., NJ) at room temperature to release the structure(figure 3(e)).

5. Cell preparation

Mouse fibroblast cells (NIH3T3) were selected for thisexperiment instead of osteoblasts due to their immediateavailability. NIH3T3 fibroblast cells were cultured in25 cm2 flasks (Fisher Scientific, GA) at 37 ◦C in anincubator with humidified 5% CO2. The culture mediumcontained Dulbecco’s Modified Eagle’s Medium (DMEM)(with 4.5 g L−1 glucose), 2 mM glutamine, 7.5% sodiumbiacarbonate, penicillin/streptomycin antibiotic, fungizoneand 10% newborn calf serum.

The NIH3T3 cells were dissociated by incubating in a0.125% (w/v) trypsin solution at 37◦ C for 10 min. Aftertrypsinization, the cells were pelleted by centrifugation at750 rpm for 5 min and were resuspended in the culturemedium. The cells were seeded at a density of 1 × 105

cells in a flask that contained 5 ml cell medium, and werefed every 36–48 h for 7 days before the next dissociation.After resuspending, the cells were transferred to the devicesfor testing using a pipette.

6. Characterization of the BioMEMS device

For the electrical characterization of the performance ofthe BioMEMS device, a custom-designed amplifier circuitwas built using an Apex PA19 operation amplifier (ApexMicrotechonology Corp., Tuson, AZ). Measurements weremade using a Suss MicroTec PM5 probe station, anddisplacement was recorded using a 3.0 megapixel CMOScamera (Moticam 2300, Motic Inc., China). Afterward,the Image Plus v2.0 ML image processing software (MoticInc., China) was used to measure the displacement fromthe captured digital images, and the optical system wascalibrated in air using a standard calibration slide (providedby Motic Inc.) so that the digital image was 150 nm perpixel. One image was taken when applying a controlledpower to the device and four measurements were extractedfrom the image. The error bars in the following figuresrepresent the standard deviation of the multiple measurementsin one image. The same measurement method was appliedto other figures in this paper unless specified differently,and the available optical resolution was ±0.3 µm. Thepolymer V-shaped ETA arrays were characterized electricallyin three different environments: air, deionized (DI) waterand cell medium. The inputs were low dc voltage(less than 2 V) or high-frequency (800 kHz) ac voltage. Themotivation for applying ac voltage was to avoid electrolysis.To avoid heating the cell medium surrounding the thermalactuator, each actuation voltage was applied as a 2.5 s pulseengaged by a function generator and relay.

Prior to the characterization tests in cell media, thepolymer device was submerged in the cell media for 2 h inorder to allow the water content in the polymer to stabilize.The device characterization experiments were conductedwithin 30 min at 23 ◦C ambient temperature. In order to

3

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

Substra

(a) (b) (c)

(d ) (e) ( f )

te (Oxidized silicon) 1st Metal Layers (Ti/Pt) Sacrificial Layer (AZ P4620) Structural Layer (SU-8) 2nd Metal Layers (Ti/Au)

Figure 3. Illustration of the fabrication flow for the BioMEMS devices to measure cell mechanical compliance: (a) deposit and pattern thefirst metal layers (Pt/Ti) on an oxidized silicon wafer, (b) deposit and pattern the sacrificial layer (AZ P4620), (c) deposit and pattern thestructural layer (SU-8), (d) deposit and pattern the second metal layers (Au/Ti), (e) release the structural layer by removing the sacrificiallayer, where (f) is the material color legend.

0

1

2

3

4

5

6

7

8

9

10

0 100 200 300 400 500

Electrical DC Power (mW)

Dis

pla

cem

ent

(µm

)

Figure 4. Experimental results of the displacement as a function of the input electrical power when the V-shaped electrothermal array wasactuated by dc voltage in air. The displacements were measured optically from top view (400× magnification).

quantify the electrical losses in conductive liquids, electricalconductivities of DI water and cell media (NIH3T3 fibroblastmedia) were tested using OAKTON R© CON6/TDS6 Hand-held Conductivity/TDS Meter at room temperature. Themeasured electrical conductivities of DI water and cell media(NIH3T3 fibroblast media + 10% newborn bovine serum) were1.48 µS cm−1 and 15.85 mS cm−1, respectively.

