Serotonin is one of the monoamine neurotransmitters in the central nervous system, and its concentration is important in the body owing to its responses to physiological and psychological functions, such as sleep, Alzheimer’s disease, appetite, pain, sexuality, desire, and mental state [1,2]. Generally, the concentration of serotonin in the human body is approximately 295 to 687 ng/mL in human urine, and irregular concentrations can occur with various disorders [3,4]. In particular, an increasing number of people have suffered depression owing to numerous stresses from their personal lives, society, and the COVID-19 pandemic. Therefore, rapid detection methodologies are required for monitoring changes in serotonin concentration.

Although many technologies, including cyclic voltammetry, radioimmunoassays, chemiluminescence, enzyme-linked immunosorbent assay, field-effect transistors, fluorescence, and the polymerase chain reaction, have been developed for serotonin analyses, these methods have limitations; they can be time-consuming or exhibit a low detection range, low sensitivity, and low selectivity [1,5–10]. Although various electronic sensors were recently developed for serotonin detection, these sensor platforms show deficiencies related to rapid detection, real-time monitoring, and on-site application systems. To overcome these limitations, technology is required to provide high performance, quick testing, a broad detection range, high-sensitivity and selective detection, and this remains a challenge.

In this study, we first demonstrate a real-time monitoring system using a portable serotonin antibody (SAb)-graphene micropatterned field-effect transistor (GMFET) device for on-site monitoring and serotonin detection. Graphene micropatterned (GM) was used with SAb as a bioprobe for the channel of the FET, and the selective serotonin detector bis(2-aminoethylene)perylene-3,4,9,10-tetracarboxyldiimide (PDA) was utilized to chemically immobilize the SAb onto the graphene. In particular, PDA was synthesized to include two amine groups at the terminus for increased conjugation with SAb [11]. The SAb-GMFET showed high sensitivity and selectivity, a limit of detection (LOD) of 10 pM and a wide range for 10 pM to 100 nM with a 10 s response time. Moreover, the K value was calculated by the Langmuir adsorption equation and was 1.63 × 10⁻¹⁰ M⁻¹, which indicated the equilibrium constant for the SAb-GMFET. To prove selective detection by the SAb-GMFET, the interfering materials cortisol, epinephrine, oxytocin, dopamine, and tryptamine were injected to compare them with the serotonin present at a 100-fold lower concentration, and selective interaction with serotonin resulted. To apply the device on-site, a portable SAb-GMFET was developed by using wire bonding and a printed circuit board (PCB) sim chip. This portable platform showed high sensitivity and a LOD similar to that of the SAb-GMFET. Based on these results, this portable SAb-GMFET platform can be utilized for rapid on-site detection and perhaps for self-diagnosis with a point-ofcare test (PoCT).

2.1. Materials and instruments

Copper foil and glutaraldehyde (GA) were purchased from Thermo Fisher and Sigma Aldrich Co., Ltd. Samples of 950 PMMA A4 4 % and serotonin antibody (ab66047) were purchased from Micro Chem Co., Ltd. and Abcam. Selected area electron diffraction (SAED) images were obtained using high-resolution transmission electron microscopy (Jeol, JEM-ARM200F), and Raman spectra were obtained with a LabRAM HR Evolution Visible_NIR system from HORIBA using a 633 nm laser.

2.2. Fabrication of monolayer graphene

Monolayer graphene (MLG) was synthesized using a chemical vapor deposition (CVD) method based on our previous reports [12–15]. Graphene film was formed on the copper foil substrate and was synthesized from the gas sources CH₄, H₂, and Ar at 1,000 °C. Then, the MLG surface was coated with PMMA for protection. The copper film was removed using copper etchant, and the MLG was transferred onto a silicon dioxide wafer using the wet-transfer method. Finally, the PMMA was removed by acetone vapor and washed with distilled water.

