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Biosensors and Bioelectronics 33 (2012) 1– 9

Contents lists available at SciVerse ScienceDirect

Biosensors and Bioelectronics

j our na l ho me page: www.elsev ier .com/ locate /b ios

eview

coustic wave based MEMS devices for biosensing applications

oana Voiculescua, Anis Nurashikin Nordinb,∗

City College of New York, New York, NY 10031, USAInternational Islamic University Malaysia, Jalan Gombak, Kuala Lumpur, 53100, Malaysia

r t i c l e i n f o

rticle history:eceived 21 September 2011eceived in revised form 9 December 2011ccepted 21 December 2011vailable online 16 January 2012

eywords:EMS

a b s t r a c t

This paper presents a review of acoustic-wave based MEMS devices that offer a promising technologyplatform for the development of sensitive, portable, real-time biosensors. MEMS fabrication of acousticwave based biosensors enables device miniaturization, power consumption reduction and integrationwith electronic circuits. For biological applications, the biosensors are integrated in a microfluidic systemand the sensing area is coated with a biospecific layer. When a bioanalyte interacts with the sensing layer,mass and viscosity variations of the biospecific layer can be detected by monitoring changes in the acousticwave properties such as velocity, attenuation, resonant frequency and delay time. Few types of acoustic

ilm bulk acoustic wave resonators (FBAR)urface acoustic waves (SAW) resonatorAW delay lineass sensitivity

iosensor

wave devices could be integrated in microfluidic systems without significant degradation of the qualityfactor. The acoustic wave based MEMS devices reported in the literature as biosensors and presented inthis review are film bulk acoustic wave resonators (FBAR), surface acoustic waves (SAW) resonators andSAW delay lines. Different approaches to the realization of FBARs, SAW resonators and SAW delay linesfor various biochemical applications are presented. Methods of integration of the acoustic wave MEMSdevices in the microfluidic systems and functionalization strategies will be also discussed.

© 2011 Elsevier B.V. All rights reserved.

ontents

1. Introduction . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 12. Acoustic wave based MEMS devices as biosensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 2

2.1. FBAR devices used for biological applications . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 32.2. SAW delay lines as biosensors . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 52.3. SAW resonators as biosensors. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 5

3. Integration of the acoustic wave MEMS devices in microfluidic systems and sensor functionalization for biological tests . . . . . . . . . . . . . . . . . . . . . . . 63.1. Integration in microfluidic systems . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 6

3.2. Sensor functionalization . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7

4. Conclusions . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 7 . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8. . . . .

hrough or on top of a piezoelectric medium. The travelling wavesre sensitive to any change both on the surface and in the piezo-lectric material. Variations such as mass loading, viscosity thatccurs in the propagation path of the waves causes its velocity and

∗ Corresponding author. Tel.: +60 3 6196 4478.E-mail address: anis.nordin@gmail.com (A.N. Nordin).

956-5663/$ – see front matter © 2011 Elsevier B.V. All rights reserved.oi:10.1016/j.bios.2011.12.041

. . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . 8

amplitude to change. Traditional usage of acoustic wave deviceshave been in the telecommunications industry, primarily in mobile

cell phones and base stations (Hashimoto, 2000; Macchiarellaand Stracca, 1982). Emerging applications for acoustic wave devicesas sensors include as torque and tire pressure sensors (Cullen andMontress, 1980; Cullen and Reeder, 1975; Ivanov et al., 1996; Pohlet al., 1997), gas sensors (Levit et al., 2002; Nakamoto et al., 1996;Staples, 1999; Wohltjen and Dessy, 1979), biosensors for medi-cal applications (Andle and Vetelino, 1995; Ballantine et al., 1997;

Acknowledgement . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .References . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . . .

1. Introduction

Electro-acoustic sensors are popular in a wide range of applica-tions due to their high sensitivity and wireless capabilities. Thesesensors have high frequency (MHz–GHz) acoustic waves travelling

Cavic et al., 1999; Janshoff et al., 2000), and industrial and com-mercial applications such as: vapor, humidity, temperature, andmass sensors (Bowers et al., 1991; Cheeke et al., 1996; Smith, 2001;Smith et al., 2003; Vellekoop and Jakoby, 1999; Vetelino et al.,

2 sors and Bioelectronics 33 (2012) 1– 9

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Fig. 1. SAW delay line biosensor integrated in a microfluidic channel. The surfacebetween the IDTs is coated with antibodies sensitive to the analyte to be detected.

I. Voiculescu, A.N. Nordin / Biosen

996; Weld et al., 1999). Additional capabilities of acoustic waveensors include remote operation and passive interrogation (Guhrt al., 2005; Reindl et al., 1996). In recent years, the interest in theevelopment of highly sensitive acoustic wave devices as biosensorlatforms has grown. Biosensors are analytical devices that com-ine a biologically sensitive element with a physical transducer toelectively and quantitatively detect trace amounts of biologicalamples (Andle and Vetelino, 1995; Ballantine et al., 1997; Cavict al., 1999; Janshoff et al., 2000).

This review is focused on acoustic wave microelectromechan-cal systems (MEMS) biosensors that are fabricated using MEMSechniques such as lithography, etching, deposition of thin metallms and sputtering of piezoelectric thin layers. The MEMS fab-ication of the acoustic wave biosensors enables the integrationf these devices with the corresponding electronic circuits. Theompatibility of the piezoelectric layer fabrication with the MEMSechnology is critical for robust sensor fabrication. Sensors fabri-ated using standard complementary metal oxide semiconductorCMOS)-MEMS technology facilitate the integration of the acous-ic wave devices with the electronic circuits (necessary for datacquisition) on a single silicon chip. The CMOS-MEMS fabricationerformed by a CMOS foundry reduces the production costs of thecoustic wave biosensors considerably.

