Surface acoustic wave (SAW) devices have been used in biochemical assays due to their high sensitivity. The device sensitivity is a function of changes in the density and viscosity of the liquid. Here, we studied the effect of fluid viscosity using a 250 MHz quartz shear-horizontal (SH)-SAW biosensor by monitoring different concentrations of binary aqueous/glycerol solutions. In this study, the sensitivity of the biosensor was determined by fitting the data to models derived from perturbation theory. Measurements in water were used as the reference. For a 0% to 50% glycerol solution, an 87°–204° separation in the phase shift was observed. The slope of the plot of the phase shift vs (ηρ)0.5 was used to indicate the sensor’s sensitivity. The sensitivity for our 250 MHz quartz SH-SAW sensors was calculated to be 3.7×10−3m2sKg⁠. The corresponding mass sensitivity was determined to be 9.25 × 105 m2Kg⁠. The limit of detection was calculated to be 36 picograms (pg), while the limit of quantification or LOQ was calculated to be 109 pg. Traditionally, liquid phase measurements have been challenging for SAW devices because liquids dampen the vibrating sensors severely. This problem has been largely solved using a transverse (shear) wave instead of the more popular longitudinal or Rayleigh waves. Liquid measurements are now possible using transverse waves, also known as shear waves, because transverse waves are only minimally attenuated by liquids. Shear-horizontal SAW sensors (SH-SAW) show great promise as label-free biosensors because of their ability to handle liquid samples. However, the viscosity of the liquid still induces loading effects and can be measured when the liquid is loaded onto the SH-SAW propagating surface (delay line). When the liquid above the delay line is perturbed by physical or chemical changes, such as binding to a receptor, it alters the propagating acoustic wave. The SH-SAW device can measure these changes in liquid properties as a change in the wave’s phase compared to the original wave. The device’s phase shift was recorded as a function of the changes in the density and viscosity of the binary glycerol solution and used to determine the sensitivity in the linear dynamic range of responses.

Topics
Surface acoustic waves, MEMS technology, MEMS devices, Sensors, Delay lines, Telecommunications, Detection limit, Medical diagnosis, Pathology, Biosensors

Biosensors are devices that can measure the concentration of biological or medically related agents at concentrations that offer information on contamination, biological identity, medical diagnosis, or disease prognosis.1 Surface acoustic wave (SAW) sensors have been widely used in several fields, such as consumer electronics and wireless communication, primarily radio frequency (RF) filters.2–6 Biosensors that use SAW technology, incorporate an acoustic transducer that generates a wave whose propagation velocity changes with mass or viscosity loading.7 The change in the wave velocity of surface acoustic waves at the sensor-media interface to monitor changes in the biochemical system under observation.8 Our SAW biosensor is capable of detecting perturbation at the solid–liquid interface induced by the binding of a receptor to a complimentary ligand in liquid samples.9 The resulting perturbation is caused by a change in the physical properties of the sensor surface due to the interaction of the targeted chemical species.10 The change in the physical property will alter the local viscosity and, therefore, the elasticity modulus of the layer just above the surface in the region where the SAW propagates.11 Minute changes in the viscoelastic properties will be measured and recorded as a function of time.12 The effects of the changes in these physical properties are recorded as changes in SAW amplitude attenuation and SAW wave velocity.13 We can then apply these biosensors to rapidly detect viral, bacterial, or fungal pathogens in complex biological matrices.14 Unfortunately, liquid samples severely dampen longitudinal surface acoustic waves on contact, as seen in quartz crystal microbalances and other piezoelectric sensors that use Rayleigh (longitudinal) waves.15 For this reason, shear-horizontal surface acoustic wave (SH-SAW) devices have been developed and are now being utilized for liquid sample measurements.16 

In our applications, we primarily focus on examining the phase shift rather than focus on the attenuation. The reason for this is the attenuation or insertion loss in the amplitude since the insertion losses are also a function of the piezoelectric material used for the fabrication of the sensor and not strictly a function of the concentration of the analyte.10 This is because the phase shift gives the device’s response to viscous liquid loading and mass loading.17 This gives good biosensor performance and enhances the robustness of the SH-SAW sensors, enabling them to be applied in medical diagnostics, environmental monitoring, food safety, or sterility testing.18,19 However, one essential parameter that needs to be well understood is the interactions of viscous fluids with the sensor. This is because viscous loading induces a phase shift with a simultaneous increase in the insertion loss. In this report, we will explore the viscous loading response of a gold-coated, reflection-type, 250 MHz quartz SH-SAW biosensor.

