Letter Cite This: Nano Lett. 2019, 19, 4004−4009
pubs.acs.org/NanoLett
Electrical Biosensing at Physiological Ionic Strength Using Graphene Field-Effect Transistor in Femtoliter Microdroplet Takao Ono,*,† Yasushi Kanai,† Koichi Inoue,† Yohei Watanabe,‡ Shin-ichi Nakakita,§ Toshio Kawahara,∥ Yasuo Suzuki,⊥ and Kazuhiko Matsumoto*,†
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†
Department of Semiconductor Electronics, The Institute of Scientific and Industrial Research, Osaka University, Ibaraki Osaka 567-0047, Japan ‡ Department of Infectious Diseases, Graduate School of Medical Science, Kyoto Prefectural University of Medicine, Kyoto 602-8566, Japan § Department of Functional Glycomics, Life Science Research Center, Kagawa University, Miki-cho, Kagawa 761-0793, Japan ∥ Department of Clinical Engineering, College of Life and Health Sciences and ⊥Research Institute for Life and Health Sciences, College of Life and Health Sciences, Chubu University, Kasugai, Aichi 487-8501, Japan S Supporting Information *
ABSTRACT: Graphene has strong potential for electrical biosensing owing to its two-dimensional nature and high carrier mobility which transduce the direct contact of a detection target with a graphene channel to a large conductivity change in a graphene field-effect transistor (GFET). However, the measurable range from the graphene surface is highly restricted by Debye screening, whose characteristic length is less than 1 nm at physiological ionic strength. Here, we demonstrated electrical biosensing utilizing the enzymatic products of the target. We achieved quantitative measurements of a target based on the site-binding model and real-time measurement of the enzyme kinetics in femtoliter microdroplets. The combination of a G-FET and microfluidics, named a “lab-on-a-graphene-FET”, detected the enzyme urease with high sensitivity in the zeptomole range in 100 mM sodium phosphate buffer. Also, the lab-on-a-graphene-FET detected the gastric cancer pathogen Helicobacter pylori captured at a distance greater than the Debye screening length from the G-FET. KEYWORDS: Graphene, droplet, enzymatic reaction, Debye screening, Helicobacter pylori where ε, R, T, F, and I are the permittivity, gas constant, temperature, Faraday constant and ionic strength, respectively. λD is less than 1 nm under physiological conditions,13 smaller than most receptors (e.g., the size of IgG is approximately 15 nm14). Debye screening is particularly significant in G-FETs, where no insulator exists and zero distance can be achieved between the graphene and the target. λD can be increased by decreasing the electrolyte concentration, for example, it is increased 10-fold by decreasing the ionic strength 100-fold. However, such dilution reduces the buffering capacity and degrades the biochemical interaction properties such as binding affinities. Electrical biosensing under physiological conditions free from the limitation of Debye screening has been eagerly awaited.11,15,16 If the target generates reaction products that freely diffuse and reach the graphene surface, the target is detected independently of Debye screening. To prove this concept, we have focused on the enzyme urease, which generates
T
wo-dimensional semiconductors and semimetals are promising materials for the electrical sensing of charged molecules since the entire material consists of surface and charged molecules directly in contact with the surface.1−5 In particular, graphene has strong potential owing to its high carrier mobility,6,7 because carrier modulation by target molecules leads to a large change in conductivity. Also, graphene is chemically stable in aqueous solution owing to its wide potential window.8 These features make graphene fieldeffect transistors (G-FETs) attractive for use as electrical biosensors.9−12 However, there is a significant limitation in electrical biosensing, that is, Debye screening. A sensor surface is generally coated with target-specific receptor molecules, such as immunoglobulin G (IgG); however, they separate the target from the sensor because of their size. In aqueous solution, the sensor and target surface are electrically shielded by an electrical double layer with the characteristic length, the Debye screening length λD, given by λD =
εRT 2F 2I
Received: April 1, 2019 Revised: May 16, 2019 Published: May 29, 2019
(1) © 2019 American Chemical Society
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Figure 1. Lab-on-a-graphene-FET device. (a) Schematic image of the device for enzymatic biosensing using G-FET with a microdroplet. The detection target (H. pylori) was captured on graphene via IgG and it generated enzymatic product in a femtoliter microdroplet. The enzymatic product was accumulated in the microdroplet and detected by the G-FET. (b,c) Bright-field images of the device with 10 μm diameter microwells. (d) Fluorescence image of microdroplets on G-FET, obtained at the same position and magnitude as those in (c). Rhodamine 6G (100 μM) was encapsulated and observed at excitation and emission wavelengths of 540 ± 10 nm and >575 nm, respectively. The inset shows a single droplet (scale bar = 50 μm).