To contain or to allow biological fluids to flow on the acousticave MEMS biosensors, micro-fluidic channels have to be placed on

op of the sensors. The method to integrate the acoustic wave MEMSiosensors with micro-fluidic depends on the type of biologicalpplication. For some acoustic wave MEMS biosensors, the sen-or and the micro-fluidic channel or reservoir are fabricated at theame time. Other acoustic wave MEMS biosensors are integratedn polymer microfluidic chambers after the sensor fabrication haseen completed.

For biological detection the biosensor is functionalized with spe-ific molecules. When a bioanalyte interacts with this sensing layer,hysical, chemical, and/or biochemical changes are produced. Massnd viscosity changes of the biospecific layer can be detected byecording changes in the acoustic wave properties such as velocity,ttenuation, resonant frequency and time delay. The main advan-age of the SAW devices is the fact that they can be operated withimple and cheap electronic components. Fully integrated systemsnclude the sensors, its readout circuitry and a microfluidic systemntegrated with the biosensor chip. The measurement of the reso-ant frequency or time delay can be performed with high degreef precision using electronic systems. In the last decade, robustnd portable acoustic wave devices used for biological applicationsave gained attention. The majority of these devices were fabri-ated on thick commercial piezoelectric substrates such as lithiumiobate (LiNbO3), lithium tantalate (LiTaO3) and quartz (Arrudat al., 2009; Bender et al., 1997, 2000; Bisoffi et al., 2008; Branchnd Brozik, 2004; Dahint and Bender, 1998; Gizeli et al., 2003; Gratend Frye, 1996; Guhr et al., 2005; Jakoby and Vellekoop, 1997; Josset al., 2001; Kovacs et al., 1992; Länge et al., 2008; Lec, 2001; Mollt al., 2007; Schweyer et al., 1997; Vellekoop et al., 1987, 1994;elsch et al., 1996; White and Voltmer, 1965).This review paper first presents acoustic wave biosensors fab-

icated using MEMS technology that could be integrated withlectronic circuitry on a single silicon chip. Piezoelectric filmsabricated using MEMS technology are thinner, can reach higherrequencies and offers the advantage of monolithic integration withhe advanced CMOS technology for peripheral readout and sig-al processing circuitry. These devices include MEMS film bulkcoustic wave resonators (FBAR), surface acoustic wave (SAW)

esonators and delay lines. Integration of these (FBAR, SAW res-nators and delay lines) in micro-fluidic systems are described next.icro-reservoirs which contain the biological fluids under test can

e placed either on top or underneath the MEMS biosensor. The

The analyte molecules binding to the immobilized antibodies on the sensor surfaceinfluence the velocity of the SAW and hence the output signal generated by thedriving electronics.

micro-fluidic systems are usually fabricated using polymers andoffer protection to the water-sensitive electronics. Sensor func-tionalization using specific antibodies or antigen is also described.The antibodies or antigen promote binding between the moleculesunder test and the sensor surface. Finally, this review articleconcludes with an analysis of successful MEMS acoustic wavebiosensors and its potential usage in arrays.

2. Acoustic wave based MEMS devices as biosensors

MEMS technology uses manufacturing techniques perfected inthe integrated circuit (IC) foundries to fabricate miniature mechan-ical and biological systems. The obvious advantage of using ICfabrication techniques is that the biosensors can be mass-produced,are low-cost, small sized and portable. The current trend in MEMStechnology is to incorporate both the acoustic wave sensor and itssignal processing circuits on a single silicon substrate. There areseveral published papers describing interesting biological applica-tions of acoustic wave based devices (Arruda et al., 2009; Benderet al., 1997, 2000; Bisoffi et al., 2008; Branch and Brozik, 2004;Dahint and Bender, 1998; Gizeli et al., 2003; Grate and Frye, 1996;Guhr et al., 2005; Jakoby and Vellekoop, 1997; Josse et al., 2001;Kovacs et al., 1992; Länge et al., 2008; Lec, 2001; Moll et al.,2007; Schweyer et al., 1997; Vellekoop et al., 1987, 1994; Welschet al., 1996; White and Voltmer, 1965). The majority of the biosen-sors presented in these papers are based on a thick piezoelectricsubstrate as; lithium niobate, lithium tantalate or AT-cut quartz(Bender et al., 2000; Bisoffi et al., 2008; Branch and Brozik, 2004;Dahint and Bender, 1998; Gizeli et al., 2003; Guhr et al., 2005; Josseet al., 2001; Moll et al., 2007; Schweyer et al., 1997; Welsch et al.,1996). Even though these acoustic wave based sensors demon-strated high sensitivities to biological analytes, the focus of thisreview paper is on acoustic wave MEMS biosensors, for whichthe thick piezoelectric layer characteristic to conventional acousticwave biosensors is replaced with thin piezoelectric layer fabricatedusing MEMS technology.