Some of the early work performed to investigate the analysis of liquid samples using quartz SAW devices was the work performed by Kanazawa and Gordon to explore the frequency shift of bulk wave resonators that were in contact with viscous liquids.20 In each case, the fundamental equation is that there exists a relationship, for a SAW sensor with a layered sensing surface (delay line), between the wave propagation velocity and the frequency of the wave,21 
(1)
They reported that of a Newtonian fluid, the relative frequency and corresponding phase shift are proportional to the absolute viscosity and density of the liquid.20 The induced phase shift caused by an increase in the liquid viscosity can be derived from the following equation:22,
(2)
where Δv is the change in the propagation velocity, v0 is the unperturbed wave velocity, Δφ is the change in the phase, and φ0 is the unperturbed phase. Perturbation theory was used by Auld to describe the acoustic waves with a mechanical perturbation.23 This theory can be applied to a transverse wave at a sensor surface and be described using the complex perturbation formula as follows: 24 
(3)
where α is the dampening coefficient term, β is the propagation constant, and i is the complex number −1⁠.
The change in the perturbation is written as follows: 24 
(4)
where P is the average power flow per unit width along the unperturbed sensor surface and Z is the surface impedance.25 The grouping v24P is proportional to the parameter β. The relative change in the propagation constant.
The relative change in the propagation constant due to mechanical perturbations is given as follows: 24 
(5)
To determine the surface impedance of the liquid just above the sensor, we have to find the stress response of the liquid due to the propagation of the shear wave. The surface impedance is given as follows:
(6)
As we can see that for a Quartz sensor where v ≥ 5000 ms⁠, ρ ≥ 1000 Kgm3⁠, η ≤ 1000 mPa s (cP), and the angular frequency, ω = 2πf, where f is the center frequency of the device. The equation under the second radical, under these conditions, approaches 1. We can then reduce the equation to the following equation:
(7)
If we also assume that the dampening of the wave is negligible, that is, Δα = 0, then consequently Eq. (4) would simplify to the following equation:
(8)
We can also express changes in surface perturbation as seen in Eq. (8). As a result, we obtain the following equation:
(9)
where S is the sensitivity of the device and A is the area of the device. For Newtonian fluids, both the damping per wavelength and the relative change of the propagation constants are proportional to the following equation:24 
(10)
Consequently, the change in phase can also be represented by the change in surface perturbation, as seen in the following equation:
(11)

A gold-coated quartz SH-SAW biosensor, with a central frequency centered at 250 MHz, was purchased from TST Biomedical Electronics Company and used without further modifications. The SH-SAW sensor has a corresponding phase velocity of v0 = 5000 m s−1 and is given by Eq. (1). This gives an acoustic wavelength of λ = 20 microns. The SAW sensor employed on our biosensor platform is a dual delay line SAW sensor that was developed by the Japan Radio Company (JRC), which later all assets transferred to TST Biomedical Electronic Company, Taiwan.15,26 The TST Biomedical Electric Company fabricates the SH-SAW device from quartz.26 Although quartz has lower electro-mechanical coupling than some sensor templates, quartz does have the unique advantage of having a near-zero, temperature coefficient.10 The major significance of this is that quartz sensors are not affected by thermal drifts as significantly as other SAW devices.26 Therefore, these devices can be used without having significant insulation around the devices or having to use thermal compensation to correct the data. The TST SAW sensors are coated with a 100 nm thin film of gold. The gold layer on the sensor plays several different roles including shielding the sensor from interferences from other devices, protecting the sensor from harsh chemicals, and enabling chemical functionalization of the sensor via gold–thiol chemistry.27 The gold layer also shorts any electrical effects. This sensor uses a single interdigital transducer (IDT) combined with a reflector to launch and receive RF pulses. The dual delay lines were identical and encircled by a wall of the photo-resist polymer, SU-8.27 The SU-8 well allows 5 µl of the sample to be incubated.26 