ammonia, since ammonia gas has been detected using GFETs.17−20 Furthermore, another concept has been introduced for enzymatic assay using G-FETs. Although the electrical detection of enzyme products using pH sensor and other devices has been reported,21−24 bulk solution of the products was used and the detection did not have significant superiority to optical biosensing using a spectrometer. In contrast to optical biosensing, electrical biosensing only requires enzymatic products near the sensor surface. Therefore, the target was encapsulated in a femtoliter microdroplet to accumulate its reaction products around a G-FET. The encapsulation of enzymes in microdroplets has been utilized for optical singlemolecule detection.25−28 On the basis of these two concepts, in this study we have overcome the limitation of Debye screening by developing a biosensing platform consisting of a G-FET and a microfluidic structure. We named it “lab-on-a-graphene-FET” after the term “lab-on-a-chip”. As a demonstration of the labon-a-graphene-FET, we focused on the diagnosis of Helicobacter pylori infection (Figure 1). H. pylori is a wellknown and main causal factor for gastric cancer.29−33 H. pylori lives in the acidic environment of the stomach, neutralizing the acid using its urease. The urease reaction is a benchmark for H. pylori infection.34,35 First, the G-FET response to ammonia in aqueous solution was measured. Ammonia increases the pH of aqueous solution, and the increase in pH has been measured as the increase in the hole carrier density of graphene and the hole current of a G-FET.1 However, the hole current of the G-FET decreased with increasing ammonia concentration from 100 nM to 100 mM (Figure 2). This result can be explained by considering the dominant species in the solution. The pH response of a GFET has been explained by the site-binding model, where protons bind to graphene defects, surface hydroxyl groups, and the increase electron carrier density of the graphene.36−38 According to the Henderson−Hasselbalch equation, in 100 mM ammonia solution the proton concentration is in the pM range, whereas the concentration of ammonia and its ion are in the mM range. Therefore, in ammonia solution, ammonia and
Figure 2. G-FET response to ammonia in aqueous solution. The blue line represents the fitting curve given by eq 2. Error bars show three standard deviation of the drain current fluctuation in the 300 s measurement for each data point.
its ion are considered to be the dominant species increasing electron carrier density of the graphene. This interpretation is comparable to that of ammonia gas sensing reported elsewhere17−20 where the ammonia molecules adsorbed on graphene decreased the hole current. The interpretation was also supported by the experiment where the graphene/metal contact was fully covered by the perfluoropolymer and only the graphene channel was exposed to the solution (Figure S1). The decrease in the drain current was well fitted by the Langmuir adsorption isotherm given by ΔI =
ΔImaxC NH3 (KD + C NH3)
(2)
where ΔI, ΔImax, CNH3, and KD are the change in the drain current of the G-FET, the maximum change in the drain current, the ammonia concentration, and the dissociation constant of ammonia from the graphene surface, respectively.2 From the fitting, KD was estimated to be 1.30 mM. Currently, graphene formed by exfoliation and other methods has a wide range of qualities in terms of, for example, carrier mobility and defect density. Also, the size of exfoliated graphene varies. 4005
DOI: 10.1021/acs.nanolett.9b01335 Nano Lett. 2019, 19, 4004−4009
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Nano Letters These variations result in diverse drain current and ΔImax values (Table S1). However, KD depends on site binding to the hydroxyl groups of graphene and provides a stable parameter for the following measurements. The G-FET was then applied to the detection of ammonia as an enzymatic reaction product from Canavalia ensiformis (jack bean) urease. Urease has its own surface charges, whereas urea is electrically neutral. Here, we first introduced 1 μM urease into the buffer solution, followed by urea to a final concentration of 400 mM. After the introduction of urea, the hole current of the G-FET decreased gradually (Figure 3a). On
