ORIGINAL PAPER

Resonant gravimetric immunosensing based on capacitivemicromachined ultrasound transducers

Darius Virzonis & Gailius Vanagas & Almira Ramanaviciene & Asta Makaraviciute &

Dovydas Barauskas & Arunas Ramanavicius & Weijia Wen & Rimantas Kodzius

Received: 1 July 2013 /Accepted: 20 March 2014# Springer-Verlag Wien 2014

Abstract High-frequency (40 MHz) and low-frequency(7 MHz) capacitive micromachined ultrasound transducers(CMUT) were fabricated and tested for use in gravimetricdetection of biomolecules. The low-frequency CMUTsensorshave a gold-coated surface, while the high-frequency sensorshave a silicon nitride surface. Both surfaces were functional-ized with bovine leukemia virus antigen gp51 acting as theantigen. On addition of an a specific antibody labeled withhorseradish peroxidase (HRP), the antigen/antibody complexis formed on the surface and quantified by HRP-catalyzed

oxidation of tetramethylbenzidine. It has been found that aconsiderably smaller quantity of immuno complex is formed onthe high frequency sensor surface. In parallel, the loading of thesurface of the CMUTwas determined via resonance frequencyand electromechanical resistance readings. Following the for-mation of the immuno complexes, the resonance frequencies ofthe low-frequency and high-frequency sensors decrease by up to420 and 440 kHz, respectively. Finite element analysis revealsthat the loading of the (gold-coated) low frequency sensors isseveral times larger than that on high frequency sensors. Theformation of the protein film with pronounced elasticity andstress on the gold surface case is discussed. We also discuss theadoption of this method for the detection of DNA using ahybridization assay following polymerase chain reaction.

Keywords Label-free biomolecule detection .

Immunosensing . Capacitivemicromachined ultrasoundtransducers . Gravimetric sensors . Sensor arrays .

DNA hybridization

Introduction

Microelectromechanical systems (MEMS) devices, such asmicro/nano cantilevers, thin film bulk resonators (FBAR),surface acoustic waves (SAW) devices and capacitivemicromachined ultrasound transducers (CMUT) nowadaysare commonly being used as resonant, acoustic or gravimetricchemical and biochemical sensors [1–6]. Application ofMEMS for label-free chemical/biochemical detection is supe-rior to classical chemical analysis due to their throughput andminiaturization potential. Also MEMS can be comparativelyeasily integrated with microelectronic circuits. This createssignificant potential for distributed, low-power, low-costchemical and biochemical sensing.Mechanical part ofMEMSis to be modified by the molecules that exhibit specific

D. Virzonis (*) :G. Vanagas :D. BarauskasDepartment of Technologies, Kaunas University of TechnologyPanevezys Faculty, Daukanto 12, 35212 Panevezys, Lithuaniae-mail: darius.virzonis@ktu.lt

A. Ramanaviciene :A. MakaraviciuteNanoTechnas – Center of Nanotechnology and Materials Science,Department of Analytical and Environmental Chemistry, Faculty ofChemistry, Vilnius University, Naugarduko 24, 03225 Vilnius,Lithuania

A. RamanaviciusDepartment of Physical Chemistry, Faculty of Chemistry, VilniusUniversity, Naugarduko 24, 03225 Vilnius, Lithuania

A. RamanaviciusLaboratory of BioNanoTechnology, Department of Material Scienceand Electrical Engineering, Semiconductor Physics Institute, StateResearch Institute Center for Physical and Technological Sciences,A. Gostauto 11, 01108 Vilnius, Lithuania

W. Wen :R. KodziusKAUST-HKUST Micro/Nanofluidic Joint Laboratory, Kowloon,Hong-Kong

R. KodziusBiological and Environmental Science and Engineering Division(BESE), Computer, Electrical and Mathematical Sciences andEngineering Division (CEMSE) and Computational BioscienceResearch Center (CBRC), King Abdullah University of Science andTechnology, Thuwal 23955-6900, Kingdom of Saudi Arabia