To perform our sensitivity measurements, we utilized three gold-coated SH-SAW sensors that were provided by TST Biomedical Electronics, Company Limited, Taiwan, ROC. The sensors were observed using a stereo microscope to look for any obvious defects. The sensors were then cleaned with an air plasma cleaner for 60 s on the high settling. The sensor was then rinsed with distilled water.

The liquid viscosity measurements for the glycerol mixtures were obtained using a C. Goldenwall NDJ-5S Rotary digital rotational viscosity meter. The viscosity meter is capable of measuring liquid viscosities over the range of 10–100 000 mPa s.

The mass sensitivity of a 250 MHz quartz SH-SAW sensor is highest at the center frequency. The delay line has a length of 200λ, which corresponds to a length of 1.8 mm.27 An image of the SH-SAW is shown in Fig. 1. The glycerol–water mixtures were pipetted into the Su-8 micro-wells, ranging in concentration from 0% to 50%.

FIG. 1.
Top view image of the dual channel (delay line) SH-SAW sensor from TST Biomedical Electronic Company, Taiwan (ROC).
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Top view image of the dual channel (delay line) SH-SAW sensor from TST Biomedical Electronic Company, Taiwan (ROC).

FIG. 1.
Top view image of the dual channel (delay line) SH-SAW sensor from TST Biomedical Electronic Company, Taiwan (ROC).
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Top view image of the dual channel (delay line) SH-SAW sensor from TST Biomedical Electronic Company, Taiwan (ROC).

Close modal
The propagational losses per wavelength ΔPL can be expressed as the difference between the insertion losses from the liquid sample and the sensor before the addition of the sample,28,
(12)
where ILsample and ILSystem are the insertion losses associated with the sample and the unperturbed system, respectively. The value for Δx is the distance between the input and output IDTs.12 The measured values for ΔPL of the quartz SH-SAW biosensor, under the loading of the glycerol solutions, are shown in Fig. 2. The propagational loss of the biosensor increases proportionally with the square root of the viscosity–density product of liquids.28 
FIG. 2.
Measured propagational losses of the 250 MHz quartz surface acoustic wave sensor. We see that the increase in the insertion losses is linearly increasing.
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Measured propagational losses of the 250 MHz quartz surface acoustic wave sensor. We see that the increase in the insertion losses is linearly increasing.

FIG. 2.
Measured propagational losses of the 250 MHz quartz surface acoustic wave sensor. We see that the increase in the insertion losses is linearly increasing.
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Measured propagational losses of the 250 MHz quartz surface acoustic wave sensor. We see that the increase in the insertion losses is linearly increasing.

Close modal

The binary glycerol–water mixtures were pipetted into the sample wells for analysis, as shown in Fig. 1. The resulting sensor response is plotted in Fig. 2.

The plot of the propagational losses vs the square root of the density–viscosity product gave a linearly increasing relation over the entire viscosity range tested. This is in good agreement with the theory and suggests that the insertion increases linearly with viscosity loading. The plot goes directly through the origin, which suggests that the sensor response is close to that of an ideal sensor. The theoretical propagation losses are given by the following equation:
(13)
where S is the sensor’s sensitivity and the mass sensitivity is defined as Eq. (13). The viscosities of the glycerol mixtures cause the sensors to be loaded with a thin liquid film of glycerol. This thin film is the source of the decay in the wave’s amplitude, which we record as the propagational losses or the insertion losses. The characteristic film depth, d, of the surface of the sensor, is given by the following equation with the resulting value tabulated in Table I:29 
(14)
TABLE I.

The values for the film thickness on the glycerol film on the surface of the SH-SAW sensor. The glycerol solution and the corresponding phase shift were recorded.

