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Sensors and Materials, Vol. 35, No. 10 (2023) 4797–4808
MYU Tokyo
S & M 3432

C-reactive Protein Detection Using an Ion-sensitive Field-effectTransistor (ISFET)-based Aptasensor with a Chemically Modified
Gate Surface for Improved Sensitivity
Siti Masturah binti Fakhruddin,1 Kumi Y. Inoue,2*Masayoshi Esashi,3 and Hitoshi Shiku4
1Graduate School of Environmental Studies, Tohoku University, Sendai 980-8579, Japan

2Center for Basic Education, Faculty of Engineering, University of Yamanashi, Kofu 400-8511, Japan
3Micro System Integration Center (μSIC), Tohoku University, Sendai 980-0845, Japan
4Graduate School of Engineering, Tohoku University, Sendai 980-8579, Japan(Received June 30, 2023; accepted October 12, 2023)

Keywords: aptamer, ion-sensitive field-effect transistor (ISFET), C-reactive protein, silane coupling, pH
sensitivity

C-reactive protein (CRP) is an inflammation biomarker that requires simple and real-timemonitoring for accurate diagnosis. Conventional CRP tests are complicated, expensive, andtime-consuming. Field-effect transistor (FET)-based affinity sensors are seen as the ideal
solution but it is difficult to obtain FET with sensitive gate structures. In this work, we propose a
simple method of chemically modifying the gate surface of a commercial ion-sensitive FET(ISFET) with (3-glycidyloxypropyl)trimethoxysilane (GPTMS) to lower the background noise
signal and then immobilize aptamers that provide significant surface potential change when theybind to CRP. The FET aptasensor was able to measure 0.002–20 μg/mL CRP in 1 × phosphatebuffered saline (PBS) with a higher sensitivity than the nonmodified ISFET sensors with their
original pH sensitivity and was on par with other FET sensors without needing expensive
nanomaterial or complicated nanofabrication.

1.

Introduction

C-reactive protein (CRP) is a pentameric plasma protein produced by the liver and can befound in blood and certain other biofluids. It is an acute-phase protein and nonspecific biomarker
that can indicate the levels of inflammation and tissue damage in the body. Monitoring CRP
levels can be useful for assessing cardiovascular risks or even COVID-19 infection severity(1 ng/mL for very mild cases to more than 100 ng/mL).(1) Plasma levels of CRP in humans canrise rapidly more than 1000-fold after an acute-phase stimulus, and the synthesis rate peaks
around 48 h.(2,3) Furthermore, the plasma half-life of CRP is about 19 h, making it difficult tomonitor accurately in patients. This makes it necessary to have a sensor that can measure thereal-time and long-term CRP levels of patients to obtain accurate diagnostic results. The real*Corresponding author: e-mail: inokumi@yamanashi.ac.jp

https://doi.org/10.18494/SAM4570

ISSN 0914-4935 © MYU K.K.
https://myukk.org/

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Sensors and Materials, Vol. 35, No. 10 (2023)