where Curea, t, vmax, and KM are the urea concentration, reaction time, maximum reaction rate, and Michaelis constant of urease, respectively. The values of vmax for urease from C. ensiformis and H. pylori were measured to be 67 s−1 and 12 amol/ (bacterial cell·s) by the indophenol method under the experimental conditions (100 mM sodium phosphate buffer, pH 8.0), respectively. Also, the values of KM for urease from C. ensiformis and H. pylori have been reported to be 3 mM39 and 0.3 mM,40 respectively. The initial value of Curea was 400 mM. Therefore, Curea was sufficiently high throughout the experiment (approximately 10 min) to keep the turnover of urease close to vmax. Additionally, the relatively low responsivity of the G-FET to ammonia in 100 mM order was a cause of the reduced change in the current during the later period of the reaction. Consequently, eq 2 can be modified to ΔI =
the other hand, the pH of the assay solution showed a different trend of its changes (Figure S2). It rapidly increased immediately after the introduction of urea, probably owing to the buffering capacity of the solution. This difference also indicates that the G-FET responded to ammonia and not to protons in the experiment. The time course of the drain current is described in terms of the enzymatic reaction kinetics and Langmuir adsorption isotherm as follows. The urease reaction is represented schematically as (3)
that is, two ammonia molecules are generated from a single urea molecule. According to the Michaelis−Menten equation, the urease reaction rate is given by −
dC NH3 dCurea v C =2 = max urea dt dt KM + Curea
(
KD 2Cureasevmax
+t
)
(5)
where Curease is the urease concentration. The time course of the drain current was well fitted by this function (Figure 3). This shows that the G-FET monitored the enzymatic reaction kinetics via the detection of the reaction product in real time. vmax estimated from the fitting curve was 22.5 s−1 at a fixed KD of 1.30 mM. The vmax values obtained from the indophenol method as the end point assay and the G-FET method as the rate assay were different. The difference can be attributed to insufficient diffusion of the substrate in the early period of the reaction. Similarly to the bulk-scale measurement using 200 μL buffer solution, a decrease in the drain current was also observed in a 78.5 fL droplet (Figure 3b) after rapidly mixing urease and urea on the G-FET and sealing the microdroplet with air. vmax was estimated to be 30.6 s−1. This result shows that the enzymatic reaction was retained and monitored in a reaction volume billions of times smaller. Next, biotinylated urease was captured on a streptavidinmodified graphene channel and its reaction was monitored in a microdroplet (Figure 4a). After urease capture, a microdroplet containing urea was repeatedly formed on the G-FET. The drain current did not change before microdroplet formation but decreased immediately after the formation. The amount of captured urease was extremely small and its reaction product was not detected before it accumulated in femtoliter volume. After droplet formation, the ammonia product decreased the drain current similarly to that in Figure 3b. When the buffer
Figure 3. G-FET response to urease in (a) 200 μL solution and (b) 78.5 fL droplet. In (b), the drain current was disturbed by the exchange of the buffer. Blue curves represent fittings by eq 5 to the drain current after urea addition. Back-gate voltage was fixed at 20 V.
(NH 2)2 CO + 2H 2O → 2NH3 + CO2 + H 2O
ΔImaxt
(4)
Figure 4. G-FET response to biotinylated urease captured on graphene channel in femtoliter microdroplet. (a) Repeated responses of G-FET to enzymatic reaction. Blue curves represent fittings by eq 5 to the drain current after each sealing of the droplet. The drain current was disturbed by the exchange of the buffer. (b) Detection of 4 pM urease using G-FET. The drain current at the sealing of the droplet (t = 175 s) was subtracted from all data. The blue curve shows the fitting by eq 5 to the drain current after sealing the droplet. 4006
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0.04 bacterial cells (ammonia production rate: 540 zmol/s, 1.7 μM/s) in 2.3 × 102 bacterial cells/mL solution and 0.12 bacterial cells (ammonia production rate: 1.50 amol/s, 4.8 μM/s) in 2.3 × 103 bacterial cells/mL solution. This result shows that the G-FET with the femtoliter microdroplet successfully monitored a very low concentration and a very small number of H. pylori cells. H. pylori solution contains many fragments containing urease that are smaller and diffuse more rapidly than the whole bacterial cell. Also, in this study the size of the graphene channel and microwells was comparable to that of the bacterial cells. Therefore, the graphene surface was considered to be occupied by these small fragments containing urease. Individual bacterial cells can be captured on a G-FET and quantized urease activity can be monitored by using a purer sample without the fragments and a larger device with graphene synthesized by chemical vapor deposition. The enzyme assay using a G-FET equipped with femtoliter microwells