Fabrication and Characterization of a Surface-Acoustic-Wave Biosensor in CMOS Technology for Cancer -eQ2.pdf

IEEE TRANSACTIONS ON BIOMEDICAL CIRCUITS AND SYSTEMS, VOL. 4, NO. 1, FEBRUARY 2010

Fabrication and Characterization of a

Surface-Acoustic-Wave Biosensor in CMOS

Technology for Cancer Biomarker Detection

Onur Tigli , Member, IEEE , Louis Bivona, Patricia Berg, and Mona E. Zaghloul , Fellow, IEEE

ABSTRACT— Design, fabrication, and characterization of a

novel surface acoustic wave (SAW) biosensor in complementary

metal–oxide semiconductor (CMOS) technology are introduced.

The biosensor employs a streptavidin/biotin-based ve-layer

immunoassay for detecting a prominent breast cancer biomarker,

mammoglobin (hMAM). There is a growing demand to develop

a sensitive and specic assay to detect biomarkers in serum that

could be used in the early detection of breast cancer, determining

prognosis and monitoring therapy. CMOS-SAW devices present

a viable alternative to the existing biosensor technologies by

providing higher sensitivity levels and better performance at low

costs. Two architectures (circular and rectangular) were developed

and respective tests were presented for performance comparison.

The sensitivities of the devices were analyzed primarily based on

center frequency shifts. A frequency sensitivity of 8.704 pg/Hz and

a mass sensitivity of 2810.25 m kg were obtained. Selectivity tests

were carried out against bovine serum albumin. Experimental

results indicate that it is possible to attach cancer biomarkers to

functionalized CMOS-SAW sensor surfaces and selectively detect

hMAM antigens with improved sensitivities, lowered costs, and

increased repeatability of fabrication.

INDEX TERMS— Biosensor, cancer, complementary metal–oxide

semiconductor

(CMOS),

microelectromechanical

systems

When compared to their competition, acoustic-wave sensors

present a solid example of the effective use of acoustic-wave

devices in diverse applications. They are versatile, highly

sensitive, reliable, reusable, small, inexpensive, can easily be

designed for responding to various measurands, have a wide

dynamic range, and they are passive devices which can also be

deployed as wireless units. Therefore, acoustic wave devices

present attractive alternatives to their counterpart technologies

in their corresponding sensor applications. The application

of interest in this paper, namely, biosensor, presents a solid

example of how effective acoustic-wave sensors can be used in

diverse applications.

Surface acoustic waves (SAWs) are generated at the free

surface of piezoelectric material. The application of a varying

voltage to the metal interdigital transducer (IDT) generates the

acoustic wave on the input side. The acoustic wave generated by

the input IDT travels through the region called the delay line and

reaches the output IDT where the mechanical displacements

due to the acoustic waves create a voltage difference between

the output IDT ngers. In order to employ mass loading on

SAW sensors, the device surfaces should be functionalized by

selective coatings that react with the entity under analysis. This

interaction, while causing a mass loading, produces a shift in

resonant frequency, which then can be measured to analyze

the entity being sensed. The rst examples of this approach

came from the eld of chemical sensors. Wohltjen and Dessy

reported the rst use of SAW devices for chemical analysis [2].

They laid out instrumentation design for chemical analysis,

which uses SAW devices.

The principle of biosensors is very similar to the chemical

vapor sensors that were listed in the literature. They detect tar-

geted biological species in uid or solid form by employing

mass loading on acoustic-wave devices. This fundamental prin-

ciple is depicted in Fig. 1 along with a typical frequency-shift-

based response. The surface of these devices also requires a spe-

cial coating to functionalize the sensors for the analytes of in-

terest. Several techniques were developed and applied for a va-

riety of analytes. These are DNA-RNA hybridization, adsorp-

tion of proteins on functionalized surfaces, lipid-protein inter-

actions on membranes and piezoimmunosensing [3]. Due to

the high specicity of antigen-antibody reactions and the well-

structured generation of antibodies against a variety of biolog-

ical materials made it possible to build immunosensors that em-

ploy acoustic-wave devices. The application example that is se-

lected for this research is essentially a piezoimmunosensor that

(MEMS), surface acoustic wave (SAW).

for more than 60 years [1]. Although the telecommunica-

tions industry has been the primary employer of these devices,

they are enjoying a large surge in several new emerging and

developing industries. These industries constitute automotive,

medical, and commercial applications. The vast majority of

these industries use acoustic-wave devices as sensors. These

sensors include but are not limited to torque, pressure, biolog-

ical, chemical, temperature, vapor, humidity, and mass sensors.