7. Results and discussion

7.1. Polymer V-shaped ETA actuation by dc voltage in air

Initially, the electrical characterization of the polymer V-shaped ETA array was conducted in an air environment. Thearray was made of four V-shaped polymer (SU-8) beams,and each beam was 200 µm long, 50 µm wide, 8 µm

thick and had an angle θ of 35◦. The experimental dataof the electrothermal array displacements as a function ofinput electrical dc power were obtained (figure 4), and theyconfirmed that these polymer V-shaped electrothermal arrayscould achieve up to 9 µm displacement with 365 mW of power.

The displacement drop in the experimental data relativeto applied power was associated with significant out-of-plane displacement. The yoke curled upward and this reducedthe in-plane displacement, which was predicted earlier duringthe finite element analysis.

7.2. Polymer V-shaped ETA actuation by dc voltage in liquidenvironments

The same actuation experiments were conducted in DIwater and cell medium (NIH3T3 fibroblast media) and the

4

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

DC Electrical Power (mW)

Dis

pla

cem

ent

(µm

)

0

1

2

3

4

5

6

7

8

9

10

0 50 100 150 200 250 300 350 400

In air

In DI water

In cell medium

Figure 5. Comparison of measured displacement of the electrothermal array as a function of the input dc power when operating in threedifferent media.

AC Power RMS (mW)

Dis

pla

cem

ent

(µm

)

0

1

2

3

4

5

6

7

8

9

10

0 200 400 600 800 1000

In airIn DI waterIn cell medium

Figure 6. Measured displacement of the electrothermal array as afunction of the input ac power.

displacement as functions of input power in different mediawere compared (figure 5). The maximum displacementswere 3 µm in liquids, which was one-third of the maximumdisplacement achieved in air (9 µm). The degradationwas primarily due to thermal loss through the surroundingliquid. Furthermore, the statistic analysis results showed nosignificant difference between the displacement in DI waterand in cell media. This was partially due to the fact thatelectrical loss through weakly conductive cell media is alimited portion of applied power—the parasitic power lossthrough the cell media was found to be 7% of the total power[13].

Compared with polysilicon ETAs, our experimental datashow that the polymer ETAs improved thermal efficiency

by five times. When operating polysilicon V-shaped ETAsin water, the polysilicon actuators deflect 6% of theirdisplacement per volt applied in air because the thermalconductivity of water is 20 times higher than that of air [15].In contrast, our experimental results indicate that the polymerETAs’ displacement was above 30% of the displacement in theair at the same electrical power. This result reveals that thesepolymer ETAs convert more thermal energy to mechanicalmovement. This increased efficiency of thermal energy usagewas a result of the low thermal conductivity of the polymer(0.3 W mK−1 versus 13.8 W mK−1 for polysilicon), so moreheat was confined inside the devices. This advantage made thepolymer electrothermal array more suitable for the operationin liquid environments.

7.3. Polymer V-shaped ETA actuation by ac voltage indifferent media

DC actuation is not suitable for underwater actuator operationdue to electrolysis. Therefore, high-frequency ac voltage(sinusoidal wave, 800 kHz, 1 Vp-p to 40 Vp-p) was applied tothe designed circuit. The high ac frequency makes the polarityof electrical field changes faster than the ions screening theelectrodes, and electrolysis is avoided [10]. The deviceswere actuated in three different media: air, DI water andcell medium. The displacement of the electrothermal arrayas a function of an input ac power RMS value was recorded(figure 6). The maximum displacements were 8 µm, 6 µm and4 µm when these devices actuated by ac voltage in air, DI waterand cell medium, respectively. The reduction of incrementalin-plane displacement at high power levels indicated that thedevices were near the maximum of their mechanical designlimits due to out-of-plane deformation.