time monitoring of CRP levels in patients is also vital for medical events such as cytokine storms
that occur in COVID-19 patients.(4–6) CRP level has even been found to be predictive of survival
rate among high-risk COVID-19 patients.(7)Conventional methods of measuring CRP levels in blood serum include enzyme-linked
immunoassay (ELISA), high-sensitivity assay, and cardiac assay, with limits of detection
varying from 0.03 to 0.2 μg/mL and detection ranges up to 10 μg/mL.(8) These methods requirecomplicated labeling materials, multistep processes, expensive equipment, and experienced
personnel. To overcome these limitations, there have been multiple different approaches indeveloping CRP sensors such as fluorescence,(9) surface plasmon resonance (SPR),(10,11)
electrochemical methods(12,13) or even acoustic profiling/quartz crystal microbalance.(14)However, these methods still have disadvantages such as requiring fluorescent labels,
complicated fabrication process, and expensive equipment with skilled personnel for operation.
They are also incapable of real-time measurements. Thus, field-effect transistor (FET)-based
sensors that offer rapid and label-free direct electrical detection, inexpensive mass production
with easy miniaturization, and relative ease of use are gaining interest to realize the need for a
CRP sensor capable of point-of-care diagnostics.
Since the invention of the metal-oxide-semiconductor FET (MOSFET) in 1959, a myriad of
FET-based sensors with different structures, sensing materials, and their target analytes havebeen developed. The ion-sensitive FET (ISFET) itself was introduced by Bergveld in 1970(15)and derived from the MOSFET structure. In practice, ISFET sensors are often operated in aconstant charge or constant drain-current mode, i.e., by setting the drain current at a fixed value
using a feedback circuit, the resulting sensor output signal from the biochemical reaction is thenproportional to the voltage shift recorded.(16) However, FET sensors have inherent limitations
such as the Debye screening length (counter-ion screening effect), which occurs during
measurements in an aqueous medium, and that the sensitive detection of biomolecules with low
or neutral charges is difficult without some form of interaction that can significantly affect thesurface potential of the FET sensor.
Some of the earliest attempts of FET-based CRP sensors were by Sohn et al. in 2007,(17)Sohn and Kim in 2008,(18) and Lyu et al. in 2009.(19) Their techniques mostly used the extendedgate FET configuration where the gate electrode is connected to the sensing layer that is
physically external to the FET itself. CRP detection occurs when negatively charged CRP binds
to the CRP antibodies immobilized on the gate surface, and this induces a change in drain–source current. To improve the sensor’s limit of detection and detection range, especially inphysiological fluids (high ionic strength), sensing methods have been proposed over the years
with varying levels of success. Examples are silicon nanowire FET CRP sensors developed by
Lee,(20) Lee et al.,(21) and Kwon et al.,(22) the nanogap-embedded FET proposed by Ahn etal.(23) and Kim et al.,(24) a hybrid MOSFET with a bipolar junction transistor,(25,26) a carbon
nanotube FET,(27) and an AlGaN/GaN high-electron-mobility transistor (HEMT),(28,29) which
was also combined with a microfluidic platform, an electrolyte-gated organic FET sensor,(30) ora different oxide material for the gate electrode.(31) These sensors were mostly focused on
nanotechnology or gate oxide material to improve their capabilities. This required complicatedfabrication steps or expensive nanomaterials that can hinder the mass production of the sensors.

Sensors and Materials, Vol. 35, No. 10 (2023)

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On the other hand, ISFETs are readily and inexpensively available on the market. ISFETs for pH
sensing have a pH sensitivity of about 56–58 mV/pH, which originated from the protonation anddeprotonation of the −OH groups of their gate material, such as Ta2O5.(32,33) This means that thesensor is easily affected by pH. Therefore, by eliminating the −OH groups on the gate surface ofa commercial ISFET and adding a CRP biorecognition layer to the gate surface, a CRP sensor
can be mass-produced by taking advantage of the commercial ISFET sensor. Furthermore, mostof the FET-based CRP sensors were immunosensors and used anti-CRP antibodies as thebiorecognition layer. Antibodies (about 10 nm) are larger than the Debye length of physiologicalfluids (about 0.7 nm), which requires the sensors to use diluted solutions or nanotechnology as
mentioned previously. Aptamer-based FET sensors for CRP detection only emerged in2015.(34–36) Aptamers are synthetic oligonucleotides that are smaller than antibodies; therefore,
their interactions with the target analyte are more likely to be within the Debye screening length.
Aptamers’ backbones are made up of negatively charged phosphate groups; therefore,biomolecules binding to the aptamers will enable signal transduction and amplification even forbiomolecules with low or neutral charges.
In this work, we propose a simple method of chemically modifying the gate surface of acommercial ISFET with silane coupling agents to eliminate the −OH groups on the gate surface
to reduce pH sensitivity, which is indirectly the background signals. CRP-binding aptamers areused as the biorecognition layer. To the best of our knowledge, this work is the first to propose
the method of decreasing the pH sensitivity of the gate material to increase the sensitivity of the
FET CRP sensor.