overcomes the limitation of Debye screening, and it is a powerful tool for high-sensitivity and rapid electrical biosensing under physiological conditions. It is currently difficult to control the quality of graphene, such as defect density, and also to control the channel size of the exfoliated graphene. In fact, in this study the initial and final drain currents varied considerably among the G-FET devices. Nevertheless, coefficient of variation (CV) for the estimated amount of urease and H. pylori was suppressed to as small as 1−4% (Table S1). It is because the amount of the detection target was determined not by the amount of the drain current but by the temporal changes in the drain current. The combination of a G-FET and a microdroplet achieved not only high sensitivity in the zeptomole range but also the quantitative measurements by using enzymatic reactions monitored in real time and in a well-defined volume on the basis of the microscopic site binding of enzymatic products to graphene. This unique feature of the lab-on-a-graphene-FET will contribute to the robust and accurate sensor readout and therefore to practical application of the graphene device. Other enzymatic products with an affinity to graphene are expected to be found in the future. Also, in combination with a second antibody using urease, the applicability of this method can be further expanded to nonenzyme targets. In this article, we demonstrated electrical biosensing under physiological conditions. We used graphene, a surface-sensitive two-dimensional material, and a femtoliter microdroplet, which effectively regulates the reaction near the graphene surface. The combination of the G-FET and microfluidics was named a “lab-on-a-graphene-FET”. It overcame the limitation of Debye screening by the real-time monitoring of the enzymatic reaction based on the site-binding model for enzymatic products in a well-defined reaction volume. The lab-on-agraphene-FET quantitatively detected urease and H. pylori at physiological ionic strength, even when they were at a distance of more than λD from the graphene channel. Two hundred seventy zeptomoles of biotinylated urease and H. pylori corresponding to 0.04 bacterial cells were successfully detected within 30 min. Our method is expected to expand the applicability of electrical biosensing under physiological conditions using the unique two-dimensional material of graphene. Methods. Materials. Urease from C. ensiformis was purchased from Toyobo Co. Ltd., Japan. Biotinylated urease was synthesized using biotinyl-N-hydroxysuccinimide ester
was injected to rupture the microdroplet on the G-FET, the drain current recovered to its initial value. This is because the accumulated ammonia was released upon the microdroplet rupture. After the formation of a second droplet, it decreased again in the same manner as before. These results show that biotinylated urease was successfully captured on the graphene and that its ammonia product was stably monitored by the GFET owing to the reproducible formation of the microdroplet. The amount of captured urease was estimated using eq 5 and the microdroplet volume of 78.5 fL. vmax for biotinylated urease was 3.1 s−1, measured by the indophenol method, and the amount of captured urease was estimated to be 360 zmol (2.2 × 105 molecules). Two hundred seventy zeptomoles (1.6 × 105 molecules) of urease was also detected using the same procedure with a low concentration (4 pM) of urease (Figure 4b). These results show that the G-FET quantitatively measured a small amount of enzyme captured on a graphene channel in a microdroplet with a well-defined volume by realtime monitoring of enzymatic reaction kinetics. Finally, the highly sensitive enzyme assay using the G-FET with microwells was applied to the detection of H. pylori. Before the G-FET experiment, H. pylori was detected by the indophenol method and by using a conventional test kit (Figure S3). The strip test barely detected 1 × 105 bacterial cells in 2.3 × 107 bacterial cells/mL solution after a 1 h reaction. The detection limit of the indophenol method was determined to be 2 × 103 bacterial cells in 2.3 × 104 bacterial cells/ml solution after 4 h reaction. On the other hand, in the G-FET experiment the urease activity of H. pylori in 2.3 × 102 bacterial cells/ml solution was monitored immediately after microdroplet formation (Figure 5). Including the time for H.
Figure 5. H. pylori detection using G-FET with femtoliter microdroplet. The drain current value at the sealing of the droplet (t = 0 s) was subtracted from all data. Dark blue and red curves represent fittings by eq 5 to the drain current values for 2.3 × 103 and 2.3 × 102 bacterial cells/mL solution, respectively.
