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Flexible Organic Electrochemical Transistors for Highly Selective Enzyme Biosensors and Used for Saliva Testing Caizhi Liao, Chunhin Mak, Meng Zhang, Helen L. W. Chan, and Feng Yan*

Organic thin-film transistors (OTFTs) have emerged as the state-of-the-art potentiometric sensing platform for flexible, versatile, and disposable biosensors.[1–9] Recently, OTFT-functionalized sensors have been successfully used in the detection of specific analyte in the complex multianalyte environments or real physiological samples,[7] which holds the great potential for point-of-care applications in the future. Organic electrochemical transistors (OECT) are a type of OTFT with a simple structure, in which a thin layer of organic semiconductor is deposited on the channel area between the source and drain electrodes and exposed to an electrolyte together with the gate electrode.[10–12] Thanks to the inherent amplification capability,[13] OECTs are highly sensitive transducers that can convert biochemical signals into electronic ones and have been successfully used as many types of sensors, including ions,[14–16] lactate,[17] glucose,[12,18] dopamine,[19,20] DNA,[21] bacteria,[22] protein,[23] and cells,[2,24] etc. OECTs can be easily fabricated by solution processes, like printing or spin-coating techniques, which make the devices potentially useful in some emerging areas, such as wearable electronics, electronic skin, and body-integrated implantable sensors.[25,26] The working mechanism of OECTbased biosensors is detailed in the Supporting Information. Enzyme biosensors for monitoring various substances, such as glucose, uric acid (UA), cholesterol, and lactic acid, have broad potential applications in clinical diagnostics, healthcare testing, food safety, and environmental monitoring, etc. Therefore, it is necessary to develop OECTs as a novel platform for high-performance enzyme biosensors. In practical applications of the biosensors, of particular importance is the selectivity that however has been rarely studied in OECT-based devices until now.[4,5] In this paper, considering that most enzyme biosensors are based on the detection of H2O2, we report highly selective OECT-based H2O2 sensors by modifying the gate electrodes with a bilayer film that can effectively block the interference from charged biomolecules in solutions and only allow H2O2 to diffuse to the gate. Then, highly sensitive and selective enzyme biosensors for different analytes were realized by immobilizing suitable enzymes on the surface of the bilayer film. UA is the end product of purine metabolism in human body. The basal concentrations of UA in physiological samples vary

C. Liao, C. Mak, M. Zhang, Prof. H. L. W. Chan, Prof. F. Yan Department of Applied Physics The Hong Kong Polytechnic University Hung Hom, Kowloon, Hong Kong E-mail: [email protected]