2.

Data, Materials, and Methods

2.1

Reagents and solutions

Human CRP full length protein (AB167710) was purchased from Abcam. The CRP-binding
aptamer with amine-terminated end 5’-GGCAGGAAGACAAACACGATGGGGGGGTATGAT
TTGATGTGGTTGTTGCATGATCGTGGTCTGTGGTGCTGT-3’(37,38) with a binding affinityo f K d = 3. 51 n M w a s m a n u f a c t u r e d b y E u r o f i n G e n o m i c s .
(3-Glycidyloxypropyl)trimethoxysilane (GPTMS) was purchased from Sigma-Aldrich and(3-aminopropyl)trimethoxysilane (APTMS) from Tokyo Chemical Industry (TCI). Phosphatebuffered saline [PBS(−)] was purchased from Nacalai Tesque. Phosphate pH standard equimolal
solution, 25% glutaraldehyde, ethanolamine, acetone, isopropanol, and ethanol were purchasedfrom Fujifilm Wako Pure Chemical. All aqueous solutions in this study were prepared using
Milli-Q water (Millipore Corp.).
2.2

Device fabrication and biofunctionalization

The ISFET sensors were purchased from ISFETCOM Co., Ltd., Japan (Fig. 1). The ISFET
sensor was first ultrasonicated up to 10 s in pure water, acetone, and isopropanol, in thissequence. It then undergoes plasma ashing at 70 W for 60 s. The cleaned ISFET sensor was4800

Sensors and Materials, Vol. 35, No. 10 (2023)

(a)

(i)
PolyimidebodyConnector

(ii)

ISFET

Aptamer-modified
gate (G)

Source (S)
n+

Ta2O5

Insulator
ChannelDrain (D)
n+

Si

(b)

Fig. 1. (Color online) (a, i) ISFET aptasensor with a closeup of the ISFET sensing chip. (a, ii) Simplified schematicof the ISFET aptasensor and detection mechanism. (b) Biofunctionalization steps of the ISFET sensor.

silanized by gas-phase silanization with either (i) 400 μL of GPTMS for 2 h at 80 ℃ in vacuumor (ii) 400 μL of APTMS overnight at room temperature in vacuum. A physically adsorbed FET
aptasensor was fabricated by (iii) incubating the cleaned ISFET with the CRP-binding aptamer(2 μM) overnight at room temperature. Silanized ISFET sensors were rinsed with anhydrous
ethanol to remove unreacted silane from the ISFET surface before the CRP-binding aptamer (2μM) was then reacted with the silanized surface overnight at room temperature. Afterwards,
PBS was used to rinse the ISFET before immersing in 50 mM ethanolamine as blocking buffer
overnight at room temperature. Further rinse using PBS was performed before the ISFET wasused for CRP detection.
2.3

ISFET pH sensitivity evaluation and CRP quantitative measurement

The ISFET is connected to the customized measuring apparatus and a commercial Ag/AgCl(sat. KCl) reference electrode (BAS Inc., USA). It is first calibrated to 0.0 V in pH 6.86 phosphate
pH standard equimolal solution at room temperature before adding 0.1 M NaOH or 0.1 M HCl to
the pH 6.86 standard buffer solution to change the pH, and thus the voltage, accordingly. The pHof the solution is recorded using a pH meter (Mettler Toledo, Belgium). Aptamer-modified
ISFETs are calibrated to 0.0 V using PBS solution to prevent harming the aptamer layer. 0.1 M
NaOH or 0.1 M HCl is then added to the PBS solution to change its pH.
CRP detection is carried out almost similarly. First, the ISFETs are calibrated to 0 V in PBS
solution (pH 7.4), i.e., 0 μg/mL solution, before being incubated for 30 min in 0.002, 0.02, 0.2, 2,
5, 10, and 20 μg/mL CRP solutions (diluted in PBS), in that order. The ISFETs are rinsed
thoroughly in PBS in between measurements. These steps are then repeated using 0.1 × PBS and0.01 × PBS solutions to investigate the Debye screening effect.