The displacement as a function of input powerwhen applying either dc or ac voltages was compared

5

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

0

1

2

3

4

5

0 200 400 600 800

Power (mW)

Dis

pla

cem

ent

(µm

)

AC/In cell mediu mDC/In cell medium

Figure 7. Comparison of the displacement as a function of inputelectrical power when either a dc or ac voltage was applied. Thesedevices were operated in cell medium.

(figure 7). The maximum displacements were 4 µm by acactuation and 3 µm by dc actuation in cell media. Eventhough both of them met the design specification (2 µmdisplacement), the electrothermal array could move further

(a) (b)

(c) (d )

Out of plane displacement observed

Figure 8. Displacement tests in air for time-dependent behavior of a polymer electrothermal array showing displacement versus the elapsedtime (top) and the corresponding applied power (bottom). The results for a low-power test are shown in (a) and (c), while the results for ahigh-power test are shown in (b) and (d).

under ac voltage actuation than under dc voltage actuation.More power could be applied to the device using ac voltage,and the electrolysis problem could be avoided as well. Thesedevices were operated in cell medium (electrical conductivity15.85 mS cm−1).

7.4. The viscoelastic behaviour of the polymer V-shapedelectrothermal array

In addition to performance tests in various environments,the viscoelastic behavior of the polymer electrothermal arraywas tested by applying a series of 20 s long pulses ofvoltage at increasing power levels, and optically measuringthe displacement in air. Unlike the previous setup, thesetests used automated image analysis and a feature detectionalgorithm in MATLAB to determine displacement whichallowed a large number of time-lapse images to be processed.The displacement resolution was ±0.3 µm and the powermeasurement resolution was 100 µW.

The tests for a six-beam electrothermal array show that atlow power levels up to 127 mW (figures 8(a) and (c)), thereis forward movement with a residual negative displacementwhich slowly relaxes to the original position. However, athigher power levels (figures 8(b) and (d)) out-of-plane plasticdeformation was observed which occurred when the power

6

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

(a) (b)

Actuator

Cell

DEP Electrodes Actuator

Cell

DEP Electrodes

16 µm 12 µm

Figure 9. Optical images of (a) a cell being trapped at the center of dielectrophoresic quadrupole electrodes, and (b) the cell beingcompressed by the actuator after being powered up (the outline of the cell is enhanced for better visibility).

level reached 270 mW (but was not present at a power level of230 mW) and which remained present even after 48 h. Thisplastic deformation was manifested as a slight upward curl ofthe pusher tip above the substrate surface, and was predicted bythe FEA simulations discussed above. Note that the residualnegative displacement is limited to 1 µm (figure 8(b)) evenas the forward displacement continues to increase. The time-dependent deformation in figure 8(a) is an interesting propertyof the polymer actuator, but it does not impact performance forits use in the cell compression system, since the displacementis measured independently (either by optical means or by anintegrated displacement sensor planned for future devices).

7.5. Compression of a cell

In order to demonstrate the electrothermal array capabilitiesfor the target biological application, the entire device wasimmersed into cell medium and NIH3T3 fibroblast cells weretransferred to the cell medium surrounding the BioMEMSdevice using a micropipette. The diameter of the fibroblastcell under test was measured to be 16 µm. First, one cellwas trapped at the center of the dielectrophoresic quadrupoleelectrodes after applying ac voltage (1 MHz, 10 Vp-p,sinusoidal) to them (figure 9(a)). The cell and the devicewere immersed in the cell media for 2 h in order toestablish the initial contacts with polymer beams, and thecell shown in figure 9 was a particularly close match tothe gap width in front of the actuator (gap width was avariable parameter in the design space, due to naturallyoccurring variance in cell diameters). Next, the actuatorwas powered using ac voltage, and the cell was compressed(Figure 9(b)). The maximum displacement was 4 µm.Currently, the displacement resolution in cell medium was±0.3 µm when a cell was present; nevertheless, it was clearfrom this demonstration that the maximum strain of 25% wasachieved, while 10% strain was the requirement [25]. Thecell could be compressed repeatedly and did not disintegrate

during or after the compression test. This confirmed thatthese BioMEMS devices could provide sufficient mechanicalstimulation to the live biological cell so that its compliancecould be measured.