Manuscript received February 12, 2009; revised May 18, 2009. First pub-

lished December 04, 2009; current version published January 27, 2010.

An earlier version of this paper was presented at IEEE Sensors Conference

28-31 Oct. 2008 and was published in the conference proceedings.

O. Tigli is with Washington State University Vancouver, School of Engi-

neering and Computer Science, Vancouver, WA 98686 USA (e-mail: tigli@wsu.

edu).

L. Bivona and P. Berg are with the George Washington University, Biochem-

istry and Molecular Biology Department, Washington, DC 20052 USA (e-mail:

lbivona@gwu.edu; bcmpeb@gwu.edu).

M. E. Zaghloul is with the George Washington University, Electrical and

Computer Engineering Department, Washington, DC 20052 USA (e-mail: za-

ghloul@gwu.edu).

Color versions of one or more of the gures in this paper are available online

at http://ieeexplore.ieee.org.

Digital Object Identier 10.1109/TBCAS.2009.2033662

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I. I NTRODUCTION

A COUSTIC wave devices have been in commercial use

TIGLI et al. : FABRICATION AND CHARACTERIZATION OF A SAW BIOSENSOR

Fig. 1. Sketch of a typical SAW-based bio/chemical sensor using the mass

loading principle. The center frequency shifts of the reference and the loaded

devices are used to detect the change in the mass that is perturbing the SAW.

reported for radiolabelled IgG adsorption [16]. As an example

of vapor phase SAW biosensing, Stubbs et al. demonstrated the

detection of uranine vapor on anti-FITC (uorescein isothio-

cyanate)-coated SAW devices [39].

This paper introduces the biosensor application through an

overview of the design fabrication and use of CMOS-SAW

sensors along with a thorough comparison of rectangular and

circular CMOS-SAW devices. After the completion of the

proof-of-concept phase of this research, the CMOS-SAW de-

vices were characterized thoroughly for performance [18]–[20].

Based on the results of this characterization, designs were im-

proved for better performance. The following sections lay out

the primary components of a biosensor system. Fabrication

and characterization of these important components are laid

out in detail. Sample results are presented for the experiments

that were carried out for developing robust testing procedures,

functionalizing gold-covered chips with streptavidin/biotin

and antibody/antigen interactions. This paper concludes with

the discussion of performance, and testing issues for the

CMOS-SAW-based biosensors.

employs antigen-antibody binding to detect cancer biomarker

mammoglobin (hMAM).

The applications of piezoelectric immunosensors are abun-

dant and reported widely in literature. They include medical ap-

plications, such as bacteria, virus, and toxin detection [4], clin-

ical diagnosis and analysis [5], as well as food industry and en-

vironmental analysis for the detection of compounds using anti-

body-antigen reactions [6]. Piezoelectric immunosensors for the

detection of viruses and bacteria associated with acute diarrhea

were reported in [7] with limits of microorganisms mL.

Zhou et al. demonstrated the detection limits of /mL for

hepatitis A and B viruses [8]. Immunoglobin M with detection

limits of 5.5 cells/mL and insulin with a detection limit of 0.17

cells/mL were reported in [9] and [10], respectively. Cocaine de-

tection with limits of 33 cells/mL and nicotine detection of 0.025

cells/mL were reported in [11] and [12], respectively. Other than

the nicotine work, all of the listed biosensors work employed

antigen-antibody binding as the assay principle. The nicotine

work used molecularly imprinted polymer.