The current biomedical applications of the acoustic wave basedMEMS devices focuses on two types of devices: FBAR and SAW. TheFBAR device is fabricated from a thin layer of piezoelectric mate-rial with excitation electrodes fabricated on both sides of the thinpiezoelectric film. The acoustic waves propagate unguided throughthe volume of the thin piezoelectric film. The detection mechanismsoccur at the opposite surfaces of the piezoelectric film. As shown

in Fig. 1, SAW devices have the interdigitated transducers (IDTs)excitation electrodes fabricated on the one side of the piezoelectricfilm. As a result, the SAW devices have the acoustic waves prop-agating along the surface of the piezoelectric substrate. The SAW

sors and Bioelectronics 33 (2012) 1– 9 3

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2005). High frequency operations improve a sensor’s detectioncapabilities, as sensitivity is a function of resonant frequency. Shearacoustic wave propagation is preferred for liquid operations sinceshear waves do not compress into the liquid, thereby reducing

I. Voiculescu, A.N. Nordin / Biosen

evice could be resonator or delay line depending of the design ofhe IDTs. For SAW resonators the IDTs are fabricated in a centralosition and reflectors are added on both sides of the input andutput IDTs to trap the acoustic energy within a cavity.

The SAW delay line does not have reflectors and there is a signif-cant distance between the input and output IDTs. The input IDTsenerates and transmits the acoustic waves while the output IDTss used to detect the acoustic waves. To transform a SAW delayine device into a biosensor, the surface between these two sets ofDTs is covered with a biological layer sensitive to the analyte to beetected, as illustrated in Fig. 1. The absorption of the analyte onhe sensitive layer will produce a time delay in the acoustic waveropagation. The SAW delay lines measure the time required forhe acoustic wave to travel from one set of IDTs to the second setf IDTs. SAW resonators detect analytes by measuring the resonantrequency. Mass loading due to the presence of analytes load thecoustic wave and causes the resonance frequency to decrease.

The acoustic wave biosensors presented in this review are fab-icated using microfabrication processes characteristic for MEMSevices and are integrated in microfluidic channels or microreser-oirs to perform biological detection. The main disadvantage of thecoustic wave MEMS biosensors is the degradation of performanceue to the damping associated with liquid loading. In liquid, theuality factor (Q) drops (usually more than 90% reduction) and neg-tively affects the device’s sensitivity. Since most of the biologicalpplications are performed in liquid only a few types of acousticave devices could be integrated in microfluidic channels, without

ignificant degradation of the sensor performance.There has been interest in applying the FBAR resonator as

iosensors because of low damping of the acoustic wave in theiquid (Mason and Thurston, 1972; Zhang et al., 2009). FBARs areompact and robust and can be integrated with CMOS technologyn silicon substrates. FBARs are fabricated from a thin central piezo-lectric layer with metal electrodes micropatterned on both sides.BAR are considered MEMS devices because the thin piezoelec-ric layer and the actuation electrodes are fabricated using MEMS

icroprocessing techniques that allow for further integration withlectronic circuits. The FBARs can be integrated in a microfluidicystem and used for biosensing applications.

.1. FBAR devices used for biological applications

In this review section five MEMS FBARs biosensors are pre-ented. Each resonator is evaluated in terms of its fabricationomplexity and performance. MEMS FBAR’s are popular due tots high quality factor, Q which range from a few hundreds to aew thousands in air and vacuum. These high Qs are achievableince the FBAR’s creates a very large acoustic impedance mismatchetween the solid materials and air, thus confining the generatedcoustic waves within a cavity (Guhr et al., 2005). This structure isompromised in biological applications since the MEMS FBARs aremmersed in liquid and the acoustic impedance mismatch at theolid–liquid interface is reduced. As a result, the acoustic waveseak at the solid–liquid interface and the Q of the FBAR is degrades.o counteract this, microfluidic channels on top of the FBAR haveo have heights comparable to the acoustic wavelength in FBARMason and Thurston, 1972). This minimizes the dissipation in theiquid and hence improves the resonator’s Q.

The first example of a MEMS FBAR biosensor is illustrated inig. 2 (Zhang et al., 2009). This MEMS FBAR biosensor operates inhe longitudinal mode and uses ZnO as its piezoelectric thin film.o enable future integration with electronic circuits on the same

ubstrate, the MEMS FBAR resonator was fabricated on a <1 0 0>ilicon wafer. To electrically isolate the silicon substrate, a thin sil-con dioxide layer is thermally grown on the wafer and coveredy a thin SiN layer which forms the membrane. The bottom FBAR

(b) Top view of the fabricated FBAR sensor; a microfluidic channel run across theFBAR sensor.

Reproduced with permission from Zhang et al. (2009). Copyright 2009 IEEE.

electrode is formed using a thin Al layer, while the top FBAR elec-trode uses Cr/Au layers for bio-compatibility. Au was chosen as thetop electrode due to its excellent conductivity and good affinityfor biomolecular binding. The silicon wafer was backside etchedusing deep reactive ion etch (DRIE) to release the SiN membrane.As shown in Fig. 2a, the SiO2 layer is wet etched using a bufferedoxide etchant to reveal a smooth SiN membrane. Isolation of theFBAR from the Si substrate improves the Q of the FBAR. The top Auelectrode of the FBAR represents the sensing electrode. For biologi-cal applications, the top electrode will be coated with a thin layer ofbiological material. The microfluidic channel is formed using a 3 �mthick parylene film. To reduce Q damping in liquid environments,the microfluidic channels’ height is designed to be comparable tothe FBAR’s acoustic wavelength. This method improved the sensor’sQ in liquid to 110 compared to 15 with shallow microchannels.