Glycerol soln
(% glycerol)Decay depth (nm)Phase shift
1 89.55 87.52 
5 95.10 99.10 
10 101.21 112.81 
20 116.16 129.92 
30 136.45 156.63 
40 165.04 180.66 
50 207.10 204.74 
Glycerol soln
(% glycerol)Decay depth (nm)Phase shift
1 89.55 87.52 
5 95.10 99.10 
10 101.21 112.81 
20 116.16 129.92 
30 136.45 156.63 
40 165.04 180.66 
50 207.10 204.74 
View Large
The viscosity-induced phase velocity shift is directly proportional to the changes in the wave velocity at the surface of the sensor and can be derived from the following equation:22 
(15)
where the phase velocity shift of the device is the unperturbed phase velocity, and the phase shift measured by the biosensor reader is the unperturbed phase between input and output IDTs of the sensor. The biosensor under viscous liquid loading, the theoretical relative phase velocity shift, was found to be proportional to the square root of the viscosity–density product, as described in the following equation and the plot of the phase shift vs the square root of the viscosity–density product is shown in Fig. 3:
(16)
FIG. 3.
The relative phase velocity shift of the 250 MHz quartz SH-SAW sensor with mass-loading from a series of binary glycerol–water solutions. (a) Plot of phase responses to glycerol mixtures from 0% to 50% glycerol–water mixtures. (b) Plot of phase responses to glycerol mixtures from 0% to 30% glycerol–water mixtures, and the linear dynamic range used to calculate the sensitivity.
View largeDownload slide

The relative phase velocity shift of the 250 MHz quartz SH-SAW sensor with mass-loading from a series of binary glycerol–water solutions. (a) Plot of phase responses to glycerol mixtures from 0% to 50% glycerol–water mixtures. (b) Plot of phase responses to glycerol mixtures from 0% to 30% glycerol–water mixtures, and the linear dynamic range used to calculate the sensitivity.

FIG. 3.
The relative phase velocity shift of the 250 MHz quartz SH-SAW sensor with mass-loading from a series of binary glycerol–water solutions. (a) Plot of phase responses to glycerol mixtures from 0% to 50% glycerol–water mixtures. (b) Plot of phase responses to glycerol mixtures from 0% to 30% glycerol–water mixtures, and the linear dynamic range used to calculate the sensitivity.
View largeDownload slide

The relative phase velocity shift of the 250 MHz quartz SH-SAW sensor with mass-loading from a series of binary glycerol–water solutions. (a) Plot of phase responses to glycerol mixtures from 0% to 50% glycerol–water mixtures. (b) Plot of phase responses to glycerol mixtures from 0% to 30% glycerol–water mixtures, and the linear dynamic range used to calculate the sensitivity.