Although most of this initial piezoelectric immunosensor

work was carried out by using (QCM) in the past, the increased

sensitivities—primarily due to much higher operating fre-

quencies—of SAW-based sensors attracted a lot of biosensor

research interest and effectively replaced the use of QCMs

over the years. Its integration possibilities with the current

cutting-edge microelectromechanical systems–very-large-scale

integrated (MEMS-VLSI) fabrication technologies and the

production of higher sensitivities at lower overall costs made

them a very attractive tool to exploit for biosensor applications.

Ganske et al. demonstrated the use of SAW as an immuno-

biosensor for the detection of human interleukin-6 (IL-6) [13].

They reported IL-6 molecule concentrations of 15 pg/mL to

100 pg/mL with a frequency shift range of 4500 to 5700 Hz.

Berkenpas et al. presented the use of a shear horizontal SAW

biosensor on langasite substrates [14]. They used biotin-modi-

ed rabbit IgG and goat antirabbit IgG antibodies as the analyte

and performed liquid medium experiments with phase change

analysis. A more recent similar work that employed the same

SAW substrate presented the detection of toxigenic E.Coli

O157:H7 bacteria with phase responses of 14 compared to 2

anti-TNP (trinitrophenyl) binding [15]. Gizeli et al. carried out

important investigations in liquid phase detection employing

Love wave devices [16], [17]. Maximum sensitivity of 430

cm g with a concentration range of 1 – 400 g

II. B IOSENSOR F UNDAMENTALS

In essence, a biosensor can also be dened as an analytical

device incorporating a biological sensing material with a trans-

ducer to detect and translate biological changes into known

electrical signals. There are three signicant components of

a biosensor: 1) sensing material; 2) interface material; and 3)

transducer. In order to successfully analyze the material under

test, all of these components should be functioning effectively.

There are several major factors that affect the performance of

a biosensor. The sensing material determines the selectivity of

the sensor. Therefore, this material should have high specicity

and afnity to the molecules under analysis. The objective for

this component is to produce a thin lm of biologically active

material on or near the transducer surface, which responds

only to the presence of one or a group of materials requiring

detection. The second component is the interface material. It

provides a medium for the transducer to come into contact with

the sensing materials by promoting direct contact and strong

binding. In biosensor technologies, gold nds a wide use as

an interface material. The third component in the system is

the transducer. This component takes the most interest and

determines the operational quality of the entire system, as it

is the primary component that takes the chemical/biological

interaction and converts it into electric signals.

III. T RANSDUCER D ESIGN

A. Mechanisms

was

There are several transducer mechanisms that are widely

used for decades in various biosensor applications. These

mechanisms and their corresponding methods can be summa-

rized as [21]: electrochemical: potentiometry, amperometry,

ISFET; electrical: surface conductivity, capacitance; thermal:

calorimetry, enzyme thermistor; magnetic: paramagnetism;

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optical: uorescence, surface plasmon resonance, scattering;

piezoelectric: QCM, SAW, SH/APM, Lamb wave, Love wave.

Each mechanism has their particular advantages/disadvan-

tages and optimal applications. Several physical transducers

are capable of measuring surface mass changes resulting from

the biological materials at the sensitive area. Although mostly

advanced optical systems are utilized, the piezoelectric and

acoustic devices present similar but signicantly less expensive

alternatives [22]. As demonstrated in the literature, the adsorp-

tion of biomolecules on functionalized surfaces turned out to be

one of the prominent applications of piezoelectric transducers.

Examples of these applications include the interaction of DNA

and RNA with complementary strands, specic recognition

of protein-ligands by immobilized receptors, the detection

of viruses, bacteria, and cells, as well as the development of

immunosensors. Piezoelectric transducers enable a label-free

detection of molecules. They are more than simple mass sen-

sors since the sensor response is also inuenced by interfacial

phenomena, viscoelastic properties of the biomaterial, surface

charges of adsorbed molecules, and surface roughness [23]. In

this research, the transduction mechanism that was employed

for the biosensor application is piezoelectricity.

Piezoelectric QCM has been the most widely used trans-

duction for biosensor applications in the past several decades.