DOI: 10.1002/adma.201404378

Adv. Mater. 2014, DOI: 10.1002/adma.201404378

across a wide range from 1 × 10−7 to 1 × 10−3 M, depending on the matrix.[27] Abnormal UA level in human body is a sign for many diseases, including gout, hyperuricemia, Lesch– Nyhan syndrome, renal diseases, and even cardiovascular disorder.[28–30] Consequently, fast and accurate determination of UA is critical to the diagnosis and treatment of these diseases. However, the conventional techniques for UA measurements usually suffer from several drawbacks like labor-intensive, expensive, time-consuming, insensitive, and unportable.[31–37] Here, we studied OECT-based UA sensors with the device structure shown in Figure 1a for the first time. The UA sensors demonstrated good selectivity and a detection limit of ≈1 × 10−8 M, which is approximately three orders of magnitudes better than conventional electrochemical methods using the same enzyme electrodes. Other high-performance enzyme biosensors, including cholesterol and glucose sensors, were realized with the same strategy. In the end, the devices were demonstrated for saliva testing and the concentrations of UA and glucose in saliva can be selectively measured. Flexible OECTs with Pt electrodes and poly(3,4-ethylene dioxythiophene):poly(styrene sulfonate) (PEDOT:PSS)) active layers were fabricated on thin polyethylene terephthalate (PET) substrates with thicknesses of about 50 µm. The flexible OECTs can be attached on various deformable surfaces and show stable performance during bending tests up to 1000 times, as shown in Figure 2a,b. The OECTs with Pt gate electrodes are sensitive to H2O2 due to the following reaction on the gate:[27] H2 O2 → O2 + 2H+ + 2e − However, the Pt gate is also sensitive to many other analytes, such as dopamine (DA), glucose, UA, and ascorbic acid (AA), etc.[12,19] An OECT with a bare Pt gate electrode shows the detection limits (signal-to-noise ratio ≥ 3) of 1 × 10−9 M, 5 × 10−9 M, 0.1 × 10−6 M, 1 × 10−6 M, and 5 × 10−6 M to H2O2, DA, AA, UA, and glucose, respectively (see Supporting Information, Figure S1). To improve the selectivity of the devices to H2O2, the Pt gate electrodes of the OECTs were modified with a thin layer of the composite of graphene flakes and Nafion, followed by a thin layer of polyaniline (PANI) conducting polymer. The use of graphene flakes is to improve the electrocatalytic activity and the conductivity of the gate electrode.[38] Figure 2c shows the channel current responses of an OECT with a PANI/Nafion-graphene/Pt gate characterized in PBS solution with additions of H2O2. The channel current of the OECT shows clear responses to additions of H2O2 and a detection limit of 3 × 10–9 M.[12,19] More importantly, we find that the selectivity of the device was dramatically improved after the gate modification. The detection limits of the OECT to DA and

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Figure 1. a) Schematic diagram of an OECT with a UOx-GO/PANI/Nafion-graphene/Pt gate. b) Potential drops between the gate and channel of the OECT before (solid line) and after (dash line) the addition of UA in the electrolyte (PBS solution).

AA are 3 × 10−6 M, which are three orders of magnitude higher than that to H2O2 (see Supporting Information, Figure S2). The device cannot show any response to the additions of UA and glucose at the concentrations up to 100 × 10−6 M (see Supporting Information, Figure S3). It is notable that the detection limits to H2O2 before and after the gate modification are very similar, while those to other analytes are dramatically increased after the gate modification. In sum, the PANI/Nafion-graphene bilayer film can greatly improve the selectivity of the OECT to H2O2. The PANI film covered on the surface of Pt gate electrode is in the H+ protonated emeraldine salt form,[39] which can

strongly repel the positively charged molecules such as DA in PBS solution by electrostatic forces.[40] On the other hand, Nafion,[41] an acid with a stable Teflon backbone and acidic sulfonic groups, is negatively charged in PBS solution, and could effectively impede the diffusion of the anionic electro-active substances, like AA and UA, through it by electrostatic interactions. In addition, it will be difficult for big biomolecules like glucose to pass through the nanochannels in PANI and Nafion films.[41] The size of H2O2 molecule is the smallest in all of the analytes and it can pass through the multilayer film most conveniently. Therefore, the PANI/Nafion-graphene bilayer

Figure 2. a) Photographs for flexible OECTs attached on different surfaces. b) The transfer characteristics of an OECT measured in PBS solution after the bending tests up to 1000 times. VDS = 0.05 V. c) Channel current responses of an OECT with PANI/Nafion-graphene/Pt gate electrode to additions of H2O2 in PBS solution. Inset: transfer curve of the device characterized in PBS solution. d) The changes of effective gate voltage (ΔVGeff ) versus the concentrations (C) of H2O2, AA, and DA.