Sensors and Materials, Vol. 35, No. 10 (2023)

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3.

Results and Discussion

3.1

Silanization of ISFET surface of Ta2O5 and immobilization of aptamer

We opted to use silanization, a method that will simultaneously lessen the density of −OH
groups and enable aptamer immobilization. The conventional silane often used for the FET
surface is aminosilane (APTMS or APTES), which has a positively charged amine group at one
end to immobilize proteins using a crosslinking agent. This is detrimental to our strategy of
reducing the pH sensitivity of the FET sensor; hence, we chose to use epoxysilane (GPTMS) that
can also facilitate aptamer immobilization without the need for a crosslinking reagent. We also
conducted silanization using APTMS as a control. AFM and XPS measurements were used toevaluate untreated and silane-treated surfaces, and the results are shown in Fig. 2. Since the FET
sensors were very small for the measurements, we used samples fabricated using Ta2O5 sputteredon Si wafers instead. The peak changes in XPS data (for Si 2p and 2s) between untreated Ta2O5
and GPTMS-treated Ta2O5 clearly indicate the formation of a silane layer on the Ta2O5 surface.
Further aptamer immobilization shows the appearance of a N peak, which can be attributed tothe amine group of the aptamer (Table 1).
A more detailed discussion can be made using the composition ratio summarized from theoverall XPS data (note that the composition ratio is only for rough comparisons and not accurate(a) Ta2O5 onlySi 2p(b) GPTMS-treated

Si 2p(c) APTMS-treated

Si 2p

Si 2s

Si 2s

Si 2s

Fig. 2. (Color online) Surface modification evaluation by XPS, Si2p, and 2s scans. (a) Untreated Ta2O5 surface, (b)GPTMS-treated Ta2O5 surface, and (c) APTMS-treated Ta2O5 surface.

4802

Table 1
Composition ratio summary from XPS data.
Ta (4f)Si (2p, 2s)
Untreated
0.27
0.01
APTMS
0.01
0.16
GPTMS
0.13
0.03
Apt
0.14
0.01
APTMS-GA-Apt
0.00
0.08
GPTMS-Apt
0.14
0.03

Sensors and Materials, Vol. 35, No. 10 (2023)