2.3. Fabrication of SAb-GMFET

The electrodes for the SAb-GMFET were fabricated by a microelectromechanical system (MEMS) process [16,17]. AZ GNX-601 positive photoresist was utilized with GM in the photolithography process, and a reactive-ion etching system was used for graphene etching. A DNR L300-40 negative photoresist was used to form the electrode (Cr/Au) and passivation layer (SiO₂). Ten micromolar PDA linker compound was used for conjugation on the GM channel for 30 min, and 2 % GA coupling agent was dripped onto the PDA/GM for 4 h. SAb with a concentration of 2 µg/mL was reacted with the modified GM channel for 4 h at 4 °C.

2.4. Electrical measurements

All electrical measurements were performed using a probe station (MS-tech, Model 4000) and a Keithley 2612 A source meter. The normalized sensitivity was calculated with the following equation:

where I indicates the immediately measured current level and I₀ is the initial current level.

3.1. Characterization of SAb-conjugated MLG

MLG has various properties, such as a zero bandgap, high carrier mobility, high conductivity, ambipolar properties, and rapid electron transport [14,18–20]. MLG has been utilized in various applications and has been applied in preparing bio/chemosensors owing to its excellent electrical properties [21,22]. To confirm the preparation of MLG, the graphene sample was analyzed using high-resolution transmission electron microscopy (HR-TEM), and the hexagonal lattice pattern was observed in a SAED image [Fig. 1(a)]. The MLG surface was modified by PDA for immobilization of the SAb, and the PDA interfacial chemical was stacked on graphene via π interactions between the sp² hybridized graphene and the perylene backbone of the PDA. The PDA-doped graphene was analyzed with Raman spectroscopy, X-ray photoelectron spectroscopy (XPS), and Fourier transform infrared (FT-IR) spectroscopy. The Raman spectrum showed ideal monolayer graphene with an approximate I_2D/I_G = 1.989 (black line), and the red line spectrum exhibited a D peak due to PDA introduction [Fig. 1(b)]. The intensities of the C 1s and O 1s XPS peaks increased after PDA conjugation on the MLG, and the N 1s core level appeared at 398 eV because of the amine group of the PDA terminus [Fig. 1(c)]. Therefore, introduction of the PDA onto the graphene was confirmed with these results. The FT-IR spectrum was obtained to confirm immobilization of the SAb on graphene, and the amide I and II bands were observed at 1,700 to 1,600 cm⁻¹ and 1,580 to 1,408 cm⁻¹, respectively [Fig. 1(d)]. The broad absorption peaks for -OH and - NH moieties were located between 3,300 and 3,200 cm⁻¹, and broad peaks ranging from 1,700 to 1,200 cm⁻¹ were attributed to bending vibrations of the amine, carbonyl, and alkyl groups. Perfect conjugation of the SAb on the MLG via the PDA linker was obviously confirmed by the increased intensities observed for amide I and II peaks.

Figure 1. Characterization of the graphene surfaces subjected to surface modification. (a) SAED image and (b) Raman spectrum of MLG (I_G/I_2D = 1.989). (c) XPS spectra before (black) and after (red) PDA conjugation. (d) FT-IR spectrum of pristine graphene (black), PDA-conjugated graphene (red), and SAb-immobilized graphene.