However, there is an important advantage of SAW devices over

the QCM. They can be designed to operate at higher frequen-

cies and, therefore, the sensitivity of these devices to the same

amount of analyte is increased remarkably. The other important

limitation of the QCMs is that they use bulk acoustic waves

and, consequently, result in reduced sensitivities than SAW

devices which employ SAWs that are constantly in contact

with the biomaterial. Recent advances in SAW device devel-

opment and the unique features of the fabrication sequences

that are developed for the CMOS-compatible SAW devices

in our research prove that CMOS-SAW devices could replace

the existing technologies being employed for biosensing by

providing higher sensitivity levels and better performance.

Fig. 2. Schematic depiction and SEM snapshots for (a) after CMOS fabrica-

tion, (b) oxide removal, (c) piezoelectric deposition, (d) and pad frame denition

for the three-step postprocessing sequence.

B. CMOS-SAW Device Fabrication

The CMOS-SAW devices were fabricated in AMI Semicon-

ductor’s 0.5- 3-metal 2-poly technology [19], [20]. A three-

step postprocessing sequence is applied to the CMOS nished

dice. All three postprocessing steps are completely compatible

with CMOS and do not disturb any of the CMOS layers. Fig.

2 presents the three-step postprocessing sequence. The post-

processing steps consist of: 1) reactive ion etch (RIE) for the

IDT denition; 2) maskless ZnO RF magnetron sputtering for

piezoelectric material deposition; and 3) pad frame denition

through shadow mask photolithography. The details of these

steps and the pertaining characterization data are detailed in

our previous work [18]–[20]. CMOS-SAW devices employ de-

sign novelties, such as embedded heaters for temperature sta-

bility analysis, acoustic absorbers, and EM feedthrough con-

tacts. Fig. 2(d) presents a nished CMOS-SAW die. One of the

major challenges of the postprocessing sequence is to obtain

highly oriented perpendicular sidewalls between the IDT ngers

for higher performance. This challenge was addressed by using

RIE, which employs the top metal layers as masking for the

underlying oxides. This step provided highly oriented IDT n-

gers without any loss of sidewall or connection metal integrity.

Fig. 2(b) demonstrates the results of the RIE step. As can be

observed from the gures, complete elimination of internger

oxide layers was achieved. Fig. 2(b) also shows a closeup of

a single IDT after the RIE step is completed. The acoustic ab-

sorber, pad connections, and IDT ngers are also displayed.

Another challenge was faced in the postprocessing sequence

due to small-size die-level fabrication. After the zinc–oxide

(ZnO) deposition, the entire die is covered with the sputtered

ZnO. Fig. 2(c) shows the ZnO layer after the RF magnetron

sputtering is completed. The detailed process parameters for

the rst three steps are reported in our previous work. In order

to access the pads for bonding or probing purposes, a nal step

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TIGLI et al. : FABRICATION AND CHARACTERIZATION OF A SAW BIOSENSOR

Fig. 3. (a) Top view of a typical CMOS-SAW die after photoresist develop-

ment. Closeup of the pads after development showing excess PR (inset). (b)

Closeup of the pad frame showing a well-dened edge masking.

IV. I NTERFACE M ATERIAL

Gold (Au) is virtually the only material that is being used in

biomedical applications as the coating for transducers. This is

mainly due to its outstanding material properties. It has out-

standing resistance to tarnishing and corrosion, and has high

electrical and thermal conductivity. It is easy to work with, and

its high malleability and its amenability to plating in very thin

lms make it ideal for use in miniaturized components in elec-

tronics, medical, and other applications. Au is known for its re-

sistance to corrosion, so it does not oxidize like other metals and

it is chemically inert.