Close modal
The plot of phase shift vs the square root of the viscosity–density product demonstrates a linear dynamic range in the region from 0% to 30% glycerol, as shown in Fig. 3(b). Beyond 30% glycerol, the response becomes highly non-linear. If we focus on the linear dynamic range and employ linear regression, we obtain the sensitivity (S) of the device, which is S=3.70×10−3m2sKg⁠. This is valid over a viscosity range between 1.0–2.5 mPa s. The mass sensitivity (Sm) is the sensitivity multiplied by the angular frequency, which gives the mass sensitivity of Sm=9.25×105m2Kg or Sm=9.25×108mm2g⁠. To calculate the theoretical lower limit of the sensor, we multiply by the smallest possible phase shift from the reader of 0.001°. The result is a theoretical lower limit of 1.0 pg based on a 1.01 mm2 area of the delay line. Therefore, this is the maximum device’s resolution, the smallest measurement that can be theoretically resolved on the device. The upper limit of the device was estimated to be associated with a 360° phase shift and a phase shift of ∼10°. The resulting value for the upper limit was 10 μg. When compared to similar a quartz SAW device, our device had a much high mass sensitivity of Sm=9.25×109mm2g compared to Sm=5.0×103mm2g for a 180 MHz quartz SAW sensor.24 A similar 255 MHz aluminum nitride (AIN)/sapphire SAW device had a mass sensitivity of 21 mm2g⁠, which was much lower than reported quartz devices. The 250 MHz quartz SAW sensor used in the project was most similar to a Love-wave 330 MHz lithium tantalate (LiTaO3) sensor with a 500 nm silicon dioxide guided layer, as reported by the Sandia National Laboratory. The report suggested that the mass sensitivity of the sensor was 4.31×109mm2g⁠. The lithium tantalate device had a higher mass sensitivity; however, this is largely due to the much higher electro-mechanical couple of lithium tantalate compared to quartz. The area of the quartz SH-SAW sensor is (0.56 × 1.8 mm2) or 1.01 mm2. If we divide the area of the sensor by the mass sensitivity, we get a theoretical mass of 1 × 10−10. The device is able to measure as little as 100 pg on the sensor surface based on the experimentally derived data. Experimentally, we determined that the standard deviation of the baseline measurement (the blank) for the SH-SAW was 0.000 49°. Based on this standard deviation in the signal, we could calculate the limit of detection, as described in the following equation:
(17)
where Sm is the mass sensitivity of the system. This gives 9.25×105m2Kg⁠. Using the standard deviation of 0.01°, we calculate a practical resolution of 36 × 10−12 g. The limit of detection (LOD) may be difficult to achieve if the instrument has a high electronic jitter or other parasitic electrical noise. For this reason, we calculated the limit of quantification (LOQ) as a more practical measurement of the limit for mass detection on the SH-SAW device. The LOQ is given in the following equation:
(18)

The LOQ was calculated to be 109 × 10−12 g (∼110 pg). This suggests that the device can precisely make measurements for analytes in increments of 110 pg.

In this article, we used a 0%–50% glycerol solution to determine the overall mass sensitivity of the 250 MHz, gold-coated quartz, reflection mode SH-SAW biosensor. The sensitivity was determined to be 3.7×10−3m2sKg⁠. The mass sensitivity was determined to be 9.25×105m2Kg⁠.

The reported values for high-sensitivity lithium tantalate SH-SAW biosensors operating at 330 MHz sensor were determined to be 4.31×109mm2g⁠. The higher mass sensitivity is expected due to the high electro-mechanical coupling of lithium tantalate when compared to quartz at ∼5% compared to ∼0.15%; however, the quartz sensors have a much smaller thermal coefficient. This results in the quartz devices being able to operate at room temperature without the need for significant insulation from the environment.

The major limitation of the 250 MHz quartz SH-SAW biosensor, compared to other high-sensitivity lithium tantalate sensors, seems to be that the unperturbed delay line insertion losses are nearly five times higher for the quartz device. The gold-coated quartz device that was used, had an unperturbed insertion loss of −27.8 dB, while the higher sensitivity lithium tantalate device reported in the literature had an unperturbed insertion loss of less than −10.0 dBm. We ascribe the much higher insertion losses of our device to the differences in material properties between quartz and lithium tantalate. The electromechanical coupling for quartz is 0.15% vs 5.0% for lithium tantalate, which is 33.33 times better at converting electrical potential to a mechanical wave.

We hypothesize that the dramatic difference in energy conversion leads to significant increases in propagational losses when the wave is excited, propagates from the IDT, and travels across the delay line. The quartz devices also operate at a frequency that is 80 MHz lower frequency the operating frequency of the lithium tantalate device, which also improves the overall signal-to-noise ratio.

It was found that the 250 MHz SH-SAW biochip has similar sensitivities when compared to the lithium tantalate SH-SAW sensor for assays where the biochip is analyzer moist but not submerged. The net result is that while the lithium tantalate device is more sensitive, the signal-to-noise ratio of the 250 MHz quartz device has a similar signal-to-noise ratio. Another positive attribute of the quartz sensor is that it is fabricated with a more stable sample handling system. Overall, the properties of quartz are highly desirable for real-world conditions. Therefore, the 250 MHz quartz SH-SAW biosensor is suited for biomedical applications for samples with viscosities that fall between 1.0–2.5 mPa s.