Microchim ActaDOI 10.1007/s00604-014-1244-3

interaction with the analyte, which preferably binds moleculesto the modified surface and thus introduces physical, chemi-cal, and/or biochemical changes. Then corresponding massand viscosity changes can be detected, such as the change ofthe resonance frequency, acoustic wave propagation velocityand attenuation [7]. Both micro/nano cantilever [2] andCMUT [1] techniques have been demonstrated to have tensof zeptogram (10−21 g) sensitivity. Basic CMUT structure isthe array of electrostatically actuated micrometer-sized plates(commonly referred to as “membranes”) supported on theisolating posts and having individually sealed vacuum cavi-ties. A sealed CMUT structure is subjected to less dampingthan a cantilever structure with correspondingly better poten-tial for higher frequencies and higher resonance quality (Q)factor. Thickness-extensional mode resonance used in FBARdevices is even more advantageous over flexural mode reso-nance (used in cantilevers and CMUTs) since it is consider-ably less damped during operation in liquids. The drawback ofFBARs is the sensitivity, which is still a few orders of mag-nitude lower. SAW delay-line type structures have similaradvantages and disadvantages [5, 7]. In addition to the sensi-tivity the advantage of the CMUT approach over FBARs andSAW is electrostatic actuation, which eliminates the use ofpiezoelectric materials and enables the potential to fabricatethe devices with standard CMOS technology [8]. CMUTs canalso be fabricated in large one-dimensional and two-dimensional arrays of individually wired elements, enablinga number of parallel measurement channels while still main-taining the compact dimensions of an entire device. Thisresults in a high degree of parallelism, which is a great advan-tage in biosensing, when a high number of different probes isrequired in a small sensor area. For example, it is recommendedto have at least three probe design for the recognition ofcomplimentary DNA strand– the first of which is 100 % com-plimentary (positive signal), second of which has one or twomismatch (a negative control) and the third one which is notcomplimentary to the target at all (100 % mismatch). CMUTbased biosensing can benefits here from mastered fabricationtechnologies of compact large arrays, two-dimensional matrix-es with bottom side contacts and dedicated electronics [9, 10].Significant efforts towards matrix sensing have been also madeby developers of optical biomolecule sensors, for example byusing resonant waveguide gratings with direct 2D imaging andexploiting micropads spatial profiles on a nanopatterned chip[11], however, CMUT platform is still superior in terms ofsimplicity of readout electronics and ease of integration ofsensor with electronic circuits.

We have reported earlier the potential use of CMUTs asimmunosensors sensors [6]. In addition to a well known reso-nance frequency change measurement we have determinedsignificant informative content in the real part of the measuredCMUT impedance. However, real-time measurement of thespecific interaction of proteins has not been demonstrated yet

due to specific requirements for the sensor surface modification,immobilization of proteins and control of the measurementconditions. As specific interaction of biomolecules, such asproteins and DNA takes place in liquid environment, it conflictswith the mass resonance sensor approach, which requiresCMUT to be dry in order to maintain the adequate resonancequality. Direct adsorption of antigen on the sensor surface fromsolution used in our previous work [6], was difficult to controland resulted in comparatively large uncertainty intervals.

In the current work we aim to analyze the relationship be-tween the CMUT surface modification methods and measure-ment conditionswith the changes ofCMUTresonance frequencyand resistance, which were identified earlier as potentially infor-mative parameters that can give adequate sensing performance.

Experimental

CMUT fabrication and critical data

Two types of CMUTs, having 7 and 40 MHz resonancefrequencies were used in this research. Low frequency(7 MHz) devices (LF-CMUT) were fabricated by the waferbonding process, where 1 μm thick monocrystalline siliconwas directly bonded to the 38×38 μm sqare, 150 nm deepcavities fabricated over highly doped and oxidized siliconsubstrate [12]. Each 0.50×0.25 mm sensor element contains50 capacitive cells electrically connected in parallel. Threehundred nm thick gold electrodes and bonding pads weredeposited and patterned on top of the membranes to providecommon conditions for immobilization of biological speciesas well as to improve electrical coupling to the sensing sur-face. Figure 1 shows 7 MHz sensor elements (a) and 13-element array wirebonded to the printed circuit board withepoxy passivation of electrical interconnects (c). High fre-quency (40 MHz) CMUTs (HF-CMUT) were fabricated bythe sacrificial release process [13, 14]. 100 nm vacuum gap of16 μm round capacitive cells was defined by depositing andpatterning the sacrificial chromium film over highly doped,oxidized and pre-patterned silicon substrate. 200 nm thickaluminum electrodes were burried between two layers ofPECVD silicon nitride films. 140 nm thick bottom and100 nm thick upper silicon nitride films establish the mem-brane. Bottom film also plays the role of an isolating layer thatprevents CMUT from short circuiting during collapse.