Au surfaces are very attractive candidates for self-assembly

due to their metallic nature, great nobility, and particular afnity

for sulphur. This aspect allows functionalization with thiols of

various types and adhesion to diverse organic molecules, which

are modied to contain a sulphur atom. These coatings, assem-

bled onto the gold surfaces, can serve as biosensors [24]. This

offers the prospect of preparing biomedical devices with inter-

esting functionalities. Besides the highly useful material prop-

erties and its attractive features for biological compatibilities,

gold provides additional advantages when used with SAW de-

vices. The piezoelectric material being employed in this work,

namely ZnO, shows a high binding to Au, which promotes effec-

tive isolation of the ZnO layer by reducing its reactive nature.

Moreover, Au thin lms on top of piezoelectric layers essen-

tially provide very strong waveguiding. Au has a relatively low

shear acoustic velocity and a high density when compared to the

ZnO layer. Therefore, it presents a very effective waveguiding

mechanism, which can be exploited for liquid phase sensing.

Gizeli et al. reported the use of Au layers in a thickness range of

10 nm and 100 nm to investigate its effectiveness in shorting the

electric eld without interfering with the propagation of Love

wave [25]. It was concluded that the Au layer does not reduce

the viscoelastic acoustic wave interaction in the three-layered

waveguide devices. Love wave devices present the most effec-

tive solution to the liquid phase-sensing problem of SAW de-

vices as evidenced in various biosensor research in the literature

[25]–[27]. The Love wave devices were shown theoretically and

experimentally to be the most sensitive acoustic structure for

liquid sensing. Therefore, using gold in this work also makes the

potential liquid phase biosensor applications possible and pro-

vides improved sensitivity even for the solid phase application

of the work presented here which employs dip and dry testing

techniques. Due to all of these important features, Au was em-

ployed in this paper as the interface material between the bio-

logical sensing medium and the transducer.

For the biosensor application on the CMOS-SAW devices,

additional postprocessing steps should be applied. After the fab-

rication steps that were detailed in Fig. 2 are completed, gold

deposition and patterning should be carried out. Fig. 4(a) de-

picts the nal step of gold deposition and the pertinent layers.

As can be clearly seen from Fig. 4(b), the pad frame is patterned

out completely and only the active sensor area is covered with

the Au layer. Thermal evaporation is used for this purpose. The

thermal evaporator employs resistive heating and very low pres-

sures to sublimate Au. A high degree of vacuum is achieved

TABLE I

P HOTOLITHOGRAPHY P ROCESS P ARAMETERS

of patterning and etching is required. The typical die area is 1.5

mm 1.5 mm and the custom designed pads are 100 m 100

m. In order to carry out patterning and etching, the dice were

mounted on microscopic glass slides to provide safe and easy

handling. The major problem encountered during the patterning

was the photoresist build up (edge bead) on the edges of the

die. Due to the small size of the die, excessive amounts of

photoresist were built up on the edges after spinning. This, in

turn, created a nonuniform distribution of the photoresist on

the areas where the pads are located. Thickness measurements

showed a 2–4- m PR buildup on the pad frame region when

the thickness on the areas closer to the center of dies was

measured to be 1 m. The primary parameters that affect the

build up are: 1) the spinner speed/acceleration; 2) the spin

time; 3) the photoresist used; 4) the exposure time under the

aligner, and, nally, 5) the time in the developer. After several

experiments, these parameters were optimized for the process

in hand. In order to completely remove the photoresist build

up covering the pad frame, a two-step process was carried out.

Table I summarizes the process parameters. In the rst step,

the dies were spun at 5000 r/min for a 40-s period. Then they

were exposed for 20 s under UV light. A 5:1 ( DI Water: Devel-

oper ) developer was used for 120 s in the rst step. Once this

step was nished, the excessive photoresist due to build up is

thinned down approximately to half its thickness. In the second

step, the same exposure and development times were used to

remove the photoresist on the edges completely. Microposit

351 Developer and Shipley 1818 photoresists were used. Fig. 3

shows the results of the photolithography steps. It is clearly

seen in Fig. 3(a) that there is excess photoresist accumulation

on the edges of the die close to the pad frame. This effect is

considerably alleviated by using the two-step process as can be

seen in Fig. 3(b).