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C ⎞ kT ⎛ VGeff = 2.30 ⎜ 1 + E−C ⎟ log[H2O2 ] + C ⎝ CG−E ⎠ 2q

(1)

where k is Boltzmann constant, T is temperature, q is the electron charge, [H2O2] is the concentration of H2O2, and C is constant. CG-E and CE-C are the capacitances of the two electric double layers (EDL) close to the gate and the channel (Figure 1b), respectively. We can read out the VGeff corresponding to different channel currents from the transfer curve of the device. As shown in Figure 2d, the effective gate voltage change, ΔVGeff, can be fitted with Equation 1 in a broad range of [H2O2] from 10−8 to 10−5 M in logarithmic axis, being similar to many typical transistor-based biosensors.[43–45] Next, UA sensors were realized based on the H2O2-sensitive OECTs with the device structure as shown in Figure 1a. The Pt gate electrodes of the OECTs were modified with PANI/ Nafion-graphene bilayer films by solution process and then the enzyme uricase (UOx) was immobilized on the surface of PANI.[27] UOx on the surface can catalyze the oxidation of UA to allantoin, H2O2, and CO2 in the following reaction: UA + 2H2O + O2 ⎯→ Allantoin+H2O2 + CO2 Various enzyme immobilization approaches, either physically or chemically, have been used to immobilize UOx onto the surface of the gate electrodes.[46] Firstly, the gate electrode of an OECT was immersed in UOx solution for 4 h at 4 °C to immobilize the enzyme by physical absorption. As shown in Figure 3a, the OECT demonstrates a detection limit of 300 × 10−9 M to UA and the VGeff increases for 143 mV when the concentration of UA is increased for one decade. Because the device is not sensitive to UA before the coating of UOx, the sensing mechanism of the device is based on the detection of H2O2 generated in the reaction of UA biocatalyzed by UOx. However, we observed that the device is unstable for long-term operations in solutions presumably due to the weak van der Waals force between the enzyme and PANI layer in absorption.[47] Chemical methods involved with covalent bonding have emerged as promising approaches for enzyme immobilization.[48] UOx was then immobilized on the gate electrode of another OECT by using a cross-linker glutaraldehyde, which is a small organic molecule with functional groups at both ends.[49] Figure 3c,d demonstrate the responses of the OECT with the UOx-glutaraldehyde/PANI/Nafion-graphene/Pt gate electrode to different UA concentrations. The device shows a much improved sensitivity and a low detection limit down to 30 × 10−9 M, which is one order of magnitude better than that of the device based on physical absorption. The significantly improved sensitivity can be attributed to the more effective enzyme immobilization on the gate electrode associated with covalent bonding.

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film can block the diffusion of both positively and negatively charged molecules to the Pt gate and improve the selectivity of the device to H2O2. The channel current response can be attributed to the change of the effective gate voltage (VGeff) of the OECT due to the reaction of H2O2 on the gate electrode and VGeff is given by (see the Supporting Information):[5,18,42]

2D nanomaterials have also emerged as promising matrix materials for the effective enzyme immobilization.[50] Graphene oxide (GO) is a 2D graphene sheet with oxygen-containing functional groups anchored on the surface of sheet. These functional groups can readily react with the amine groups of the protein enzyme and the reactive moieties of the conductive PANI layer, leading to the optimized enzyme immobilization via the effective chemical covalent bonding.[51] In addition, the physical van der Waals forces between the benzene rings of the PANI and the GO sheets can also improve the enzyme-anchoring effect on the working electrode.[52] As shown in Figure 3e,f, the UA sensor shows the detection limit of 10 × 10−9 M and the effective gate voltage shift (ΔVGeff) of 147 mV per decade of UA concentration, which are better than those of the aforementioned two type UA sensors. The device shows a good linearity in a wide range of UA level (100 × 10−9 M to 500 × 10−6 M), which covers the normal levels of UA in human body.[27] The remarkably improved sensitivity can be ascribed to the introduction of GO for the efficient enzyme immobilization. It is notable that the sensitivity of the organic device is much higher than those of the UA sensors based on inorganic field effect transistors because of the ultrahigh transconductance of OECTs.[13,27,55] The impressively high sensitivity of the OECT-based UA sensor even can guarantee a sufficient response to a trace amount of the analyte in highly diluted biological samples, in which the concentrations of most interferents are diluted to a negligible level. In addition, the device able to maintain a stable performance after several weeks of storage in refrigerator (4 °C). The selectivity of the OECT-based UA sensor is key to practical applications.[53,54] Human body fluids is a complex biological system containing many interferents that may cause significant error signals in UA detections. As shown in Figure 3f, the responses of the device to UA are much bigger than those to the common interferents, including glucose, AA, and DA (see Supporting Information, Figure S4), which is consistent with the excellent selectivity of the device to H2O2 as shown in Figure 2d. Typically, a transistor-based sensor is a combination of a sensor and an amplifier.[5] The high transconductance of the OECTs (approximately 200 µS, when VDS = 0.05 V) can significantly enhance the signals induced by analytes.[13] For a comparative study, we also performed the conventional cyclic voltammograms (CV) and amperometric measurements on the UOx-GO/PANI/Nafion-graphene/Pt electrodes (see Supporting Information, Figure S5), which show a detection limit of around 3 × 10−6 M. So the OECT-based UA sensor is much more sensitive than conventional electrochemical methods. To the best of our knowledge, the concept of using a positively/negatively charged bilayer film to improve the selectivity of electrochemical biosensors has never been reported before. To better understand the effect of the bilayer film, the amperometric measurements of the UOx-GO/PANI/Nafiongraphene/Pt electrode were also performed in AA, DA, and H2O2 solutions and showed the detection limits of 3 × 10−6 M, 1 × 10−6 M, and 1 × 10−6 M, respectively (see Supporting Information, Figure S6), indicating that the bilayer film cannot improve the selectivity of the enzyme sensor in amperometric measurements. It is reasonable to find this result because