An additional 300 nm layer of PECVD silicon nitride wasdeposited after sacrificial etching of chromium and criticalpoint drying to seal the etching vias, thus making total platethickness 740 nm in the electrode area and 540 nm in the restof the membrane area. Finally bonding pads for top andbottom electrodes were opened. A high frequency sensorcontains 2,500 capacitive cells electrically connected inparrallel. Figure 1 shows a 40 MHz sensor element (b) and a

D. Virzonis et al.

row of five sensors mounted to the printed circuit board withepoxy passivated wirebonding (d). Detailed structure of thecapacitive cells is explained in Fig. 2, which also shows thestructure of the finite element analysis (FEA) model.

Obtaining sensor readings

Resonance frequency and the values of the real part of theimpedance (“resistance”) weremeasured by a network analyzer(Agilent 4395A) equiped with the impedance measurement kit.Forty volt direct current bias was used during all measure-ments. The resonance frequency and the peak resistance value

measured at the resonance were extracted from the measuredfrequency spectra. Ten elements of LF-CMUT were involvedin tests to have more parrallel readings. The HF-CMUT read-ings were obtained from a single sensor. Measured resonancefrequency data initially were interpreted based on the parallel-plate equivalent circuit model described in [14], in relation tothe modification of the sensors surface with an increasedmoving mass of the membrane. For more adequate esti-mation of non-linear effects induced by the sensor mod-ification and immune complex formation we used thefinite element analysis, which is described in the sub-chapter of this paper.

Fig. 1 Micrographs of CMUT sensors: a 7 MHz 0.5×0.25 mm multiel-ement sensor with external gold electrodes; b 40MHz 1.3×1.3 mm singlesensor with burried aluminum electrode. Photos of mounted and

wirebonded sensors: c 7 MHz multielement sensor; d five 40 MHzsensors in a row

Fig. 2 CMUTcapacitive cell structure used in finite element analysis: a low frequency sensor cell with surface gold electrode; b high frequency sensorcell with burried aluminum electrode

Resonant gravimetric immunosensing by CMUT

Modification of CMUTwith antigen for sensingthe interaction with specific anti-gp51-HRP antibody

For the immunoassay and confirmation of the immune com-plex formation on CMUT we used bovine leukemia virusantigen gp51 (BLV gp51) obtained from BIOK (Kursk, Rus-sia, http://www.biok.ru) and a monoclonal BLV gp51 specificantibody-horseradish peroxidase conjugate (anti-gp51-HRP)from enzootic bovine leukemia virus gp51 antibody test kit(IDEXX Laboratories, Westbrook, Maine, USA, http://www.idexx.com). BLV gp51 (33 μg mL–1) was deposited on thegold coated LF-CMUT surface and on the silicon nitridesurface of HF-CMUT. After drying, the deposit was cross-linked covalently by exposing it to glutaraldehyde vapour.Covalent cross-linking of BLV gp51 molecules was per-formed exploiting the volatility of the glutaraldehyde accord-ing to the procedure described earlier [15]. The CMUTsurfacemodified with BLV gp51was exposed to 25% glutaraldehydevapour for 15 min at 20 °C. By using this method covalentbonds were formed between the proteins using bifunctionalreagent with aldehyde groups that react with primary aminesof amino acid residues. Chemical modification of solubleproteins due to the cross-linking result in minimal changeson the original protein structure [16]. As a result, a thin BLVgp51 layer was formed on the surface allowing further inter-action with specific antibodies. The same modification proce-dure was used for gold and silicon nitride surfaces. Then, thesensor surface modified by BLVantigen gp51 was allowed toreact with anti-gp51-HRP for 20 min (concentrated anti-gp51-HRP conjugate was diluted in water at a ratio of 1:100 v.v.).After each deposition, immobilization or interaction step thesensor surface was rinsed with deionized water and dried inthe air at the room temperature for 10 min or until no remain-ing wet residues were observed. Both resonance frequencyand electromechanical resistance of the sensors were mea-sured after each step by the network analyzer, so the sensorloading was completely traceable. In order to addition-ally confirm the formation of immune complexes on thesurface (BLV gp51/anti-gp51-HRP), tetramethylbenzydine(TMB) was added as an enzyme substrate. In the presenceof an enzyme, the substrate was oxidized resulting in blue

colour. The changes of colour were observed in the solutionvisually, and higher colour intensity was interpreted as ahigher complex formation rate. The differences of substratesolution colours collected from all elements with gold andsilicon nitride surfaces was very clear and visually easy todistinguish. Thus, in our experiments TMB was used only forqualitative comparison of different surfaces. To evaluate howthe readings of the sensors were impacted by complete drying,after the tests, sensors were rinsed and dried as usual and thensubjected to 4-day aging in room conditions. After the aging,measurements of the resonance frequency and electromechan-ical resistance were repeated.

