Introduction

People have felt that their mind controls the body. Advances in immunology and neuroscience are scientifically elucidating this experience. For example, it has been clarified that the mechanism by which changes in the activity of the central nervous system due to stress regulate the immune response through sympathetic nerves. If these latest medical knowledge and electronic advances can be used to provide a simple monitoring system for stress-related substances, it is hoped that it will help prevent mental and physical diseases. Nevertheless, performing a blood draw to detect the stress hormones cortisol and catecholamine has the challenge that the act of drawing blood itself causes stress. Instead, cortisol concentration measurement from saliva has been developed as a non-invasive detection, but catecholamines cannot be detected from saliva. Salivary cortisol concentration measurements have been developed as an alternative non-invasive detection method, but catecholamines are unable to obtain the necessary information from saliva. Salivary α-amylase and chromogranin A (CgA) have also been studied and used in part as alternatives to catecholamines. Secretary immunoglobulin A (s-IgA) in saliva is also a good stress marker that reflects suppression of the immune system by stress. Simultaneous monitoring of the time-dependence of these stress markers of different origins is expected to help elucidate the complex mental stress mechanisms [1].

Sensor module platform

Accurate and inexpensive biomaterial detectors are required for IoT biosensing systems and monitoring over time has not been realized until now. Field-effect transistor (FET) biosensors are small devices that detect various types of biomarkers at low power consumption without disturbing the system under test [2]. The biggest challenge of FET biosensors when used in electronics systems is instability due to current drift. We have succeeded in developing a method that minimizes drift using only the normal silicon fab process. This manufacturing process does not use tantalum pentoxide or other special materials.

Figure 1 shows a picture of a newly developed four element FET sensor chip with extremely low instability. The electronics part of the developed biosensor module consists of this chip and a Bluetooth Low Energy (BLE)-type communication circuit. We selected four types of aptamers as sensor receptors on the gate insulator on the chip [3]. The aptamers can be stored and used at room temperature for a long period of time. They also have the advantages of being reversible to thermal denaturation and can be produced inexpensively and industrially. Finally, as a technique for producing biosensors with less variation, such as commercial physical sensors, we developed a tool for uniformly immobilizing the receptor monolayer in a narrow range of fixed positions on the chip.

Simultaneous detection of multiple stress markers in saliva

When n types of receptors are immobilized on n FET elements, n-1 types of independent signals can be extracted. The effects of non-specifically adsorbed substances and pH in saliva, temperature fluctuation, optical noise, and crosstalk between elements can be eliminated from the original signals of multiple FET elements. Using this technique, we succeeded in obtaining multiple stress marker concentrations such as cortisol, a-amylase, s-IgA, and CgA [4], at the same time just by dropping saliva on the sensor. The signal time constant is less than 1 minute, which indicates that a continuous monitor is realized substantially.

Operability equivalent to physical sensor

Internet of Things (IoT) systems require many low-cost sensors. FET sensor chips manufactured using only the conventional silicon fab process can achieve a low cost of about $1 per chip. However, even with such cheap biosensors, if they are disposable, the cost burden on the user will increase significantly in the long run. As a result, it becomes difficult to secure good customers as fixed users. In addition, disposable chips are not suitable for continuous monitoring required for medically important data. We are developing biosensors that are as easy to operate as conventional physical sensors by introducing reusable cleaning methods and recycled precision cleaning methods.

References

[1] L. Steinman, Annu. Rev. Immunol., 32, 257-281, (2014).

[2] K. Ohashi, T. Osaka, ECS Transactions, 75, 39, 1-9, (2017).

[3] N Kaneko, H Minagawa, J Akitomi, K Ohashi, S Kuroiwa, S Wustoni, S Hideshima, T Osaka, K Horii, I Waga, The 43rd International Symposium on Nucleic Acids Chemistry, 2P-55, (2016).

[4] S. Kuroiwa, R. Takibuchi, A. Matsuzaka, S. Hideshima, N. Kaneko, H. Minagawa, K. Horii, I. Waga, T. Nakanishi, K. Ohashi, T. Momma, T. Osaka, 232nd ECS Meeting, 2115, (2017).