The authors acknowledged TST Biomedical Electric Company Limited for providing the iprotin reader and 250 MHz SH-SAW quartz sensors. Funding for this work was provided by the United States National Science Foundation (NSF) (Grant No. 1817282) and the United States Department of Energy (DOE) (Grant No. DE-EM0005266).

The authors have no conflicts to disclose.

Ololade Adetula: Data curation (equal); Investigation (equal); Visualization (equal). Aimofumhe Eshiobomhe Sigmus: Data curation (equal); Investigation (equal). Favour Badewole: Formal analysis (equal); Investigation (equal). Collins Ijale: Data curation (equal). Marlon Thomas: Conceptualization (lead); Data curation (lead); Formal analysis (lead); Funding acquisition (lead); Investigation (lead); Methodology (lead); Project administration (lead); Supervision (lead); Validation (lead); Visualization (lead); Writing – original draft (lead); Writing – review & editing (lead).

Data for these experiments can be provided upon request. Experimental data is stored in the office of the lead author, Dr. Marlon Thomas at Fisk University in Nashville, TN; United States of America.

1.
I.
Plikusiene
 and
A.
Ramanaviciene
, “
Investigation of biomolecule interactions: Optical-, electrochemical-, and acoustic-based biosensors
,”
Biosensors
13
(
2
),
292
(
2023
).
https://doi.org/10.3390/bios13020292
Google Scholar
Crossref
Search ADS
PubMed
 
2.
W.
Akstaller
,
R.
Weigel
, and
A.
Hagelauer
, “
Electrothermal modeling of surface acoustic wave resonators and filters
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
67
(
11
),
2423
–
2432
(
2020
).
https://doi.org/10.1109/tuffc.2020.2999545
Google Scholar
Crossref
Search ADS
PubMed
 
3.
R. L.
Baer
, “
Phase noise in surface-acoustic-wave filters and resonators
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
35
(
3
),
421
–
425
(
1988
).
https://doi.org/10.1109/58.20463
Google Scholar
Crossref
Search ADS
PubMed
 
4.
R. W.
Cohn
 and
R. S.
Wagers
, “
Multiparameter tuning of surface-acoustic-wave filters
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
37
(
3
),
255
–
264
(
1990
).
https://doi.org/10.1109/58.55316
Google Scholar
Crossref
Search ADS
PubMed
 
5.
S.
Yang
,
Y.
Ai
,
Z.
Cheng
,
L.
Zhang
,
L.
Jia
,
B.
Dong
,
B.
Zhang
,
J.
Wang
, and
Y.
Zhang
, “
Method of the out-of-band rejection improvement of the AlN based surface acoustic wave filters
,”
Ultrasonics
91
,
30
–
33
(
2019
).
https://doi.org/10.1016/j.ultras.2018.07.008
Google Scholar
Crossref
Search ADS
PubMed
 
6.
C.
Campbell
,
Surface Acoustic Wave Devices for Mobile and Wireless Communications
(
Academic Press
,
San Diego
,
1998
), p
xxvi
.
Google Scholar
 
7.
J.
Kondoh
, “
A liquid-phase sensor using shear horizontal surface acoustic wave devices
,”
Electron. Commun. Jpn.
96
(
2
),
41
–
49
(
2013
).
https://doi.org/10.1002/ecj.10407
Google Scholar
Crossref
Search ADS
 
8.
Z.
Li
,
Y.
Jones
,
J.
Hossenlopp
,
R.
Cernosek
, and
F.
Josse
, “
Analysis of liquid-phase chemical detection using guided shear horizontal-surface acoustic wave sensors
,”
Anal. Chem.
77
(
14
),
4595
–
4603
(
2005
).
https://doi.org/10.1021/ac0504621
Google Scholar
Crossref
Search ADS
PubMed
 