3.2. Electrical properties of SAb-GMFET

Current-voltage (I-V) curves were measured over the range −2 V to 2 V, and the graphs showed gradually decreasing slopes that depended on the surface modifications of the MLG [Fig. 2(a)]. However, the current was slightly decreased by π-π stacking due to PDA introduction (physical attachment) and conjugation of the SAb (chemical attachment), which resulted in obvious ohmic properties and followed Ohm’s law. An illustrated scheme is presented for the liquid-ion gated FET system with the SAb conjugated-GMFET, which consisted of source (S), drain (D), and gate (G) electrodes [Fig. 2(b)]. In addition, PBS solution was added into the chamber as a dielectric material, which was formed by the field effect of the voltage supplied from the gate electrode. Following our previous studies, the conductance of the GMFET showed ambipolar properties. The liquid-ion gated FET system was suitably verified of the operation in the p-type region because electron transfer was induced by oxygen, and the biomolecules were damaged by the radicals generated in the n-type region in the liquid phase. [23]. The transfer curves showed the shifted Dirac point as a minimum gating effect depending on the surface modifications of the graphene surface, which indicated a change in the charge carrier density caused by PDA and SAb [Fig. 2(c)]. The Dirac point was negatively shifted from 0.9 V for GM to 0.85 V for PDA-conjugated GM because of the amine group of the PDA terminus (red line), and the SAb-GMFET exhibited a negative shift of approximately −0.1 V due to the negative charge at pH 7.4 (blue line) [24]. The output curves of the SAb-GMFET were measured over the V_ds range 0 V to −1 V (sweep of V_g = −0.1 V) to investigate the dependence of p-type properties on the gating effect [Fig. 2(d)]. Changes in the I_ds were observed to increase the current depending on the negative increase in V_g, which indicated a typical p-type transistor based on hole-transfer behavior [25–30].

Figure 2. Electrical properties of the SAb-GMFET. (a) I-V curves depending on the SAb conjugated on the GM. (b) Schematic diagram for the liquid-ion gated FET system (including source, drain, gate electrodes, serotonin antibody, and dielectric material). (c) Transfer curves for the GMFET and SAb-GMFET. (d) Output curves (I_ds-V_ds) for the SAb-GMFET with P-type properties.

3.3. Real-time monitoring of SAb-GMFET

To confirm the serotonin detection performance of the SAb-GMFET, an electrode was fabricated by using the MEMS process. The liquidion gated SAb-GMFET was tested during exposure in real time, and the sensitivity of the detection performance was calculated with the normalization based following equation:

The real-time responses of the SAb-GMFET to various concentrations of serotonin are displayed in Fig. 3(a). Before monitoring the current change, a buffer solution without serotonin was injected to confirm the buffer effect, and there was no electrical change. On the other hand, changes in the current were immediately observed after exposure to serotonin in the concentration range 10 pM to 100 nM, and the response time was less than 10 s. The LOD of the SAb-GMFET was 10 pM, which was 30 times higher than those seen in other studies using SAb (Table I). The concentration curve was obtained from the normalized sensitivity shown in Fig. 3(a). A linear response was displayed in the concentration range 10 pM to 10 nM, and the current level was saturated at concentrations over 100 nM [Fig. 3(b)]. To investigate the binding affinity, the K value was calculated with Langmuir’s adsorption isotherm equation by using the calibration curve:

Table 1 . Performance comparison of various biosensors used for serotonin detection..

Categorization	Materials	LOD	Ref.
Cyclic voltammetry	GI-WO3 NP	1.42 nM	[32]
Cyclic voltammetry	Au NPs/rGO	387 nM	[33]
Cyclic voltammetry	Co3O4/rGO	1.128 µM	[34]
Cyclic voltammetry	CNFs	250 nM	[35]
Field-effect transistor	β-Bi2O3 nanofiber	290 pM	[1]
Field-effect transistor	CVD-graphene	10 pM	In this study

Figure 3. erformance measurements with the SAb-GMFET. (a) Real-time monitoring toward various concentrations of serotonin (1 pM to 10 µM). (b) Concentration curve for the SAb-GMFET based on real-time monitoring results. (c) Selectivity test of the SAb-GMFET with serotonin. (d) Long-term stability of the SAb-GMFET platform on 10 pM serotonin (n = 5).