Synthesized oligonucleotides

DNA probes and target representing the sex determiningregion Y (SRY) gene marker synthesized for hybridizationtests are shown in a Table 1 (see also [17]). Positive probes are100 % complementary to the target DNA, while single mis-match negative probes have a non-complementary base in themiddle of the sequence and totally negative probe is non-complementary to the target at all.

Sensor surface regeneration

The surface of CMUT after measurements and tests wasbrought to the initial state by cleaning the sensing surfaceswith oxygen plasma. Cleaning was performed in 3-min stepsin reactive ion etching process chamber (Vision 320RIE) at250 W, 60 mTorr pressure and 80 sccm O2 flow. Each etchingstep was followed by the resonance frequency and resistancevalues measurement to gather the information of gradualunloading of the CMUT structure.

Finite element analysis

To estimate the resonance behavior of CMUT and explain thereadings obtained during sensoric CMUToperation we devel-oped the finite element model; its structure is schematicallyshown in Fig. 2. A single capacitive CMUT cell, which is thebasic element of the sensors, was simulated as a complex

Table 1 DNA probes and target representing the SRY gene marker

Oligo name 5′-modification Sequence 3′-modification Purpose

Thiol-SRY23pos disulfide linkage ATAAGTATCGACCTCGTCGGAAG None Positive probe

Thiol-SRY23neg disulfide linkage ATAAGTATCGAACTCGTCGGAAG None Single mismatch negative probe

SRY24-ThiolPos None GCGAAGATGCTGCCGAAGAATTGC disulfide linkage Positive probe

SRY24-ThiolNeg None GCGAAGATGCTACCGAAGAATTGC disulfide linkage Single mismatch negative probe

Thiol-NonComp100neg disulfide linkage CGAGTCCGCGTGCCTACTATTACT None Total negative probe

3-47SRY None GCAATTCTTCGGCAGCATCTTCGCCTTCCGACGAGGTCGATACTTAT

None Target DNA

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structure: layered membrane supported over the vacuum cav-ity with isolating posts. Finite element analysis packageANSYS (Canonsburg, PA) was used for simulation. Twomembrane types, corresponding to two kinds of CMUT sen-sors used in this work, were simulated. LF-CMUT (Fig. 2a):square, 1 μm thick monocrystalline silicon membrane with38μm side dimension and 300 nm thick gold electrode on top.HF-CMUT (Fig. 2b): 16 μm diameter disc shaped membranecomposed of three layers – 140 nm thick silicon nitrideisolation layer, 200 nm thick aluminum electrode and400 nm thick upper membrane layer. In both cases membraneswere perimeter-fixed. Each CMUTstructure layer was dividedby 3D SOLID45 elements, but the very top layer, representingprotein film, was divided by 3D SOLID185 elements. Thislayer mimicked the layer of proteins established after surfacemodification and immune complex formation. For increasedadequacy of the model we made several assumptions,explained in the corresponding paragraph. The vacuum gapwas divided by TRANS126 elements. Bias voltage andambient pressure were preset as initial boundary conditions.

Simulated electromechanical impedance of a CMUT cellwas calculated using harmonic analysis results obtained as anoutput of the model. Generic equation for complex impedancecalculation was used [18]:

Zm ¼ ImΔdeffωε0εgAm

þ jdg−ReΔdeffωε0εgAm

þ jdm

ωε0εmAmþ j

diωε0εiAm

ð1Þ

Here ω=2πf is cyclic frequency; f stands for the oscillationfrequency; Δdeff is effective deflection of the membrane; dgstands for the nominal height of the vacuum gap; dm is thethickness of the membrane; di stands for the isolation layerthickness; Am is an area of the membrane, which in our case isequal to the electrode area; ε0, εg, εm and εi are vacuumdielectric constant and dielectric permitivities of the gap,membrane and isolation layer respectively. Calculated sole-cell impedance spectra were extrapolated to the entire senorarea dividing calculated values by the number of the cells inthe sensor.