The second example illustrates a high frequency (GHz) MEMSshear mode FBAR biosensor as shown in Fig. 3 (Wingqvist et al.,

Fig. 3. Schematic of a shear wave FBAR with a microfluidic transport system.

4 I. Voiculescu, A.N. Nordin / Biosensors and Bioelectronics 33 (2012) 1– 9

Fig. 4. Schematic illustration of the FBAR sensor. The sensor is part of the bottomp

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late of a liquid reservoir.

eproduced with permission from Mertig et al. (2009) Copyright 2003 IEEE.

nergy losses. This MEMS biosensor uses 2 �m thick AlN film as itsiezoelectric material with Al top and bottom electrodes. Acousticaves are generated in the region where both the top and bottoml electrodes overlap. To generate shear acoustic waves, 30◦ c-axis

ilted AlN thin films were sputtered on the Al electrode. To furthereduce liquid damping, the silicon substrate was back-side etchedo acoustically isolate acoustic wave resonator and to create a cav-ty. When connected to a series of vertical channels, the cavity forms

microfluidic channel. The microfluidic channel allows biologi-al analytes to flow underneath the bottom electrode of the FBAR.nlike conventional FBARs sensors, which have its sensing elec-

rodes at the top, this unique sensor uses the bottom electrode asts active sensing surface. A thin, bio-compatible gold layer covershe bottom electrode to enhance its suitability for biological test-ng. The sensor was tested with different concentrations of albuminn solution and the detection limit was 0.3 ng/cm2.

Another effort to improve the FBAR’s Q factor was implementedy placing the ZnO FBAR on top of an acoustic mirror as shown inig. 4 (Farnell and Adler, 1972; Gabl et al., 2003; Weber et al., 2006).uch structures are known as solidly mounted FBARs. The acousticirror is formed using 3-fold alternating ZnO and Pt layers. Acous-

ic isolation from the substrate is achieved when a Bragg reflector,omprised of layer pairs with contrasting acoustic impedance islaced underneath the FBAR. The FBAR includes a bottom elec-rode, the ZnO film and a top-electrode, which are all integratedbove a silicon substrate. The 850 MHz FBAR has shown Qs of 312n air and 192 in water (Weber et al., 2006) while the 2.2 GHz FBARas shown Q of 400 in air. The FBAR sensors have been fabricatedn silicon substrates employing reactive magnetron sputtering ofnO, with the crystallographic axis inclined with an angle of 16◦

elative to the surface normal. This crystallization of ZnO with ailt angle allows the operation of the device in the shear mode.he shear mode is preferred for the operation in the liquid of theBAR resonator because of minimum damping of the quality factor,. The 100 nm Au top-electrode is used for the binding of biore-eptor molecules. The active vibrating region of the resonator isoated with a receptor layer sensitive to the biological analyte toe detected. The attachment of the analyte molecules leads to an

ncrease of the resonator mass load and a decrease of the resonantrequency which can be electrically determined. To verify the usef this type of FBAR in biosensing, an assay of avidin–anti avidinas used (Weber et al., 2006). The detection limit demonstrated

2

ith this sensor was 2.3 ng/cm .A novel structure to isolate the FBAR from liquid damping and

degradation is illustrated in Figs. 5 and 6 (Pottigari and Kwon,009). A thin vacuum gap is inserted between the FBAR and liquid,

Fig. 5. Schematic illustration of the cross section of the V-FBAR.

effectively protecting the FBAR from liquid damping and acousticenergy loss. The vacuum gap is formed by placing several parylenemicroposts on the FBAR’s top electrode. The microposts support aparylene diaphragm, which is in contact with liquid. During oper-ation, the liquid loads the diaphragm and this mass loading istransferred to the FBAR through the microposts. Using this mech-anism, the liquid damping effect is separated from the operatingfrequency of the resonator. This results in high Qs of 140 in liq-uid without losing mass sensitivity. Similar to the first example,this vacuum (V)-FBAR is fabricated on a silicon substrate. A thin(0.6 �m) silicon nitride layer is deposited on top of silicon. Thesilicon substrate is back-side etched to form a membrane and toacoustically isolate the FBAR from the lossy silicon layer. Both thetop and bottom electrodes of the V-FBAR are formed using Al layers.A 0.7 �m thick ZnO film is used as the piezoelectric material. Thetop sensing diaphragm was fabricated from a 1.6 �m thick pary-lene layer. The height of the vacuum space is 2 �m and the activesensing area is 200 �m × 200 �m (Pottigari and Kwon, 2009).

A nano-FBAR biosensor that uses nanomaterials to increase thesurface of the sensitive area was presented in the literature (Chenet al., 2009). To achieve both increased surface area and high Q,this device is built on Si substrates with an acoustic mirror consist-ing of alternating quarter-wavelength silicon dioxide (SiO2) andtungsten (W) layers. Fig. 7 illustrates the MgxZn1 − xO FBAR biosen-sor with ZnO nanostructures. The SiO2/W Bragg reflector layersacoustically isolate the FBAR from the lossy Si substrate. High-quality, piezoelectric MgxZn1 − xO thin films is RF-sputtered on the

Fig. 6. Micrograph of fabricated V-FBAR.Reproduced with permission from Pottigari and Kwon (2009). Copyright 2009 IEEE.