pylori binding to the antibody on the G-FET (15 min), the entire assay was completed within 30 min. The size of IgG for H. pylori is approximately 15 nm,14 which is much larger than λD under the experimental conditions, that is, less than 1 nm. H. pylori bound to the antibody was detected beyond the limitation of Debye screening by its enzymatic reaction. Also, note that although urease reaction changed the pH in the microdroplet after sealing, which degraded the binding affinity of IgG to H. pylori, H. pylori and its reaction products remained in the microdroplet and the measurement was not disturbed. Further, we measured the reaction time course after target binding, and therefore, nonspecific adsorption on the G-FET did not affect the measurement. From the fitting of the draincurrent decay, the amount of captured urease corresponded to 4007
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(Sigma-Aldrich Co., LLC, U.S.A.) and purified by gel filtration. The concentrations of urease and biotinylated urease were determined by bicinchoninic acid assay.41 1-Pyrenebutanoic acid succinimidyl ester (PBASE), streptavidin, Rhodamine 6G, and urea were purchased from Sigma-Aldrich, Nacalai Tesque, Inc. (Japan), Tokyo Chemical Industry Co., Ltd. (Japan), and Fujifilm Wako Pure Chemical Corp. (Japan), respectively. H. pylori (JCM12093) was provided by RIKEN BRC through the National Bio-Resource Project of MEXT, Japan. An anti-H. pylori antibody was purchased from Kirkegaard & Perry Laboratories, Inc. (U.S.A.). Sodium phosphate buffer (100 mM, pH 8.0) was used as a buffer solution. Ammonia sensing shown in Figure 2 was carried out in pure water to eliminate any effect of other electrolytes. Device Fabrication. Graphene was exfoliated from kish graphite and transferred onto a p-type Si substrate with a 290 nm thick SiO2 layer. Source and drain electrodes (Au/Ni, 30 nm/10 nm) were fabricated on the substrate by electron-beam lithography, electron-beam deposition and a lift-off process. The channel length was 4 μm. The G-FET chip surface was sequentially coated with 30 nm thick Al2O3 by atomic-layer deposition and 4 μm thick perfluoropolymer (CytopTM, Asahi Glass Co. Ltd. (Japan)) by spin-coating. Following the photolithography of a microwell pattern on the G-FET, the perfluoropolymer was etched using oxygen plasma.42,43 Microwells were completed after resist removal and alkaline etching of Al2O3. The microwells fully covered the graphene channel and had a diameter of 5 or 10 μm and a depth of 4 μm, and therefore a volume of 78.5 or 314 fL. The 10 μm diameter microwells were used only for H. pylori detection (Figure 1). Graphene Surface Modification. In experiments where enzyme molecules or H. pylori were captured on the graphene channel, the graphene surface was modified by receptor molecules via linkers. First, the graphene channel was immersed in 500 μM PBASE linker in 2-methoxyethanol for 1 h at room temperature. Then, the graphene channel was briefly rinsed with the buffer solution and immersed in 200 nM streptavidin or 20 μg/mL antibody in the buffer overnight at 4 °C. In the modification of the antibody, 100 mM sodium phosphate buffer (pH 6.0) was used. A silicone rubber reservoir was placed on the device to store the sample solution. Enzyme Assay. To capture biotinylated urease and H. pylori on graphene before enzyme assay, 200 μL of their solution was held on the G-FET at room temperature for 15 min. Then the solution was changed to the reaction buffer with 400 mM urea. Immediately after the change, microdroplets were formed by an airflow through the slow suction of the buffer from the rubber reservoir. Electrical Measurement Setup. All electrical measurements of the G-FET were carried out using a semiconductor device analyzer (B1500A, Keysight Technologies, Inc., U.S.A.). The drain voltage was fixed at 0.1 V and gate voltage was applied from the back side of the G-FET chip. The drain current was monitored at a fixed gate voltage of 0 V unless otherwise mentioned. Our graphene on SiO2 was generally hole-doped as reported elsewhere (Table S1),44,45 and the value of the gate voltage was chosen to enable monitoring of the hole current.
Letter
ASSOCIATED CONTENT
S Supporting Information *
The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.nanolett.9b01335.
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Effect of ammonia in water on the G-FET characteristics, electrical characteristics of the devices, pH changes in urease reaction, urease activity test, and binding assay of H. pylori (PDF)
AUTHOR INFORMATION
Corresponding Authors
*E-mail:
[email protected],. *E-mail:
[email protected]. ORCID
Takao Ono: 0000-0002-6717-123X Author Contributions
T.O. and K.M. conceived and designed the study. T.O., Y.W., S.N., T.K., and Y.S. carried out the experiments using H. pylori. T.O. carried out other experiments. T.O., Y.K., and K.I. analyzed the data. T.O. and K.M. wrote the manuscript with input from other authors. Notes
The authors declare the following competing financial interest(s): T.O. and K.M. of Osaka University hold a patent (application number 2016-169810) on the method for measuring the enzymatic activity in a microdroplet using the G-FET.
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ACKNOWLEDGMENTS This work was supported by JST-CREST (JPMJCR15F4). T.O. acknowledges support from JSPS KAKENHI (16K13638 and 18K14107).
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REFERENCES
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DOI: 10.1021/acs.nanolett.9b01335 Nano Lett. 2019, 19, 4004−4009