Several assumptions regarding the morphological and me-chanical properties of the established layer based on biomol-ecules were made. First, analyzed molecules were thought toinherit elasticity and mass density properties of bovine serumalbumine, which are 13.2 GPa and 1,170 kg·m−3 correspond-ingly, according to [19]. Second, when analyzing the initialmodification step of LF-CMUT based on BLV gp51 adsorp-tion on the gold surface followed by cross-linking with glu-taraldehyde, it was thought that porous layer based on ran-domly oriented/distributed biomolecules structure havingthickness from 300 to 500 nm, which is equivalent to severaltens of molecules, was established. Such assumption explainsthe increase of resonant frequency observed by the LF-CMUTafter the initial modification step. The morphology of

biomolecule-based layer was supposed to be uniform in thecase of HF-CMUT. The difference in behaviours of LF-CMUT and HF-CMUT was caused by different coverage ofresonators of these CMUTs, because the adsorption of BLVgp51 protein based layer on gold-coated surface was muchhigher. Such assumption explains the increase of resonantfrequency observed by the LF-CMUT after the initial modifi-cation step (this was not the case in HF-CMUT).

The porosity of the protein layer was estimated during FEAby substituting the original elasticity modulus and mass den-sity values with the values characteristic of porous material.The porosity factor, describing the ratio between the massdensity of the porous material to the mass density of the bulkmaterial [20] was taken 0.7, which resulted in effective elas-ticity modulus of 6.5 GPa and mass density of 819 kg·m−3.This filling factor was increased to 1.0 to reflect the massincrease due to the immune complex formation.

We also assumed that in the case of LF-CMUT formed BLVgp51 layer has developed additional intrinsic stress after drying,which was in the range of 10 MPa. This assumption proved tomatch the experimental data. To distinguish between the newlyformed and fully dried (4-day aging) immune complex wesimulated an extra 100 nm layer with 1,000 kg·m−3mass density(equivalent to water) for the newly formed immune complex.

Results and discussion

Testing of fabricated CMUTs

After fabrication the sensors underwent functionality tests,common for CMUT. One of the tests, namely frequencyspectra evolution of the HF-CMUT in respect of the biasvoltage, is presented in Fig. 3. The shift of the resonancefrequency and increase of the peak resistance is related tothe decreasing gap; the increasing efficiency is due to theincreasing bias voltage [12, 14].

Immunoassay sensing tests of LF-CMUT

Low frequency 10 channel multisensor surface modificationwith glutaraldehyde crosslinking of BLV gp51 and subse-quent interaction with anti-gp51-HRP gave CMUT resonancefrequency and peak resistance readings shown in Fig. 4a. Forbetter representation we excluded from the diagram the read-ings of four sensor elements, which exhibited very little or noshift of the resonance frequency and/or resistance. Readingswere normalized to the natural resonance frequency values ofunmodified sensor to eliminate the scatter of initial parame-ters. The term “natural resonance frequency” used here canalso be understood as a background signal of unmodifiedsensor. One can observe the increase of resonance frequencyand corresponding decrease of the peak resistance after the

Resonant gravimetric immunosensing by CMUT

surface modification step. This conflicts with the linear modelstating that extra mass deposited on the CMUT surface de-creases the resonance frequency [14]. We attribute this to thepossibility that the modified surface after crosslinking, rinsingand drying develops extra stress and stiffness, which increasesoverall spring constant of CMUT membrane and competeswith the effect of the increased moving mass. We simulatedthis behaviour by finite element analysis as described in theExperimental section. The effect of simulated sensor surfacemodification and immune complex formation on the imped-ance frequency spectra is shown in Fig. 4b). These spectrawere obtained by adding 300 nm thick layer with 13 GPaYoung modulus, 1,170 kg·m−3 mass density and 0.7 fill factor(see explanation in Experimental section) over the CMUTstructure to simulate the initial surface modification step.The filling factor was increased to 1.0 to reflect the immunecomplex formation. In Fig. 4a) FEA output illustrates eachsensing step by dashed and dotted lines. The line ends showthe limits of simulated sensor readings corresponding to dif-ferent protein layer thickness, from 300 to 500 nm. Also theaverage mass of the protein layer corresponding to the singleCMUT cell, as derived from the model output, is indicated.For the best match to the experimental results we also adjustedintrinsic stress and mechanical loss factor in the protein layerfor each simulated case. After interaction with anti-gp51-HRPresonance frequency decreased by 3 % (204 kHz) on average;maximum decrease is 6% (420 kHz). The resistance increasedby 38 % (4.2 Ω) on average with maximum increase by 51 %(5.8 Ω). The dark-blue colour of the solution after TMBwetting of the CMUT surface and after the measurementconfirmed the immune complex formation.