I. Voiculescu, A.N. Nordin / Biosensors and Bioelectronics 33 (2012) 1– 9 5

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microfluidic systems for biosensing is the damping of the acous-

ig. 7. Schematic diagram of the MgxZn1-xO FBAR structure. Au top electrode is usedo facilitate MOCVD growth of ZnO nanostructures.

BAR’s top Au electrode using metal organic chemical vapor depo-ition (MOCVD). The nanostructures provide a very large sensingrea, faster response, and higher sensitivities over the planar sensoronfiguration. Due to the large sensitive area, the mass sensitivityf this biosensor is higher than 1720 Hz cm2/ng. In order to employhis nano-FBAR for biosensing, the nanostructured ZnO surface wasunctionalized to selectively immobilize DNA. The total frequencyhift was about 26 MHz for 16.25 ng of hybridized DNA and linkerolecules combined.

.2. SAW delay lines as biosensors

Different from bulk wave resonators, SAW devices focus theirnergy on the surface instead of dispersing the energy through theiezoelectric substrate to the opposite surface. SAW delay linesave been successfully employed for biosensing and can be inte-rated in microfluidic systems. A MEMS-SAW delay line consistsf two IDTs that are electrode pairs on the same side of a thiniezoelectric film. There is a space or delay line between the inputnd output IDTs. When a sinusoidal electrical voltage is appliedo the input IDTs, oscillating mechanical strains in the piezoelec-ric thin film form travelling surface acoustic waves. The SAW isetected and converted back into a sinusoidal voltage of differ-nt phase and amplitude at the output IDTs. These differences areelated to changes in the velocity of the SAW and can be corre-ated to changes in the mass loading, viscosity and temperature ofhe substrate. For biosensors, a layer of bio-molecules consistingf protein cross-linkers and antibodies is coated on the sensitiveurface of the sensing device in the path of the travelling waves,etween the two sets of IDTs. If specific target proteins (antigens)re present, they bind to the antibodies, creating the mass load-ng on the surface of the substrate. As a result, a time delays in theropagation of the SAWs will occur (Ballantine et al., 1997; Gratend Frye, 1996; Jakoby and Vellekoop, 1997; Kovacs et al., 1992;ellekoop et al., 1987, 1994).

The effect of the temperature on the piezoelectric material coulde compensated by using a dual delay line configuration (Arrudat al., 2009). In this case only one of the SAW devices will beunctionalized with biological molecules. The reference SAW delayine is not functionalized with biological molecules and it is usednly for temperature compensation. Both SAW delay lines willperate at the same temperature. Measurements comparing the

xperimental delay line with the reference delay line are used toompensate the effect of temperature on the biosensor (Kovacst al., 1992).

Fig. 8. SEM of a CMOS-SAW die.Reproduced with permission from Tigli and Zaghloul (2007). Copyright 2007 IEEE.

Recently a SAW delay line biosensor fabricated in standardCMOS-MEMS technology was presented in the literature. Due tothe CMOS fabrication these biosensors, shown in Fig. 8, the biosen-sors are inexpensive (Tigli et al., 2010; Tigli and Zaghloul, 2007).The CMOS-SAW devices were fabricated in AMI Semiconductor’s0.5 �m 3-metal 2-poly technology. The piezoelectric material wasZnO, which is a CMOS compatible material but not a standard CMOSlayer. Typically, CMOS fabricated chips do not have ZnO as one ofits layers. Post-CMOS microfabrication steps were used to fabri-cate the additional ZnO layer. After the CMOS chip was receivedfrom the AMI semiconductor foundry the ZnO piezoelectric layerwas RF magnetron sputtered on top of the CMOS chip. This CMOSSAW biosensor was designed to detect cancer markers. This biosen-sor employs a streptavidin/biotin-based five-layer immunoassayfor detecting the breast cancer biomarker, mammoglobin (hMAM)(Tigli et al., 2008, 2010). The mass sensitivity of this CMOS SAWresonator reported in literature was 8.7 pg/Hz.

Surface acoustic wave delay line technology can be applied tocreate highly sensitive biosensors due to its extreme sensitivity tosurface perturbation. In SAW delay lines, the velocity of the acous-tic wave is perturbed by the binding of antigens with antibodies,immobilized in the path of the travelling wave. The mass loadingchanges the velocity of the wave and consequently changes thetime required for the acoustic wave to travel between the inputand output IDTs.

2.3. SAW resonators as biosensors

To improve the performance of the SAW delay lines, reflectorsare placed beside the input and output IDTs. The acoustic wave isgenerated by the central IDTs and reflected by reflector bars fabri-cated on both sides of the actuation IDTs. The reflector bars containthe acoustic waves within a cavity, reducing losses of the deviceand create a distinct resonant peak (series resonance frequency).Such devices are termed as SAW resonators. SAW resonators havealso been demonstrated as successful biosensors. Selective bindingof antigens with antibodies generate mass change in the path of thetravelling wave. The mass loading changes the velocity of the waveand consequently shifts the frequency of the travelling SAW. Moni-toring resonant peaks of resonators are easier than monitoring timedelays of delay lines and produce sensors with better sensitivity.

One problem of the SAW resonators when integrated in

tic wave. To overcome this, shear-horizontal (SH)-SAW resonatorshave been designed and have the advantage of less energyloss when used in liquids. The piezoelectric film used for the

6 I. Voiculescu, A.N. Nordin / Biosensors and Bioelectronics 33 (2012) 1– 9

Fig. 9. Schematic representation of the SAW biosensor device showing the sensitivedetection window between the input and output interdigitated transducers (IDTs)fw

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or protein immobilization. This biosensor is based on a Love wave mode deviceith the ZnO film as piezoelectric layer and the SiO2 film as the waveguide layer.