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Figure 3. a) Current responses and b) the corresponding effective gate voltage changes of an OECT with a UOx/PANI/Nafion-graphene/Pt gate electrode to the additions of UA with different concentrations. c) Current responses and d) the effective gate voltage change of an OECT with UOxglutaraldehyde /PANI/Nafion-graphene/Pt gate electrodes to the additions of UA. e) Current responses and f) the effective gate voltage change of an OECT with UOx-GO/PANI/Nafion-graphene/Pt gate electrodes to the additions of UA. VDS = 0.05 V and VG = 0.6 V. Insets: transfer curves of the devices characterized in PBS solution.

the amperometric measurements need to have high enough redox currents, which should be maintained by electric field in the electrolytes to drive the diffusion of the analytes to the working electrodes.[42] Consequently, the electrostatic interactions between the bilayer film and the charged molecules, such as AA and DA, can be overcome by the electric field close to the interface in amperometric measurements. Therefore, the blocking effect of the bilayer film to charged molecules is pronounced exclusively in potentiometric transducers, like OECTs. Similarly, highly sensitive and selective cholesterol sensors and glucose sensors have been realized with the same strategy. The enzyme cholesterol oxidase was immobilized on the PANI/Nafion-graphene/Pt gate electrode of an OECT together with GO for cholesterol sensing (see Supporting Information, Figure S7 and Table S1).[56] The cholesterol sensor shows a low detection limit of 100 × 10−9 M, which is sensitive enough to 4

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detect cholesterol levels in human body. Since the normal range of cholesterol level is 1.3–2.6 mg mL–1 (3.36–6.72 × 10−3 M) in the normal human plasma, accurate detection of cholesterol can be realized by diluting human plasma to a lower concentration, which can also eliminate the interference from many other biomolecules. Glucose sensors were realized by immobilizing the enzyme glucose oxidase (GOx) on the PANI/Nafiongraphene/Pt gate electrode by using GO.[27] The devices show a detection limit of about 30 × 10−9 M (see Supporting Information, Figure S8 and Table S1), which is two orders of magnitude better than that of the device to AA and DA (≈3 × 10−6 M). The glucose sensors also show an excellent selectivity for practical applications. Promisingly, highly sensitive glucose sensors can be used for sensing glucose levels in body fluids such as saliva, which may provide a viable way for non-invasive glucose detection.