Figure 1

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The dosage of 5-fluorouracil (5-FU), which is a widely used for cancer medication, is determined based on body surface area, although efficacy largely depends on the liver function of the individual, resulting that only 21% patients are given an optimal dose of 5-FU in these years. The measurement of the 5-FU concentration in the blood enables us to adjust individual dose adjustment. Conventional methods for 5-FU detection such as an enzyme-linked immunosorbent assay (ELISA) and liquid chromatography are not very suitable for clinical applications because they need time-consuming procedure with expensive equipments. Detection of 5-FU using a field effect transistor (FET) biosensor, which enables rapid and simple measurement, is expected to solve such problems. However, FET biosensor, which detects changes in its surface density due to the adsorption of charged molecules, was unable to detect uncharged 5-FU. In this study, a method for the FET biosensing to detect 5-FU exploiting sequential adsorption of 5-FU modified bovine serum albumin (BSA/5-FU) was proposed. By using this method, FET responses caused by the adsorption of negatively charged BSA/5-FU depending on the 5-FU concentration were detected.

The SiO₂ surface of the FET gate insulator was exposed to O₂ plasma to introduce hydroxyl groups. After the exposure, the surface was exposed to 3-aminopropyltriethoixysilane (APTES), followed by the modification of the cross-linker, glutaraldehyde (GA). A single chain variable fragments (ScFv) and antigen binding fragments (Fab) were allowed to react with each activated GA-modified FET. After the immobilization, the residual aldehyde-groups were treated by ethanolamine to suppress the non-specific adsorption. V _g-I _d characteristics were measured before and after dripping of both 5-FU and BSA/5-FU on the ScFv- and Fab-immobilized FET biosensors. Finally, threshold voltage shifts (∆V _g) caused by the adsorption of BSA/5-FU were obtained.

To compare the capture capability of ScFv and Fab, the electrical responses of the FET biosensors functionalized with the two receptors due to the adsorption of BSA/5-FU were measured. The responses of ScFv- and Fab-immobilized FET biosensor caused by dripping of 25 μg/mL BSA/5-FU were +25 mV and +40 mV, respectively. The difference between ΔV _g values for these two FET sensors using ScFv or Fab can be ascribed to the difference of affinity ^[1]. To verify the specificity of Fab-immobilized FET biosensor, ∆V _g was measured when human serum albumin (HSA) was dropped on the FET biosensor, and the ∆V _g was hardly observed. Additionally, the atomic force microscopic (AFM) images on the FET gate surface shows that the size of observed particles matches the size of BSA/5-FU, while the surface morphology and roughness are not significantly changed. These results indicated that Fab-immobilized surface specifically captured BSA/5-FU. To investigate the quantitative detectability of the Fab-immobilized FET biosensor, we measured the FET responses corresponding to the amount of adsorbed BSA/5-FU, which was related with the concentrations of 5-FU. As a result, the magnitude of ∆V _g by dripping of 1000 ng/mL 5-FU and 25 μg/mL BSA/5-FU was reduced to +12 mV compared with the response of 25 µg/mL BSA/5-FU (Figure 1). These results can be attributed to that the adsorbed 5-FU inhibited the adsorption of BSA/5-FU to the Fab-immobilized surface. Therefore, we conclude that the detection of 5-FU using the FET biosensors by applying the charged BSA/5-FU is a promising simple method for monitoring the concentration of 5-FU.

Reference:

[1] Y. Reiter, et al., J. Biol. Chem., 269, 28, 18327-18331 (1994).

Figure 1 V _g-I _d characteristics of Fab-immobilized FET biosensor before and after dripping of (a) 0 ng/mL 5-FU and 25 µg/mL BSA/5-FU or (b) 1000 ng/mL 5-FU and 25 µg/mL BSA/5-FU.



Figure 1

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A field effect transistor (FET) biosensor is a promising device for various applications such as medical diagnosis and environmental monitoring. Because characteristics of FET biosensors are directly influenced by the change of gate-insulator surface potential induced by the adsorption of charged molecules, FET biosensors could provide the rapid and label-free biomolecular detection. Recently, mental stress-related diseases, such as integration disorder syndrome and depression, affect people's health, resulting that simple stress monitoring is expected for early stage detection of the disease. Previously, the relation between concentration of stress markers and mental stress has been reported ^[1], and the monitoring of circadian concentration of the markers is found to be important for prediction of the stress condition. Especially, secretory immunoglobulin A (s-IgA), which is an immunity-related molecule present in the human mucus, is one of the candidates to be monitored as a stress marker. However, conventional methods for measuring concentration of s-IgA are restricted in daily use due to complex protocol, time-consuming and expensive equipment.