9.
F.
Josse
,
F.
Bender
, and
R. W.
Cernosek
, “
Guided shear horizontal surface acoustic wave sensors for chemical and biochemical detection in liquids
,”
Anal. Chem.
73
(
24
),
5937
–
5944
(
2001
).
https://doi.org/10.1021/ac010859e
Google Scholar
Crossref
Search ADS
PubMed
 
10.
D. W.
Branch
,
T. L.
Edwards
,
D.
Huber
, and
S. M.
Brozik
, Shear Horizontal Surface Acoustic Wave Microsensor for Class A Viral and Bacterial Detection,
Sandia National Laboratories, Issued by Sandia National Laboratories, operated for the United States Department of Energy by Sandia Corporation
,
2008
.
11.
F.
Bender
,
R. E.
Mohler
,
A. J.
Ricco
, and
F.
Josse
, “
Analysis of binary mixtures of aqueous aromatic hydrocarbons with low-phase-noise shear-horizontal surface acoustic wave sensors using multielectrode transducer designs
,”
Anal. Chem.
86
(
22
),
11464
–
11471
(
2014
).
https://doi.org/10.1021/ac503701b
Google Scholar
Crossref
Search ADS
PubMed
 
12.
G.
Hu
,
J.
Xu
,
G. W.
Auner
,
J.
Smolinski
, and
H.
Ying
, “
Viscosity response of shear horizontal surface acoustic wave on AlN/sapphire structure
,”
Electron. Lett.
43
(
18
),
1006
(
2013
).
https://doi.org/10.1049/el:20070769
Google Scholar
Crossref
Search ADS
 
13.
E.
Gizeli
, “
Study of the sensitivity of the acoustic waveguide sensor
,”
Anal. Chem.
72
(
24
),
5967
–
5972
(
2000
).
https://doi.org/10.1021/ac000694u
Google Scholar
Crossref
Search ADS
PubMed
 
14.
E.
Gizeli
,
F.
Bender
,
A.
Rasmusson
,
K.
Saha
,
F.
Josse
, and
R.
Cernosek
, “
Sensitivity of the acoustic waveguide biosensor to protein binding as a function of the waveguide properties
,”
Biosens. Bioelectron.
18
(
11
),
1399
–
1406
(
2003
).
https://doi.org/10.1016/s0956-5663(03)00080-0
Google Scholar
Crossref
Search ADS
PubMed
 
15.
T.
Kogai
,
H.
Yatsuda
, and
J.
Kondoh
, “
Rayleigh SAW-assisted SH-SAW immunosensor on X-cut 148-Y LiTaO3
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
64
(
9
),
1375
–
1381
(
2017
).
https://doi.org/10.1109/tuffc.2017.2734282
Google Scholar
Crossref
Search ADS
PubMed
 
16.
F.
Martin
,
M. I.
Newton
,
G.
McHale
,
K. A.
Melzak
, and
E.
Gizeli
, “
Pulse mode shear horizontal-surface acoustic wave (SH-SAW) system for liquid based sensing applications
,”
Biosens. Bioelectron.
19
(
6
),
627
–
632
(
2004
).
https://doi.org/10.1016/s0956-5663(03)00257-4
Google Scholar
Crossref
Search ADS
PubMed
 
17.
M. S.
Brugger
,
L. G.
Schnitzler
,
T.
Nieberle
,
A.
Wixforth
, and
C.
Westerhausen
, “
Shear-horizontal surface acoustic wave sensor for non-invasive monitoring of dynamic cell spreading and attachment in wound healing assays
,”
Biosens. Bioelectron.
173
,
112807
(
2021
).
https://doi.org/10.1016/j.bios.2020.112807
Google Scholar
Crossref
Search ADS
 
18.
M.
Csete
 and
W. D.
Hunt
, “
Potential of surface acoustic wave biosensors for early sepsis diagnosis
,”
J. Clin. Monit. Comput.
27
(
4
),
427
–
431
(
2013
).
https://doi.org/10.1007/s10877-013-9446-5
Google Scholar
Crossref
Search ADS
PubMed
 