H-SAW resonators is fabricated with a specific crystallographicrientation that generates waves that could propagate in thehear horizontal motion. When SH-SAW resonators are placednderneath a liquid chamber, acoustic energy is not radiated hori-ontally into the liquid. Instead, the acoustic waves have transverseisplacements polarized parallel to the surface resulting in lowernergy loss (Drafts, 2001). Shear-horizontal waves can be guided bylacing a thin guiding layer on a SH-SAW sensor. Waves that prop-gate through the guiding layer are known as Love waves, havearticle displacements which are parallel to the device surface andormal to the wave propagation (Branch and Brozik, 2004). To con-ne the acoustic energy to the sensing surface, the acoustic shearelocity in the waveguide layer has to be less than the shear velocityn the substrate (Gizeli et al., 2003). Acoustic efficiency is improvedn such cases, provided that the waveguide layer is thin and doesot load or attenuate the travelling acoustic wave.

A high sensitivity ZnO/SiO2/Si Love mode SAW resonatoriosensor for the detection of interleukin-6 (IL-6) in the humanerum, was reported in the literature (Krishnamoorthy and Iliadis,006) and it is illustrated in Fig. 9. This biosensor is based on aove mode device with the ZnO film as piezoelectric layer and theiO2 film as wave guide layer. Love mode acoustic waves are shearorizontal waves that concentrate the acoustic energy in a guiding

ayer grown on an appropriate substrate, thus offering the opportu-ity of higher mass sensitivities (Jakoby and Vellekoop, 1997). Thisode of acoustic wave propagation depends on the orientation of

he guiding layer film. In this device, using the coordinates shown,he shear waves propagate in the y-axis direction and the guidinglm is SiO2. To achieve Love mode propagation in ZnO, the grownlm should be along the c-axis and the IDTs must be patterned per-endicular to the c-axis of the film (Gorla et al., 2000). The effectivelectromechanical coupling coefficient of the Love mode waves fornO exhibits its maximum when it is grown perpendicular to the-axis in the [1 1 0 0] orientation. This biosensor is compatible withMOS technology and is used as a portable real time detection sys-em for the interleukin family of proteins in human serum. Therowth of ZnO film on SiO2 layers has been reported in the liter-ture (Gizeli et al., 1996), but the effect of SiO2 layer thickness onhe quality of ZnO films has not been studied.

This Love mode SAW biosensor was fabricated by photolitho-raphic processing of ZnO thin films grown on SiO2/Si. The ZnOhin films were grown by pulsed laser deposition (PLD) on a00 A thick SiO2 layer that was grown on a (1 0 0) Si substrate byry oxidation in an atmospheric furnace at 1025 ◦C, as describedreviously (Krishnamoorthy and Iliadis, 2006). The pulsed laser

blation system for the ZnO deposition consisted of a KrF lasert a wavelength of 248 nm, operating with an energy density of.7 J/cm2 at a frequency of 10 Hz. The growth temperature and oxy-en partial pressure during growth were 250 ◦C and 1 × 10−4 Torr,

Fig. 10. Integration of an array of FBAR sensors with microfluidic chamber for bio-logical measurements. The FBAR sensors are completely immersed in the liquid.

respectively. These were optimal growth conditions that resulted inhigh quality c-axis oriented ZnO thin films on the SiO2/Si substrate.

When ZnO in epitaxial thin film form is deposited on SiO2/Sisubstrates, it offers the advantage of monolithic integration withthe advanced CMOS technology for peripheral readout and signalprocessing circuitry. At this moment, ZnO is the only piezoelec-tric material that could be integrated in CMOS technology becauseit can be grown at low temperatures, a condition imposed not toadversely affect the on-Si circuits. Despite its promise ZnO grow-ing thin film that is c-axis oriented wurtzite crystal structure onthe cubic Si substrates, is a challenging task. Yet, high quality ZnOc-axis oriented thin films suitable for device development havebeen successfully grown on Si and SiO2/Si substrates by pulsed laserdeposition, providing a new material template for sensor develop-ment.

The input and output IDTs and the reflector bars were fabri-cated by e-beam evaporation of Al with the thickness 500 A. Twobiosensors operating at 747.7 MHz and 1.586 GHz were fabricated.The first device is designed to be a resonator with finger width andspacing of 1 �m and 2.4 �m, respectively. The second device alsoa resonator was designed for higher frequency of operation up to586 GHz with a thinner finger width and spacing of 0.8 �m each.The number of fingers in the input and output IDTs was 15. Thelength of the IDTs was 45 �m with an aperture of 40 �m for bothdevices.

This biosensor was used for the detection of IL-6 in humanserum. The sensitive area between the two sets of input and outputIDTs was functionalized by immobilizing the monoclonal IL-6 anti-body onto the ZnO biosensor surface both through direct surfaceadsorption and through covalent binding on gluteraldehyde. Themass sensitivity was demonstrated to be 4.456 �m2/pg.