O (1s)
0.64
0.22
0.39
0.39
0.19
0.38

C (1s)
0.08
0.44
0.10
0.04
0.60
0.07

P (2p)
—
0.00
0.00
0.02
0.01
0.00

N (1s)
—
0.10
0.33
0.36
0.08
0.37

descriptions of the surfaces measured). Silanization using either APTMS or GPTMS results in adecrease in Ta composition rate and an increase in Si, O, or C composition ratio, which indicatethat Ta-OH is now replaced by Ta-O-Si-CH3 instead. Hence, it can be considered that bothsilanes can reduce the density of −OH groups on the Ta2O5 surface. However, APTMS showed amore marked Ta decrease and Si increase in composition ratio compared with GPTMS. We
consider that this is because the reaction time for APTMS is much longer than that for GPTMS(overnight compared to 2), which led to an increase in the vertical polymerization of APTMS
leading to the formation of 3D “islands” compared with mostly horizontal polymerization for
GPTMS.(39,40) APTMS also has a short carbon chain, which meant that it is more prone to 3Dpolymerization as reaction time is increased Therefore, GPTMS is more successful in forming athin silane layer than APTMS, which is also effective in lowering pH sensitivity as explained in
the next section. P and N data for both APTMS and GTPMS are also shown. Since they are notmodified with aptamers, P is undetectable for both APTMS- and GPTMS-treated surfaces. A
slight increase in N composition ratio is seen in the APTMS-treated surface, which originatedfrom the amine groups of APTMS itself. However, the unexpected increase in N for the GPTMStreated surface can only be attributed to possible cross-contamination from the vacuum
container due to prior aminosilane treatment, since GPTMS nor the chemicals used during theepoxysilane treatment have any N group. (This vacuum container was not used for the next
GPTMS-treated and aptamer-immobilized Ta2O5 surface; hence, the results can be properly
attributed to the modification treatment.)
Aptamer immobilization is then confirmed by scanning for P and N, which are the phosphategroups in the aptamer’s backbone and the amine group at one end of the aptamer (Table 1). For
the APTMS-treated and aptamer-immobilized Ta2O5 surface, P and N were both confirmed;
however, interestingly, the Ta composition ratio was effectively zero. We attributed this to theformation of the said “islands” and the aptamer immobilization (physical adsorption of aptameron the untreated Ta2O5 surface also reduced the Ta and O composition ratios) having physicallycovered the Ta2O5 surface such that Ta was undetectable (for that specific point of measurement).
Note that the N composition ratio was lower than that of the physically adsorbed aptamer on theuntreated Ta2O5 surface. Considering that the APTMS itself also has amine groups, weconcluded that the formation of 3D islands reduced the number of viable sites for aptamer
immobilization; hence, a lower number of aptamers were immobilized compared with physicallyadsorbed and GPTMS-treated aptamer modifications.

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In contrast, while N was detected (at about the same level as for physically adsorbed aptamers
on the untreated Ta2O5 surface) for the GPTMS-treated and aptamer-immobilized Ta2O5 surface,
P was effectively undetectable (too low to be detected but not exactly zero). However, P can bedifficult to detect since the aptamer backbone where the phosphate groups are located can be ina rolled-up or loose configuration, as indicated by the low ratios for both physically adsorbed
and APTMS-treated aptamer modifications. Hence, the GPTMS-treated aptamer modification
can be considered to have a structure different from those of the other two methods.
3.2

Reducing pH sensitivity for lower background signals

We also investigated how the silane surface treatment affected the pH sensitivity bymeasuring the pH dependence on the surface potential of the ISFET, i.e., its output voltage. As
shown in Fig. 3, GPTMS reduced the pH sensitivity better than APTMS. From the untreated
FET sensor with a sensitivity of about 56.2 mV/pH, it dropped to about 54.4 mV/pH after the
GPTMS treatment. This can be attributed to the −OH groups becoming reduced by thesilanization reaction, combined with the slightly negative charge of the epoxy group that did notparticularly affect the pH sensitivity of the FET sensor. The subsequent treatment with the NH2terminated CRP-binding aptamer and blocking buffer ethanolamine caused further increase anddecrease in pH sensitivity, which can be considered as the effects of the following:i. the amine-terminated aptamer bonding with the epoxy group to create neutral charges andii. the ethanolamine that has its amine side bonded with unreacted epoxy groups but has a free
−OH side itself.
Overall, the GPTMS treatment helped reduce the pH sensitivity of the FET sensor.
Conversely, even with −OH groups reduced by silanization, the APTMS treatment had slightly
increased the pH sensitivity instead, which is due to the positive charges of the amine group.
Further functionalization steps did not significantly change the pH sensitivity of the pH sensor
either.
(a)(b)

Fig. 3. (Color online) Gate surface potential shift of the FET by pH change from pH 7.0 for pH sensitivity
evaluation of the (a) APTMS-treated and (b) GPTMS-treated FETs.