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Saliva testing is a convenient, safe, and cost-effective way to monitor several diseases.[57,58] For example, glucose level in saliva is much higher in diabetic patients than that in healthy persons.[59] Salivary uric acid is a non-invasive biomarker for metabolic syndrome, cardiometabolic risk, and end-stage renal disease.[60,61] To demonstrate the potential applications of the enzyme sensors for saliva testing, different devices were prepared and used for testing the uric acid level and glucose level in saliva. First, an UA-sensitive OECT with a UOx-GO/PANI/ Nafion-graphene/Pt gate electrode was characterized in PBS solution and then saliva was added in the PBS solution with different volumes. Figure 4a shows the response of the OECT to the additions of saliva. Then, the UA sensor was calibrated in UA PBS solution to obtain the relationship between UA concentration and the shift of effective gate voltage ΔVGeff. The concentration of UA in the saliva sample was then calculated to be (173 ± 20) × 10−6 M in three measurements, being consistent with the UA level in human saliva reported before.[62] In a similar way, a glucose-sensitive OECT with the GOx-GO/ PANI/Nafion-graphene/Pt gate electrode was characterized in PBS solution before and after the additions of the saliva sample with different volume. As shown in Figure 4b, the device shows significant responses to the additions of saliva. After the calibration of the device in glucose PBS solution, the glucose

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Figure 4. a) Channel current responses of an UA-sensitive OECT with a UOx-GO/PANI/Nafion-graphene/Pt gate electrode characterized in PBS solution before and after the additions of saliva of different volumes. Inset: transfer characteristic of the OECT measured in PBS solution. b) Channel current responses of a glucose-sensitive OECT with a GOx-GO/PANI/ Nafion-graphene/Pt gate electrode characterized in PBS solution before and after the additions of saliva of different volumes. Inset: transfer characteristic of the OECT measured in PBS solution.

concentration in saliva was estimated to be (103 ± 10) × 10−6 M, which is in a reasonable region for a healthy person.[63] To prove that the enzyme-based OECT is selectively sensitive to the targeted analytes, a control OECT without enzyme on the PANI/Nafion-graphene/Pt gate electrode was characterized in PBS solution with additions of the same saliva sample. We noticed that the device did not show any response even when 100 µL of saliva were added into 10 mL of PBS solution (see Supporting Information, Figure S9), indicating that the OECT without enzyme on the gate surface was not sensitive to UA and glucose, etc., in saliva. Thus, the saliva testing by the use of enzyme-modified OECTs shows excellent selectivity and sensitivity simultaneously. To our best knowledge, this is the first work to demonstrate the selective and direct measurements of glucose level and uric acid level in saliva by a convenient electrochemical method. In summary, the flexible OECTs with the gate electrodes modified with a PANI/Nafion-graphene bilayer film show excellent selectivity to H2O2, considering that the interferences are extensively blocked by the positively/negatively charged bilayer film due to electrostatic interactions. The devices are then modified with suitable enzymes and GO on the gate electrodes and successfully used for sensing UA, cholesterol, and glucose with low detection limits and good selectivity. In principle, other types of enzyme biosensors can be conveniently realized with the similar approach. Considering that most of the cells and biomolecules, such as DNAs, proteins, UA, AA, and DA, have positive or negative charges in aqueous solutions, the bilayer film modified on the gate electrodes can dramatically improve the selectivity of the OECT-based biosensors in many applications. The biosensors have been successfully used for sensing UA and glucose levels in saliva, indicating that the devices are promising for non-invasive detection of bio-markers in human body.

Supporting Information Supporting Information is available from the Wiley Online Library or from the author.

Acknowledgements This work was financially supported by the Research Grants Council (RGC) of Hong Kong, China (project number: N_PolyU506/13) and The Hong Kong Polytechnic University (project numbers: G-SB07, A-PL49, and G-YM45). Received: September 23, 2014 Revised: October 25, 2014 Published online:

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Adv. Mater. 2014, DOI: 10.1002/adma.201404378

Flexible organic electrochemical transistors for highly selective enzyme biosensors and used for saliva testing.

Flexible organic electrochemical transistors (OECTs) are successfully used as high-performance enzyme biosensors, such as uric acid (UA) and cholester...
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