3. Integration of the acoustic wave MEMS devices inmicrofluidic systems and sensor functionalization forbiological tests

3.1. Integration in microfluidic systems

The integration of the acoustic wave MEMS biosensor in themicrofluidic system is a complex process. The integration techniqueis determined by the sensor fabrication process and the type of thebiological applications. In some applications the sensor could beembedded in a microfluidic reservoir. When the biological appli-cation requires several types of solutions to be introduced on thesensor sensitive area and extracted from this system, the biosensorwill be integrated in a microfluidic channel. In other applications,the acoustic wave device is covered with polymer to protect the

electronics from water and a small detection window is patternedin the polymer between the excitation fingers. This structure allowsthe possibility to deposit drops of the biological solution on thebiosensor’s active surface for detection.

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The FBAR biosensor can be integrated in a microfluidic chan-el fabricated on the Si substrate or fabricated from an additional

ayer of polymer. When an array of FBARs biosensors is used, thehole array of biosensors could be placed in a microfluidic cham-

er, as illustrated in Fig. 10 (Weber et al., 2006). The part of thehip containing the FBAR array is embedded in a fluidic chamberade from acrylic glass with inflow and outflow tubes that can be

onnected to a microfluidic system. A viton strip ensures that theart with the biosensors is perfectly sealed and the fluid flow isontained inside the cell. The FBAR array was designed such thathe resonator active area is placed inside the microfluidic chamberhereas the electrical contacts are located outside.

In other applications, the FBAR sensor is monolithically fabri-ated with the microfluidic reservoir, as shown in Fig. 4. The upperlectrode of the FBAR represents the detection area and is placedt the bottom of the microfluidic reservoir. In other configurations,he FBAR sensor is integrated in microfluidic channel that are fabri-ated using MEMS processes, as illustrated in Figs. 2 and 3.The FBARensor shown in Fig. 3 has the microfluidic channel fabricated in theilicon substrate and placed below the AlN piezoelectric layer. Forhe FBAR sensor illustrated in Fig. 2, a parylene layer is fabricatedver the FBAR sensor and the inlet and outlet of the microfluidicystem were fabricated in the silicon substrate.

For biological applications, the Love mode SAW resonatorhown in Fig. 9 has a detection window fabricated on the sensitiverea of the sensor (Krishnamoorthy et al., 2008). A polymer layer iseposited such that it covers the upper side of the Love mode SAWesonator. A window of 20 �m × 20 �m was opened in the polymerayer in the sensitive area between the IDTs. This window repre-ents a microfluidic reservoir for testing the SAW resonator withiquid solution containing biological molecules.

For biological detection with the CMOS-SAW delay lines shownn Fig. 8, microdrops of biological fluids were dispensed with

icropipettes on the sensitive area of the biosensor. In this case,he CMOS chip was not embedded in a microfluidic channel andid not have a reservoir fabricated on the chip. The commercialMOS chip has miniature dimensions of 2 mm × 2 mm, making

t difficult to handle this chip. To enable easy manipulation, theMOS chip was glued on a glass slide using wax (Tigli et al., 2010).icropipettes connected to a six-well plate were used to dispense

ifferent biological microdrops to the sensitive area of CMOS chip.he commercial six-well plate contained different biological liquidsnd was micromachined to accommodate the micropipette tips.

.2. Sensor functionalization

Device functionalization is a difficult task which depends on theype of biological analyte to be detected and the material used forhe fabrication of the detection area. The majority of the acous-ic wave based MEMS biosensors sense proteins secreted by cellss cancer markers, avidin or interleukin. The detection process isased on antibody/antigen sensing. The sensor detection area isunctionalized with antibodies corresponding to the specific targetrotein to be detected.

The shear mode FBARs biosensor integrated in a fluidic cham-er shown in Fig. 10 was demonstrated for avidin detectionWeber et al., 2006). Anti-avidin was used as an antibody becauset binds to the sensor’s gold detection surface. Once the goldetection area has been covered with anti-avidin molecules,ovine serum albumin (BSA) was introduced in the microfluidicystem to remove the anti-avidin molecules that were looselyound. BSA is a small molecule that is used to fill the gaps between

he anti-avidin molecules. When target molecules of avidin arentroduced in the microfluidic system, the BSA molecules does notllow the avidin molecules to bind directly to the sensor surface.he FBAR sensor is sensitive to each injection of anti-avidin, BSA

nd Bioelectronics 33 (2012) 1– 9 7

or avidin in the microfluidic system, causing a significant sensorresponse manifested as distinct variations of the resonant fre-quency.

The ZnO/SiO2/Si Love mode SAW has a window fabricated in thesensitive area to allow the detection of IL-6. For the functionaliza-tion of the ZnO sensitive area, a special surface activation techniquewas developed. The sensor surface was activated with hydroxyla-tion, followed by silanization with 3-aminopropyltriethoxysilane(ATES), used as substrate for the binding of gluteraldehyde. Thesensitive surface was washed several times in ultra pure water toremove the excess gluteraldehyde. The ATES solution was appliedon the ZnO surface at room temperature using low temperaturevapor priming technique. The monoclonal IL-6 antibody covalentlybinds with gluteraldehyde and monoclonal recombinant human-Escherichia. coli derived from IL-6 antibody-500 �g. The IL-6 proteinwas then immobilized on the IL-6 antibody. When the target ana-lyte IL-6 was immobilized onto the antibody layer, the resonantfrequency shift can be measured and the target protein mass eval-uated.