The results of this experiment confirm the informativity ofthe sensor and adequacy of the surface modification tech-nique. Sensor readings were repeatedly measured after 4-dayaging.

Repeated measurements showed only partial degradationof the sensor readings, namely 17% decrease of the resistanceif compared to the readings of as-formed complex and lessthan 1 % increase of the resonance frequency.

Immunoassay sensing tests of HF-CMUT

Surface modification and test measurements of the HF-CMUT was performed simultaneously with previously de-scribed LF-CMUT tests and the same previously describedmethod with only one difference was used: the high frequencysensor provided silicon nitride surface instead of gold. Sensorreadings are presented in Fig. 5a) in the form of the spectraobtained by the network analyzer. Here we observe continu-ous decrease of the resonance frequency and increase of theresistance due to the surface modification and immune com-plex formation. After surface modification by BLV gp51 theresonance frequency decreases by 210 kHz (0.5 %) and theresistance increases by 0.5 Ω (6.2 %). As expected, this is theeffect of increased mass on CMUTsurface without significantbuild-up of extra stress and elasticity.

Fig. 3 Evolution of the spectra of the high frequency (40MHz) sensor inrespect the bias voltage

Fig. 4 Low frequency multisensor readings: a obtained from 6 elementsafter surface modification with antigen (BLV gp51) and after interactionwith antibody (anti-gp51-HRP); b simulated sensor spectra

D. Virzonis et al.

After interaction with the analyte (anti-gp51-HRP) thetendency of further decrease of the resonance frequency andsignificant increase of the resistance prevails: frequency isfurther decreased by 440 kHz (1.1 %) and resistance is in-creased by 0.2 Ω (2.5 %).

The intensity of the blue colour of TMB deposited on theHF-CMUT surface after the immune complex formation waslower than in the LF-CMUTcase, confirming lower density ofimmobilized antigens and correspondingly a lower ammountof complexes formed.

Figure 5b) shows the evolution of the HF-CMUT imped-ance frequency spectra simulated by finite element analysis.We used the same simulation technique as for the LF-CMUT,except the representation of the protein layer. Surface modifi-cation was simulated by adding a continuous 120 nm thicklayer with 1,170 kg·m−3 mass density and 3.1 MPa elasticitymodulus; the immune complex formation was simulated byadding extra 130 nm of the same layer. Thickness of thesimulated layer was chosen to provide the best match to themeasured resonance frequency changes. For improved matchto the measured impedance spectra we also adjusted thecoefficient of mechanical losses in the simulated layer. Re-maining mismatch of simulated and measured spectra canbe attributed to the non-informative parameters of theinterconnects and measurement hardware.

CMUT sensor capability to detect hybridizationof immobilized molecules

Oligonucleotides are commercially available, with their pri-mary application as primers for PCR. The oligonucleotidescan function in a wide range of modifications, such as amine,biotin, thiol functional groups or fluorophore. Whilesilanization of glass permits to attach thiol-bearing DNA, adirect attachment of disulfide-bearing DNA molecules toCMUTsurface allows an easy and simple way to functionalizethe sensor. A simple assay can be set up to prove CMUTmicrofluidic systems capability to detect and measure the pres-ence of target specific DNA molecules in the solution. Positiveand negative oligonucleotide probes with disulfide linkagegroup can be used for CMUT surface functionalization: Thiol-SRY23pos, Thiol-SRY23neg, SRY24-ThiolPos, SRY24-ThiolNeg and Thiol-NonComp100neg. The functionalizedpositive probes will be able to recognize the amplified PCRproduct [17], or as for simulation purposes, the single strandedDNA oligonucleotide 3-47SRY, while the target single strandedDNA molecule 3-47SRY will not hybridize to the negativeprobes and will be washed away (Table 1). Therefore sensorwill measure the positive signal only for the positive probes.The DNA sensor can be regenerated, as hybridized targetmolecules can be removed by lowering the pH or by increasingtemperature, reconstituting the single stranded probe moleculeson the gold surface. As demonstrated previously [21] the DNAtarget concentrations as low as 4.16 μM could be detected byquarz crystal microbalance working at 50 MHz, and this valueprobably can be better for CMUT sensors due to much highersensitivity potential [1, 6].