Future measurements of cell mechanical compliancewould involve a force sensor which would be used to measurereaction force during the cell compression test. During theenvisioned testing method, a very small preloading strainwould be applied to the cell by the actuator which wouldmark the starting point of the compression, and the integratedforce sensor would be used to detect this point of contact. Inthis experiment, the change in cell diameter was measuredoptically without consideration for this detail, which is thesubject of ongoing research.

8. Conclusion

In order to measure the mechanical compliance of livebiological cells, a polymer-based MEMS device that integratedan ETA array, a cell-positioning system, a force sensor and athermal sensor on a single chip has been built. This researchaimed to develop a novel technique to measure the compliancemore rapidly and cost effectively.

This paper focuses on the core of this BioMEMS device:a polymer V-shaped in-plane ETA array. The electrothermalactuation mechanism was selected due to superior underwateroperation. A solid model of the MEMS device was createdand finite element analysis of the actuator displacement asa function of increasing input voltage was determined usingthe CoventorwareTM MemMech solver. The finite elementanalysis results predicted that this device design could meetthe design specifications primarily due to the large CTE andlow thermal conductivity of a polymer.

The MEMS devices were fabricated using surfacemicromachining techniques, which includedphotolithography, metallization, lift-off and sacrificiallayer release. The fabrication process was dictated by the

7

J. Micromech. Microeng. 18 (2008) 075020 W Zhang et al

tight thermal budget of polymer materials, which was themain limitation of processing integration. Characterization ofthe device was conducted in air, DI water and cell media. Theexperimental data revealed that these BioETAs had five timeshigher thermal efficiency than their polysilicon counterpartsdocumented in the recent literature [15]. They demonstratedup to 4 µm displacement in cell media (NIH3T3 fibroblastmedia), where the displacement is the function of the inputelectrical power. The BioMEMS device described hereprovided 25% mechanical strain to a live NIH3T3 fibroblastcell, which exceeds the design requirement.

This actuator can be applied for compression of differenttypes of biological cells. It is a key component in a BioMEMSsystem for measuring time-independent mechanical propertiesof cells [24]. These measurements are useful to studymechanotransduction, cell adhesion and other phenomena ofinterest in tissue engineering.

Acknowledgments

This project is supported by National Aeronautics and SpaceAdministration (NASA), grant no. NNX06AD01A. Thefabrication was conducted at Sherman Fairchild Laboratory,Lehigh University. The authors are grateful to ApexMicrotechnology Corporation, Tucson, AZ, for providingthe operational amplifier, and to Professor Susan F Perry,Dr Weisong Wang, Dr Shawn Cunningham from WiSpry,Inc., and Mr Subham Sett from SIMULIA for valuablediscussions.

References

[1] Turner C H, Takano Y and Owan I 1995 Aging changesmechanical loading thresholds for bone formation in ratsJ. Bone Miner. Res. 10 1544–9

[2] Fritton S P, McLeod K J and Rubin C T 2000 Quantifying thestrain history of bone: spatial uniformity and self-similarityof low-magnitude strains J. Biomechan. 33 317–25

[3] Voloshin A, Wosk J and Brull M 1981 Force wavetransmission through the human locomotor system J. Trans.Biomed. Eng. 103 48–50

[4] Vinckier A and Semenza G 1998 Measuring elasticity ofbiological materials by atomic force microscopyFEBS Lett. 430 12–6

[5] Brown T D 2000 Techniques for mechanical stimulation ofcells in vitro: a review J. Biomechan. 33 3–14

[6] Wang N, Butler J P and Ingber D E 1993 Mechanotransductionacross the cell surface and through the cytoskeletonScience 260 1124–7