Nowadays, we have investigated sensitive detection method for various targets by using the FET biosensor ^[2,3]. To achieve improvement of the sensitivity, small receptors have been applied to increase electrical responses owing to the effective use of a charge-recognition region from FET gate surface, Debye length ^[4,5]. In this study, we selected a small plant lectin, jacalin (66 kDa), which specifically binds glycan of hinge region of IgA, as a receptor. Additionally, jacalin was inexpensive compared with antibody due to the purification from jackfruits seeds. From these points, jacalin-immobilized FET biosensor was worth to be investigated to realize a simple, sensitive and low-cost stress monitoring device for stress marker. Thus, we investigated the usefulness of the jacalin as a FET receptor.

The SiO₂ gate insulator of the FET was exposed to O₂ plasma (200 W for 1 min) for introduction of hydroxyl groups reacting with triethoxysilane groups of self-assembled monolayer (SAM). Then, the FET chip was immersed in toluene solvent including 1%(v/w) 3-aminopropyltriethoxysilane in an argon atmosphere (60ºC for 7 min.). Following by the cross-linking by glutaraldehyde, jacalin was immobilized on FET gate surfaces. Finally, ethanolamine capping was performed to prevent the non-specific adsorption of contaminating molecules in analyzed samples, resulting in the fabrication of the jacalin-immobilized FET biosensor. The FET characteristics were measured by sweeping the gate-voltage (V _g) from -2.0 V to 0 V with 0.1 V drain voltage (V _d) in 0.01 × phosphate buffered saline (pH 7.4). Then, the electrical responses (ΔV _g) were analyzed from the FET characteristics before and after the addition of analyte to gate surface.

To evaluate the specificity of jacalin-immobilized FET biosensor, ΔV _g caused by the addition of s-IgA and human serum albumin (HSA) were measured. The FET charactristics was shifted in a positive direction (+53 mV) due to the adsorption of negative-charged s-IgA (Figure 1a), while the responses related with HSA addition were scarcely observed. Thus, specific capture of the s-IgA molecules by the jacalin-immobilized surface was indicated. Moreover, to evaluate the advantage of jacalin, we compared ΔV _g with FET functionalized by antigen binding fragment (Fab), which was obtained by cleaving the anti-s-IgA antibody. An electrical response of Fab-immobilized FET was +24 mV (Figure 1b). The change in ΔV _g values for these two FET sensors using jacalin or Fab could be discussed as follows. Jacalin could capture more s-IgA molecules within Debye length from the gate surface of FET. In addition, the jacalin-immobilized FET responded linearly to s-IgA in a concentration range from 0.1 μg/mL to 100 μg/mL. Finally, sweat samples collected from healthy persons were examined with the developed jacalin-immobilized FET biosensor, and clear responses were obtained. From these results, jacalin was found to be useful as a receptor for FET biosensors to achieve a sensitive, simple and non-invasive detection of s-IgA.

[1] K. Obayashi, Clin. Chim. Acta, 425, 196-201 (2013).

[2] S. Hideshima, M. Kobayashi, T. Wada, S. Kuroiwa, T. Nakanishi, N. Sawamura, T. Asahi, T. Osaka, Chem. Commun., 50, 3476-3479 (2014).

[3] S. Hideshima, K. Fujita, Y. Harada, M. Tsuna, Y. Seto, S. Sekiguchi, S. Kuroiwa, T. Nakanishi, T. Osaka, Sens. Bio-Sens. Res., 7, 90–94 (2016).

[4] S. Cheng, K. Hotani, S. Hideshima, S. Kuroiwa, T. Nakanishi, M. Hashimoto, Y. Mori, T. Osaka, Materials, 7, (4), 2490-2500 (2014).

[5] S. Hideshima, H. Hayashi, H. Hinou, S. Nambuya, S. Kuroiwa, T. Nakanishi, T. Momma, S.-I. Nishimura, Y. Sakoda, T. Osaka, Sci. Rep., 9, 11616 (2019).

Figure 1 V _g-I _d characteristics of (a) jacalin or (b) Fab-immobilized FET biosensor before and after the addition of 100 μg/mL s-IgA.



Figure 1

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