Sensor surface regeneration

Oxygen plasma cleaning was carried out to test if sensorreadings could be brought to the initial state and to have thereadings during gradual sensor surface unloading as additional

Fig. 5 High frequency (40 MHz) sensor readings (a) and simulatedreadings (b)

Fig. 6 Sensor unloading by the oxygen plasma reactive ion etching process

Resonant gravimetric immunosensing by CMUT

reference. Between the oxygen plasma processing steps(3 min each), sensor resonance frequency and resistance weremeasured. Obtained readings show that sensor resonancefrequency and resistance reached initial value, saturating after6–12 min. The cleaning process is illustrated with one of thelow frequency multisensor elements in Fig. 6.

Conclusions

High and low frequency CMUTwere fabricated and used formonitoring the specific antigen-antibody interaction. LF-CMUT measurements of BLV gp51/anti-BLV gp51-HRPcomplex formation have shown that the used concentrationof antigen BLV gp51 (33 μg mL−1) layer coated on goldsurface after glutaraldehyde cross-linking, rinsing and dryingis likely to develop the elasticity value of 13.2 GPa, whichcauses the LF-CMUT resonance frequency increase by sever-al hundreds of kHz. The influence of protein layer elasticitywas not observed in the HF-CMUT case, because, as identi-fied in simulation, the equivalent thickness of protein layerwas over two times thinner. We explain it as a result of weakadsorption of BLV gp51 to the silicon nitride surface. Thehigher BLV gp51 immobilization rate on the gold surface wasqualitatively confirmed by the TMB oxidation intensity ob-served after immune complex formation with anti-gp51-HRPas higher intensity of the blue colour. According to the modelparameters fitted to the experimental results we estimate thesensitivity of the LF-CMUT to be not worse than4.0 fg (Hz cell)−1 and HF-CMUT not worse than205 ag (Hz cell)−1. Better protein adsorption to the gold sur-face in LF-CMUT case provides the opportunity to lower theconcentration of BLV gp51 during initial sensor modification.At the same time, HF-CMUT has better mass sensitivity; thisproperty can be successfully used for sensing low analyteconcentrations. We demonstrate that both LF-CMUT andHF-CMUT types of biosensor can be used at room tempera-ture, perform measurements in the solution, are highly sensi-tive, selective, reproducible and stable. The promising appli-cations of the sensor are determined by biomolecules, such asproteins and DNA in solution.

Acknowledgement This research is funded by a grant (No. MIP-059/2012) from the Research Council of Lithuania.

References

1. Lee HJ, Park KK, Kupnik M, Khuri-Yakub BT (2012)Functionalization layers for CO2 sensing using capacitivemicromachined ultrasonic transducers. Sensors Actuators B Chem174:87–93

2. Long Z, Kou L, Sepaniak MJ, Hou XD (2013) Recent advances ingas phase microcantilever-based sensing. Rev Anal Chem 32(2):135–158. doi:10.1515/revac-2012-0034

3. Johnston ML, Kymissis I, Shepard KL (2010) FBAR-CMOS oscil-lator array for mass-sensing applications. Ieee Sens J 10(6):1042–1047. doi:10.1109/Jsen.2010.2042711

4. Tigli O, Bivona L, Berg P, Zaghloul ME (2010) Fabrication andcharacterization of a surface-acoustic-wave biosensor in CMOS tech-nology for cancer biomarker detection. Ieee Trans Biomed Circ Syst4(1):62–73. doi:10.1109/Tbcas.2009.2033662

5. Senveli SU, Tigli O (2013) Biosensors in the small scale: methodsand technology trends. Iet Nanobiotechnol 7(1):7–21. doi:10.1049/iet-nbt.2012.0005

6. Ramanaviciene A, Virzonis D, Vanagas G, Ramanavicius A (2010)Capacitive micromachined ultrasound transducer (cMUT) forimmunosensor design. Analyst 135(7):1531–1534. doi:10.1039/C0an00104j