[7] Shin D and Athanasiou K 1999 Cytoindentation for obtainingcell biomechanical properties J. Orthop. Res.17 880–90

[8] Miyazaki H, Hasegawa Y and Hayashi K 2000 A newlydesigned tensile tester for cells and its application forfibroblasts J. Biomechan. 33 97–104

[9] McNally H A and Borgens R B 2004 Three-dimensionalimaging of living and dying neurons with atomic forcemicroscopy J. Neurocytol. 33 251–8

[10] Domke J, Dannohl S, Parak W J, Muller O, Aicher W K andRadmacher M 2000 Substrate dependent differences inmorphology and elasticity of living osteoblasts investigatedby atomic force microscopy Colloids Surf. 19 367–79

[11] Yang S and Saif T 2006 MEMS based sensors for the study ofindentation response of single cells Tech. Dig. of MEMS2006 (Istanbul, Turkey, January 2006) pp 20–3

[12] Mukundan V and Pruitt B L 2006 Experimentalcharacterization of frequency dependent electrostaticactuator for aqueous media Tech. Dig. of Solid-StateSensors, Actuators, and Microsystems Workshop(Hilton Head Island, SC, USA, 4–8 June 2006)pp 167–71

[13] Zhang W-Y 2007 Design, modeling, fabrication andcharacterization of a MEMS device for measuring themechanical compliance of a biological cell PhDDissertation Lehigh University

[14] Smela E, Kallenbach M and Holdenried J 1999Electrochemically driven polypyrrole bilayers for movingand positioning bulk micromachined silicon plates J.Microelectromech. Syst. 8 373–83

[15] Sameoto D, Hubbard T and Kujath M 2004 Operation ofelectrothermal and electrostatic MUMPs microactuatorsunderwater J. Micromech. Microeng. 14 1359–66

[16] Chronis N and Lee L P 2005 Electrothermally activated SU-8microgripper for single cell manipulation in solution J.Microelectromech. Syst. 14 857–63

[17] Hickey R, Sameoto D, Hubbard T and Kujath M 2003 Timeand frequency response of two-arm micromachined thermalactuators J. Micromech. Microeng. 13 40–6

[18] Selvaganapathy P, Leung Ki Y-S, Renaud P andMastrangelo C H 2002 Bubble-free electrokinetic pumpingJ. Microelectromech. Syst. 11 448–53

[19] Lai Y, McDonald J, Kujath M and Hubbard T 2004 Force,deflection and power measurements of toggledmicrothermal actuators J. Micromech. Microeng. 14 49–56

[20] MicroChem Corp. 2007 Product data sheet of SU-8 200–2015series http://www.microchem.com/products/pdf/SU-82000DataSheet2000 5thru2015Ver4.pdf

[21] Que L, Park J-S and Gianchandani Y B 2001 Bent-beamelectrothermal actuators: Part I. Single beam and cascadeddevices J. Microelectromech. Syst. 10 247–61

[22] Reid J R, Bright V M and Comtois J H 1996 Forcemeasurements of polysilicon thermal micro-actuators Proc.SPIE 2882 296–306

[23] Dellmann L, Roth S, Beuret C, Racine G-A, Lorenz H,Despont M, Renaud P, Vettiger P and de Rooij N F 1998Fabrication process of high respect ratio elastic and SU-8structure for piezoelectric motor applications SensorsActuators A 70 42–47

[24] Zhang W-Y, Gnerlich M, Sun Y, Jing G, Paly J, Voloshin Aand Tatic-Lucic S 2008 Polymer MEMS system formeasuring the mechanical modulus of a biological cellProc. Int. Conf. on Biomedical Electronics and Devices(BIODEVICES 2008), (Funchal, Madeira, Portugal,28–31 January) vol 2 146–50

[25] You L, Cowin S C, Schaffler M B and Weinbaum S 2001 Amodel for strain amplification in the actin cytoskeleton ofosteocytes due to fluid drag on pericellular matrix J.Biomechanics. 34 1375–86

8