19.
P.
Gaehtgens
,
H. J.
Meiselman
, and
H.
Wayland
, “
Velocity profiles of human blood at normal and reduced hematocrit in glass tubes up to 130 μ diameter
,”
Microvasc. Res.
2
(
1
),
13
–
23
(
1970
).
https://doi.org/10.1016/0026-2862(70)90049-x
Google Scholar
Crossref
Search ADS
PubMed
 
20.
K. K.
Kanazawa
 and
J. G.
Gordon
, II, “
The oscillation frequency of a quartz resonator in contact with a liquid
,”
Anal. Chim. Acta
175
,
99
–
105
(
1985
).
https://doi.org/10.1016/S0003-2670(00)82721-X
Google Scholar
Crossref
Search ADS
 
21.
E.
Danicki
,
A.
Nowicki
, and
Y.
Tasinkevych
, “
Correspondence: Interdigitated interdigital transducer for surface elastometry of soft damping tissue
,”
IEEE Trans. Ultrason. Ferroelectr. Freq. Control
60
(
6
),
1260
–
1262
(
2013
).
https://doi.org/10.1109/tuffc.2013.2690
Google Scholar
Crossref
Search ADS
PubMed
 
22.
D. S.
Ballantine
,
R. M.
White
,
S. J.
Marti
,
A. J.
Ricco
,
E. T.
Zeller
,
H.
Frye
, and
H.
Wohltjen
,
Acoustic Wave Sensors: Theory, Design and Physico-Chemical Applications
(
Academic Press
,
San Diego
,
1997
).
Google Scholar
Crossref
Search ADS
 
23.
B. A.
Auld
,
Acoustic Fields and Waves in Solids
(
Wiley
,
New York
,
1973
).
Google Scholar
 
24.
G.
Kovacs
,
M.
Vellekoop
,
R.
Haueis
,
G. W.
Lubking
, and
A.
Venema
, “
A love wave sensor for (bio)chemical sensing in liquids
,”
Sens. Actuators, A
43
,
38
–
43
(
1994
).
https://doi.org/10.1016/0924-4247(93)00660-v
Google Scholar
Crossref
Search ADS
 
25.
J.
Kondoh
 and
S.
Shiokawa
, “
A liquid sensor based on a shear horizontal SAW device
,”
Electron. Commun. Jpn.
76
(
2
),
69
–
82
(
1993
).
https://doi.org/10.1002/ecjb.4420760208
Google Scholar
Crossref
Search ADS
 
26.
Y. C.
Peng
,
C. H.
Cheng
,
H.
Yatsuda
,
S. H.
Liu
,
S. J.
Liu
,
T.
Kogai
,
C. Y.
Kuo
, and
R. Y. L.
Wang
, “
A novel rapid test to detect anti-SARS-CoV-2 N protein IgG based on shear horizontal surface acoustic wave (SH-SAW)
,”
Diagnostics
11
(
10
),
1838
(
2021
).
https://doi.org/10.3390/diagnostics11101838
Google Scholar
Crossref
Search ADS
PubMed
 
27.
K.
Kano
,
H.
Yatsuda
, and
J.
Kondoh
, “
Evaluation of shear horizontal surface acoustic wave biosensors using “layer parameter” obtained from sensor responses during immunoreaction
,”
Sensors
21
(
14
),
4924
(
2021
).
https://doi.org/10.3390/s21144924
Google Scholar
Crossref
Search ADS
PubMed
 
28.
F. H.
Herrmann
,
D.
Hahn
, and
S.
Buttgenbach
, “
Separate determination of liquid density and viscosity with sagittally corrugated love-mode sensors
,”
Sens. Actuators, A
78
,
99
–
107
(
1999
).
https://doi.org/10.1016/s0924-4247(99)00224-1
Google Scholar
Crossref
Search ADS
 
29.
S. J.
Martin
,
A.
Ricco
,
T. M.
Niemczyk
, and
G. C.
Frye
, “
Characterization of SH acoustic plate mode liquid sensors
,”
Sens. Actuators
20
,
253
–
268
(
1989
).
https://doi.org/10.1016/0250-6874(89)80124-6
Google Scholar
Crossref
Search ADS
 
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