The CMOS-SAW biosensor was used for the detection ofbreast cancer biomarker, mammoglobin (hMAM) (Tigli et al.,2010). The CMOS-SAW biosensor employs a streptavidin/biotinimmunoassay for the detection of the cancer biomarker. This sen-sor was not integrated in a microfluidic channel and microdropsof biological liquids were dispensed on the detection area. Thefluorophore-conjugated streptavidin was thiolated and immobi-lized onto the gold sensitive surface of the sensor. The primarygoal of this step was to attain the highest possible surface cov-erage by the streptavidin molecules. PBS was used for washingthe biological molecules that were not bounded to the sub-strate. The biotinylation of hMAM antibody (Ab) was performed toallow the antibody immobilization onto the streptavidin surface.A rabbit polyclonal Ab against hMAM was used for biotinyla-tion with the biotin. The unbound biotinylated Ab was removedwith several washes with PBS. After the immobilization of thehMAM Ab was completed, the binding was confirmed usingoptical and fluorescence microscopy. This immobilization tech-nique provides high-efficiency functionalization of the Au sensitivesurface. The hMAM protein binds to the hMAM Ab and wasdetected by variations in the CMOS-SAW sensor’s resonant fre-quency.

As observed from these applications the integration of theacoustic wave MEMS biosensors with biological fluids is influ-enced by the type of biological applications. For some acousticwave MEMS biosensors, the sensor fabrication and the microflu-idic channel fabrication were performed at the same time andthe microchannels were fabricated in silicon. Other acoustic waveMEMS biosensors were integrated in polymer microfluidic chan-nels fabricated after the sensor fabrication was completed. Asimple, square detection window at microscale dimensions couldbe fabricated from photoresist on the sensitive area of the sen-sor. In the case of CMOS-SAW biosensors, microdrops of biologicalsolution were dispersed on the sensor surface for detection. All theMEMS biosensors presented in this review had their sensor’s sur-face functionalized with a specific antibody corresponding to thetarget antigen protein to be detected.

4. Conclusions

Acoustic waves based MEMS devices offer a promising tech-nology platform for the development of biosensors. This review

presents acoustic wave devices fabricated using piezoelectric thinfilms characteristic to MEMS technology that can be integratedwith electronic circuitry on a single silicon chip. The thin piezo-electric films fabricated by MEMS technology can reach higher

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requencies and offers the advantage of monolithic integrationith CMOS technology for peripheral readout and signal pro-

essing circuitry. For biological applications these MEMS devicesre integrated in microfluidic system as microfluidic channels, orhambers. The main problem of the acoustic wave based MEMSevices when used as biosensors is the degradation of their per-ormance due to liquid damping. Several types of FBAR resonatorsere reported in the literature as biosensors with high sensitiv-

ty. These sensors were MEMS fabricated using piezoelectric thinlms of AlN and ZnO. In order to minimize the damping effectnd the reduction of quality factor in liquid, shear mode of oper-tion was also considered. The shear mode AlN FBAR biosensoremonstrated a detection limit of 0.3 ng/cm2. In comparison, thehear mode ZnO biosensor has a detection limit of 2.3 ng/cm2. Aano-FBAR biosensor that uses ZnO nanostructures to increase theensitive area can also be used as a biosensor. These nanostructuresffer a very large sensing area, faster response, and higher sensi-ivities over the planar sensor configuration. However, fabricatingnO nanostructures on the FBAR resonator are very difficult, mak-ng the device unreliable and commercialization of this biosensorrohibitive.

To improve sensitivity, an original FBAR sensor was fabricatedith a vacuum gap between a parylene sensing membrane and

he resonator. This design enhanced the sensitivity of the FBARy separating the liquid damping effect from the operating fre-uency of the device. The addition of a vacuum gap increases theetection capacity of the FBAR while maintaining high Q in the

iquid.The SAW based biosensors presented in review were fabricated

sing MEMS micro-processing techniques and also CMOS-MEMSabrication. CMOS technology is attractive for the fabrication ofAW based devices due to its potential of batch/mass produc-ion and simple integration of the SAW sensor with standardMOS technology. Only ZnO, and to some extent AlN, piezoelec-ric materials are CMOS-compatible because it can be sputtered as

CMOS-post processing step at a low temperatures. Fabrication ofortable acoustic wave MEMS devices enables system minimiza-ion through integration of the sensor and the electric circuits on aingle chip and the reduction of the power consumption. WirelessF signals can be used to remotely measure or monitor physi-al, chemical and biological quantities using acoustic wave MEMSevices, without the need of a power supply.

In summary, acoustic wave based MEMS devices can be success-ully used as biosensors, based on antibody/antigen recognitionystem. Among these biosensors, FBAR devices using inclinediezoelectric films, SAW and Love-wave devices show great poten-ial for applications in highly sensitive bio-detection systems foroth dry and liquid environments. Based on the published liter-ture in this area, the FBAR biosensors and especially the shearode FBARs are best suited for biological applications because

hese devices maintain high quality factor, Q during the operation iniquids. Arrays of FBAR biosensors can be fabricated, with each FBARunctionalized with different biological molecules for the detectionf complex analytes. Simple fabrication processes for the integra-ion of FBARs in Si wafers with readout circuitry and the possibilityf fabrication of arrays of FBARs make these devices interesting foriosensing applications.

cknowledgement

This work was supported in part by the NSF-EAGER award

umber 0940111, the US Army Research Office, Brain Gainalaysia Incentive: MOSTI/BGM/R&D/500-2/6(2) and the Funda-ental Research Grant: FRGS 0308-90 provided by the Ministry ofigher Education of Malaysia.

nd Bioelectronics 33 (2012) 1– 9

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