7. Voiculescu I, Nordin AN (2012) Acoustic wave based MEMS de-vices for biosensing applications. Biosens Bioelectron 33(1):1–9.doi:10.1016/j.bios.2011.12.041

8. Lemmerhirt DF, Cheng XY, White RD, Rich CA, Zhang M,Fowlkes JB, Kripfgans OD (2012) A 32 x 32 capacitivemicromachined ultrasonic transducer array manufactured instandard CMOS. Ieee Trans Ultrason Ferr 59(7):1521–1536.doi:10.1109/Tuffc.2012.2352

9. Zhang XF, Wygant IO, Lin DS, Kupnik M, Oralkan O, Khuri-Yakub BT (2009) Wafer-bonded 2-D CMUT arrays incorpo-rating through-wafer trench-isolated interconnects with asupporting frame. Ieee Trans Ultrason Ferr 56(1):182–192. doi:10.1109/Tuffc.2009.1018

10. Vanagas G, Virzonis D, Paukstaitis V, Baranauskas D, Cerviakov S(2010) Integrated front end electronics design for micromachinedultrasound transducers. Elektron Elektrotech 7:117–120. doi:10.5755/J01.Eee.18.9.2804

11. Bougot-Robin K, Kodzius R, Yue W, Chen L, Li S, Zhang XX,Benisty H, Wen W (2013) Real time hybridization studies by reso-nant waveguide gratings using nanopattern imaging for SingleNucleotide Polymorphism detection. Biomed Microdevices. doi:10.1007/s10544-013-9832-2

12. Vanagas G, Virzonis D, Mikolajunas M (2008) Testing the capacitivemicromachined transducers produced with wafer bonding process.Itelm’2008: Intelligent Technologies in Logistics and MechatronicsSystems:132–139

13. Morkvenaite-Vilkonciene I, Virzonis D, Vanagas G, Krikscikas V(2012) Operating point of capacitive micromachined ultrasonic trans-ducers with sub-structural elements. Elektron Elektrotech 9:43–46.doi:10.5755/J01.Eee.18.9.2804

14. Viržonis D, Vanagas G, Kodzius R (2013) Integration of ca-pacitive micromachined ultrasound transducers to microfluidicdevices. In: Panzarella S, Maroni W (eds) Microfluidics: con-trol, manipulation and behavioral applications. Nova, NewYork City

15. GermanN, Ramanavicius A, Voronovic J, Oztekin Y, RamanavicieneA (2011) The effect of gold nanoparticle colloidal solution on per-formance of glucose oxidase modified carbon electrode. MicrochimActa 172(1–2):185–191

16. Barbani N, Giusti P, Lazzeri L, Polacco G, Pizzirani G (1995)Bioartificial materials based on collagen. 1. Collagen cross-linking with gaseous glutaraldehyde. J Biomater Sci-Polym E7(6):461–469

17. Wu JB, Kodzius R, XiaoK, Qin JH,WenWJ (2012) Fast detection ofgenetic information by an optimized PCR in an interchangeable chip.Biomed Microdevices 14(1):179–186. doi:10.1007/s10544-011-9595-6

18. Ballandras S, Wilm M, Daniau W, Reinhardt A, Laude V, Armati R(2005) Periodic finite element/boundary element modeling of

D. Virzonis et al.

capacitive micromachined ultrasonic transducers. J Appl Phys 97(3).doi:10.1063/1.1839634

19. Perticaroli S, Nickels JD, Ehlers G, O’Neill H, Zhang Q, Sokolov AP(2013) Secondary structure and rigidity in model proteins. SoftMatter 9(40):9548–9556. doi:10.1039/C3sm50807b

20. Lee D,Wei X, Chen X, Zhao M, Jun SC, Hone J, Herbert EG, OliverWC, Kysar JW (2007)Microfabrication and mechanical properties of

nanoporous gold at the nanoscale. Scri Mater 56(5):437–440. doi:10.1016/j.scriptamat.2006.08.069

21. Bustabad EA, Garcia G, Rodriguez-Pardo L, Farina J, Perrot H,Gabrielli C, Bucur B, Lazerges M, Rose D, Compere C, Arnau A(2009) A biosensor for detection of DNA sequences based on a50MHz QCM electronic oscillator circuit. 2009 Ieee Sensors 1–3:687–690. doi:10.1109/Icsens.2009.5398346

Resonant gravimetric immunosensing by CMUT