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3.3

Sensors and Materials, Vol. 35, No. 10 (2023)

CRP detection in 1 × PBS

Modified and nonmodified FET sensors or aptasensors (aptamer-modified sensors) were thenused to detect 0, 0.002, 0.02, 0.2, 2, 5, 10, and 20 μg/mL CRP, and the results are shown in Fig. 4.
A physiological ionic strength solution was used to provide practical results in terms offeasibility as a point of care (POC) biosensor. The epoxysilane-treated FET aptasensor (GPTMSapt) successfully detected down to 0.002 μg/mL CRP with a more distinguishable and higher
signal strength between CRP concentrations than the other FET sensors. The aminosilanetreated FET aptasensor (APTMS-apt) was only able to clearly detect CRP from 2 μg/mL
onwards. The nonsilanized physically adsorbed aptasensor (bFET-apt) and untreated blank FET
sensor (blank-FET) were only able to start detecting CRP from 5 μg/mL onwards. Since the
CRP level between 1 and 3 μg/mL is considered as medium risk and more than 3 μg/mL isalready high risk, APTMS-apt, bFET-apt, and blank-FET were not practical for POC biosensor
use. Instead, GPTMS-apt that was able to detect CRP levels ranging from 0.002 to 20 μg/mL in
physiological ionic strength solution showed its feasibility as a POC CRP biosensor. To reiterate,
this GPTMS-treated FET aptasensor did not need expensive nanomaterial modifications or acomplicated nanostructure fabrication to achieve its current sensitivity and measurement range,
compared with other FET sensors for CRP detection (Table 2). Furthermore, the GPTMS-treated
FET aptasensor was capable of measuring the current CRP concentration range in 1 × PBS, ahigh-ionic-strength solution, compared with other FET biosensors for CRP, which resorted tomeasuring under dry conditions(24,27) to circumvent the Debye screening effect.
3.4

CRP detection in solutions with different Debye screening lengths

The performance of the GPTMS-treated FET aptasensor was evaluated in PBS with different
ionic strengths and the results are shown in Fig. 5. The FET aptasensor shows increased signal
strength in solutions with low ionic concentrations. This is because when the solution has fewer
buffer ions, the counter-ion screening effect becomes weaker and the effective measuring

Fig. 4. (Color online) Gate voltage change during CRP detection after calibration to 0 V in 0 µg/mL solution forcomparison of CRP detection sensitivity between ISFETs in 1 × PBS. Error bars represent the standard deviations(n = 3).

Sensors and Materials, Vol. 35, No. 10 (2023)

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Table 2This work in comparison with other FET and non-FET biosensors for CRP.
Experimental
Limit of
Sensor type
LigandDetection time*1 Detection range
conditions
detection*2
This work
aptamer
1 × PBS35 min0.002–20 μg/mL 0.002 μg/mL
antibody
unidentified—
3–10 μg/mL—
Extended gate FET(17)
Na nogap convent ional SiO2
antibody
dry
20 min0–100 ng/mL
0.1 ng/mL
FET(24)
AlGaN/GaN HEMT with nullantibody
1 × PBS3h
10–1000 ng/mL
10 ng/mLbalancing circuit(28)C y s t e i n e - t a g ge d p r o t e i n G
antibody
unidentified—
3–20 µg/mL—
modified Au/NiCr gated FET(18)(31)
SiO2-CeO2 ISFET(biotin /antibody
1 × PBS30 min0.1–2.5 μg/mL
0.1 μg/mL
streptavidin/Ab)
Carbon nanofiber immunoassay
antibody
1 × PBS80 min0.05–5 μg/mL 0.011 μg/mL+ EIS(41)Magnetic NP + micro fluxgate
antibody
1 × PBS30 min0.002–10 μg/mL 0.002 μg/mL
sensor(42) (magnetic field)(27)
antibody
dry
20 min0.001–100 μg/mL 0.001 μg/mL
SWCNT SiO2 Ti/Au FET
Pol y m e r- m o d i f i e d o r g a n i c
0.4 ng/mL–
antibody
1 × PBS15 min0.22 ng/mL
FET(30) (physical adsorption)
0.2 mg/mL*1: Estimated as a total reaction time including incubation time for CRP and ligand binding, washing time, andmeasurement time from the experimental procedure in each paper.
*2: Value as written in each reference was adopted. Generally, the limit of detection is the lowest analyte concentration
where the signal exceeds "BG + 3σ", derived from background (BG: an averaged signal at zero analyte) and standard
deviation at zero analyte (σ). In this work, the limit of detection is defined as the lowest concentration where the entire
error bar is within the positive region.