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Fig. 4. (a) Postprocessed biosensor layers after Au evaporation. (b) Snapshot

that shows the GSG-congured probe tips in contact with the pads of a

CMOS-SAW biosensor that is patterned for an active gold surface.

through a combination of a mechanical roughing pump and a

high-vacuum pump. Airco Temescal CV-8 power supply with a

130-V input was used to apply current through the boat. The

chamber pressure of 1.1 was set and monitored

with an ion gauge. The thickness of the Au is aimed to be 1000

Å. A deposition rate of 330 Å/min was attained for the samples.

Multiple samples were deposited simultaneously to provide uni-

formity and repeatability between the samples.

Once the Au deposition is nalized, the samples were taken

through the patterning process. This process followed the same

mask and lithography that was detailed in [20] for the ZnO layer.

As for the nal step of etching, Metex auto strip gold etchant

solution was used. This etchant offers a highly selective etch at

a temperature range of 50–60 C. The samples were etched at

an average temperature of 55 in two steps to control the etch

rate. A total of a 1000-Å-thick Au layer was etched completely

in 1 min and 25 s which translates into an etch rate of 11.76

Å/min.

Since the quality of the Au thin lms plays a crucial role in

the overall sensitivity of the devices, thorough characterization

is required. As in the case of the ZnO piezoelectric layer, the

surface morphology, surface roughness, and the crystal orienta-

tion/structure are the subject of interest. To address these inter-

ests, optical microscopy, X-ray diffraction (XRD), and atomic-

force microscopy (AFM) analyses were carried out. For compar-

ative analysis, scans of a dummy die (Au-Si) and a patterned

die (Au-ZnO- -Si) were carried out. As can be seen from

Fig. 5(a), the Au lm shows its peak at 38.921, which

agrees with the tabulated data in the literature [38]. The same

peak is also obtained in the case of the patterned die. However,

due to highly expressed perpendicular ZnO crystallography, the

gold peak is suppressed in comparison. Fig. 5(b) shows the XRD

curve obtained for this case. As demonstrated in this curve, the

ZnO (002) peak at 34.84 presents much higher intensity

than the Au peak. This is expected due to the highly perpendic-

ular -axis orientation of ZnO. Multiple layers cause the inten-

sity values to vary when compared to a single layer. The ratio

of Au to ZnO thickness is 1/30, which results in overexpression

of ZnO peaks in this XRD curve.

Surface roughness plays an important role in the overall func-

tionality of the gold surface for biosensor applications. Perfectly

smooth surfaces with the lowest possible grain sizes are desir-

Fig. 5. (a) XRD curve for the Au thin lm on an Si dummy chip. Note the peak

38.921 for Au. (b) XRD curve for the Au thin lm on the ZnO/ /Si

chip. Note the much higher peak value of ZnO crystal (002) at 34.84

suppressing the Au peak.

able to promote strong binding and effective immobilization on

the Au thin lms. In order to analyze the quality of the Au sur-

face after evaporation, AFM analyses were carried out. Fig. 6

presents the results for the surface roughness and grain size

analyses. As shown in the surface roughness prole of the sur-

face in Fig. 6(c), the Au surface presents an average grain size

of 3.66 nm with a maximum of 6.63 nm and a minimum of

1.23 nm. As evidenced by these gures, the thermally evapo-

rated Au thin-lm surface shows an extremely smooth surface

with very low grain sizes. This aspect plays an important role in

dening and applying the immobilization technique.

V. B IOLOGICAL S ENSING M ATERIAL

Piezoelectric immunosensors are accepted to be one of the

most sensitive analytical instruments, being capable of detecting

antigens in the picogram range [28]. Moreover, they have the

potential to detect antigens in the gas phase as well as in the

liquid phase. In order to achieve this sort of high sensitivity, the

sensing material of the biosensor system plays a very important

role. There are several conditions that must be satised for such

a system [21]: 1) the biological component must retain substan-

tial biological activity when attached to the sensor surface; 2)

the biological lm must remain associated with the sensor sur-

face while retaining its structure and function; 3) the immobi-

lized biological lm needs to have long-term stability and dura-

bility; and 4) the biological material needs to have a high degree

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