where N is the normalized sensitivity, C is the concentration of serotonin, and the obtained K value for serotonin was 1.63 × 10⁻¹⁰ M⁻¹ [31]. To demonstrate selective detection, real-time monitoring was carried out by using the interfering materials cortisol, epinephrine, oxytocin, and dopamine as neurotransmitters, and tryptamine as a precursor [Fig. 3(c)]. Although no significant response was observed for the buffer and with 100-fold higher concentrations of the interfering materials relative to that of serotonin, the current level was obviously changed by serotonin with a 10 pM concentration. Therefore, this SAb-GMFET platform clearly detected serotonin with no effect from potential interfering substances. The reproducibility and repeatability were evaluated over 5 replicates, and the sensitivity consistently exhibited an error range of 2.5 % [Fig. 3(d)]. Moreover, the long-term stability of the SAb-GMFET was assessed to look for variations in sensitivity over 7 days of exposure to 10 pM serotonin. Although the sensitivity was maintained for 4 days, the performance was slightly decreased over 5 days. However, the performance level of the SAb exceeded 50 % of the initial performance.

3.4. Portable SAb-GMFET device

Although many serotonin detection technologies have been developed, these analysis methodologies exhibit limitations such as low sensitivity, low selectivity, and long measurement times. The SAb-GMFET platform was utilized in a portable device for PoCT health care applications [Fig. 4(a)]. The portable platform consisted of a hardware platform based on a PCB, such as a microcontroller unit (MCU), power supply, communication module for outgoing data, and chip socket for docking of the chip integrated with the SAb-GMFET electrode. In addition, the sim chip with the integrated SAb-GMFET was connected by wire bonding between the PCB and electrode to supply the voltage. The performance of the portable SAb-GMFET was evaluated by injecting the LOD concentration of serotonin [Fig. 4(b)]. The observed sensitivity was similar to that seen for the SAb-GMFET, as shown in Fig. 3(a), which indicates the possibility of on-site utilization. In addition, a portable SAb-GMFET was used to carry out real-time monitoring of the diluted serotonin in artificial urine for possible selfdiagnosis [Fig. 4(c)]. The results were normalized and compared with those in Fig. 4(b), and they showed similar results for the same concentrations.

Figure 4. (a) Optical image of a portable SAb-GMFET device with the inserted sim chip. (b) Real-time monitoring of various concentrations using the portable device. (c) Real-time response toward serotonin in artificial urine.

In summary, we have demonstrated a sensitive and selective SAb-GMFET platform and developed a portable device. The SAb-GMFET achieved rapid detection and real-time monitoring of serotonin and comprised SAb for selective detection and a liquid-ion gated FET as a highly sensitive monitoring system. The SAb-GMFET exhibited excellent sensitivity with an LOD of 10 pM. In addition, the device demonstrated selective detection of serotonin at concentrations 100 times lower than those of various interfering materials, such as neurotransmitters and hormones. Moreover, a portable SAb-GMFET device was constructed by integration of the SAb-GMFET with a sim and connection using wire bonding. The portable device showed excellent detection of serotonin in artificial urine. Therefore, this portable SAb- GMFET platform can be utilized for PoCT and self-diagnosis.

This research was supported by Research Program to Solve Urgent Safety Issues of the National Research Foundation of Korea (NRF) funded by the Korean government (Ministry of Science and ICT (MSIT)) (NRF-2020M3E9A1111636); Korea Institute of Planning and Evaluation for Technology in Food, Agriculture and Forestry (IPET) and Korea Smart Farm R&D Foundation (KosFarm) through Smart Farm Innovation Technology Development Program, funded by Ministry of Agriculture, Food and Rural Affairs (MAFRA) and Ministry of Science and ICT (MSIT), Rural Development Administration (RDA) (421020- 03); the National Research Council of Science & Technology (NST) grant by the Korea government (MSIT) (No. CAP22011-000); the National R&D Program of National Research Foundation of Korea (NRF) funded by Ministry of Science and ICT (2021M3H4A4079276); the Korea Research Institute of Bioscience and Biotechnology (KRIBB) Research Initiative Program (1711170578).