Approximate composition of 1x PBS buffer: 138 mM
NaCl, 2.7 mM KCl, 1.9 mM NaH2PO4, 8.1 mM Na2HPO4

Fig. 5. (Color online) Gate voltage change during CRP detection in PBS with different ionic strengths (1×, 0.1×,
0.01×) using GPTMS-treated FET aptasensors. Error bars represent the standard deviations (n = 3). The FET
aptasensors were first calibrated to 0 V in 0 μg/mL solution.

distance becomes shorter. Interestingly, although the FET aptasensors were calibrated to 0 V in
0 μg/mL CRP (either 1×, 0.1×, or 0.01 × PBS), low concentrations of CRP in 0.1× and 0.01 × PBS
gave a positive potential output that grows increasingly negative with increasing CRP
concentration, instead of starting with a negative potential output. This phenomenon can be
attributed to the fact that the same concentration of CRP is present with increasingly fewer

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buffer ions in each ionic strength solution, which meant increased effective measuring distance.
Hence, the slightly negative charges of the CRP itself also affected the gate surface potential.
This and combined with the CRP concentration being too low to cause significant conformationalchanges in the aptamer thus gave the positive potential output for low CRP concentrations.
Therefore, as the CRP concentration increased and the aptamer conformational changes
significantly contributed to gate surface potential changes, the potential output becomesincreasingly negative. Furthermore, the output signals for 0.1 × PBS did not significantly
increase compared with the output signals for 1 × PBS. This may have been due to fewer bufferions affecting the binding of CRP to the aptamer. Since CRP is slightly negatively charged,
without some shielding effect from the buffer ions, electrostatic repulsion may have occurred
between the negatively charged phosphate backbone of the aptamer and the CRP itself, affecting
the FET aptasensor’s detection sensitivity. In other words, aptamer binding is largely mediatedby electrostatic forces.(42–44) Hence, the limits of detection for 0.1 × PBS and 0.01 × PBS differ
from that of 1 × PBS (5 µg/mL for 0.1 × PBS and 2 µg/mL for 0.01 × PBS).

4.

Conclusions

In this work, a novel concept for increasing the signal sensitivity of a FET aptasensor using
easy and low-cost modification methods is proposed. The FET aptasensor treated with GPTMS
by gas-phase silanization and then aptamer-functionalized was able to measure 0–20 µg/mL
CRP in 1 × PBS solution (physiological ionic strength) with a higher signal sensitivity than thoseof other (modified or nonmodified) FET sensors. Furthermore, the FET aptasensor did not
require expensive nanomaterial modifications or a complicated nanostructure fabrication toachieve its current sensitivity and measurement range, compared with other FET sensors for
CRP detection. However, the confirmation of specificity and investigation using clinical samples
such as spike and recovery tests using blood and serum remain. We will proceed with further
studies including these evaluations to realize a low-cost and practical CRP sensor for accuratedisease monitoring.

Acknowledgments
This research was supported by JST COI (no. JPMJCE1303) from the Japan Science andTechnology Agency (JST). This research was also supported partly by AMED Grant Number
JP21zf0127001 from the Japan Agency for Medical Research and Development (AMED). It wasalso partially supported by a Grant-in-Aid for Scientific Research (B) (no. 22H02105) and a
Grant-in-Aid for Challenging Research (Exploratory) (no. 22K18941) from the Japan Society for
the Promotion of Science (JSPS) and JST FOREST Program (no. JPMJFR213C) from JST. We
would like to express our thanks to Professor Nagatomi Ryoichi and Professor Yoshida Shinya
for their input on the need for CRP monitoring and on using FET biosensing. We are also
grateful to the Nishizawa Memorial Centre for allowing the use of their sputtering machineryand to the Instrumental Analysis Group of Tohoku University for the XPS and AFM
measurements and their invaluable insight.

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