review OTFTs

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Organic Semiconductors in Organic Thin-Film Transistor-Based Chemical andBiological SensorsCaizhi Liao a & Feng Yan aa Department of Applied Physics , Hong Kong Polytechnic University ,Kowloon , Hong Kong , ChinaPublished online: 08 Aug 2013.

To cite this article: Caizhi Liao & Feng Yan (2013) Organic Semiconductors in Organic Thin-Film Transistor-Based Chemical and Biological Sensors, Polymer Reviews, 53:3, 352-406, DOI:10.1080/15583724.2013.808665

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Polymer Reviews, 53:352–406, 2013Copyright © Taylor & Francis Group, LLCISSN: 1558-3724 print / 1558-3716 onlineDOI: 10.1080/15583724.2013.808665

Organic Semiconductors in Organic Thin-FilmTransistor-Based Chemical and Biological Sensors

CAIZHI LIAO AND FENG YAN

Department of Applied Physics, Hong Kong Polytechnic University, Kowloon,Hong Kong, China

The dramatic progress in developing organic semiconductors with many virtues, in-cluding flexibility, high conductivity, solution processability, and biocompatibility, hasaroused great research interest recently. Organic thin-film transistors (OTFTs), beingan important type of organic semiconductor devices, have been successfully used in var-ious high-performance sensors. This review will give a comprehensive description onthe state-of-the-art chemical and biological OTFT-based sensors, including humidity,pH, ions, glucose, DNA, antibody-antigen, dopamine, bacteria, and cell-based sensors.Several essential aspects in this research field, including the properties of organic semi-conductors, device physics, and the performance of different types of sensors, will beintroduced in details.

Keywords organic semiconductor, OTFT, OFET, OECT, chemical sensor, biosensor

1. Introduction

1.1 Organic Semiconductors

Since the serendipitous discovery of the highly conductive polymer polyacetylene by Shi-rakawa et al. in the late 1970s,1 organic electronics has become the “new darling” of theacademic community and the industry. As the essential materials for the organic electronics,organic semiconductors have attracted huge attention for over three decades, and emergedas the commercial alternative to conventional inorganic semiconductors such as silicon.Unlike inorganic materials, the organic semiconductors have the physical and chemicalproperties that can be easily tailored by the incorporation of functional groups or the ma-nipulation of physical conditions to meet the specific requirements. The versatility of thesematerials affords a brand new era for the development of novel electronic devices andpoises to revolutionize the society. Numerous π -conjugated organic semiconductors, in-cluding poly(3-hexylthiophene) (P3HT),2 polyaniline (PANI)3 and polypyrrole,4 etc., havebeen synthesized and successfully used in organic electronic devices (Fig. 1).

One advantage of the organic semiconductors over the inorganic ones is the low re-quirements of processing conditions. Most inorganic semiconductor devices are preparedon the high-purity crystalline substrates at high temperature, normally relying on the uti-lization of expensive facilities in clean rooms. In contrast, organic electronic devices can

Received March 27, 2013; accepted May 14, 2013.Address correspondence to Feng Yan, Department of Applied Physics, Hong Kong Polytechnic

University, 11 Yucai Road, Kowloon, Hong Kong, China. E-mail: [email protected]

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Organic Semiconductors in OTFT-Based Sensors 353

Figure 1. Chemical structures of the organic semiconductors commonly used in OTFT-based sensors.(a) Poly(3,4-ethylenedioxythiophene):poly(styrene sulfonic acid) (PEDOT:PSS). (b) Polypyrrole.(c) 1,4,5,8-naphthalene-tetracarboxylicdianhydride (NTCDA). (d) Poly(3-hexylthiophene) (P3HT).(e) Polyaniline. (f) Pentacene. (g) Polycarbazole. (h) 5,5′-bis-(7-dodecyl-9H-fluoren-2-yl)-2,2′-bithiophene (DDFTTF). (i) Copper phthalocyanine (CuPc).

be fabricated by the convenient solution-processable techniques or thermal evaporationat room-temperature, enabling the integration of devices on low-cost substrates such asglass, flexible plastics, metal foils, and even papers.5 For instance, one of the mostly re-searched alkyl substituted polythiophene, P3HT,6 shows excellent solubility in a variety oforganic solvents, which affords the easy fabrication of the devices by solution-processabletechniques, including spin-coating, inkjet printing, or screen printing, etc.5,7

Organic semiconductors normally have conjugated-molecular structures and exhibitthe electronic and optical properties similar to that of their inorganic counterparts.8 Electron-donating organic semiconductors involving high and highest occupied molecular orbital(HOMO) levels are good candidates for p-type semiconductors, while electron-acceptingones with low HOMO levels are normally used as n-type semiconductors. Both types of thematerials are important to this area. Organic semiconductors, including polymer semicon-ductors and small molecule organic semiconductors, have been successfully used in manysensing applications. Polymer semiconductors normally contain a π - conjugated backboneand side chains that can increase the solubility of these materials in solvents. Therefore,most of the polymer-based devices can be fabricated by solution process. Some conjugatedpolymers are referred to as “synthetic metal,” primarily due to their electronic properties(e.g., high electrical conductivity) similar to that of metals.9 The intrinsically conductive

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354 C. Liao and F. Yan

polymers have also been widely used in the sensing applications by virtue of their inher-ent charge transport properties.10,11 Most of the small molecule organic semiconductorscan form ordered crystal structures. So, some of the materials have the carrier mobilitiescomparable to or even higher than that of amorphous silicon.12,13 Small molecule organicdevices can be prepared by thermal evaporation and are thus compatible with high through-put fabrication techniques.5,14 Compared with their polymer counterparts, small moleculeorganic semiconductors exhibit some other advantages, including well-defined molecularstructure, controllable molecular-weight, and high purity.15,16

Nearly all organic semiconductors rely on the π -conjugated systems for electrical con-duction, in which the continuous sp2 hybridized carbon centers along the molecules cangive rise to partially delocalized π -orbital states that will significantly affect the electricaland optical properties.17 Owing to the complexity of the conduction mechanisms in suchmaterials, organic semiconductors exhibit conductivities covering a wide range of fifteenorders of magnitudes, which could be explained by quite a few models built on the carrierhopping process.18 The electrical conductivity of organic semiconductors is influenced bytwo significant factors: the carrier (electron or hole) density and the carrier mobility. Nor-mally, organic semiconductors have limited carrier densities (106 to1018/cm3) and relativelow carrier mobilities19,20 which are affected by many issues, such as the nature of thesemiconductors21 and the morphology of the solid films.22 Organic semiconductors haveweak intermolecular interactions,23 for example, van der Waals and dipole-dipole interac-tions, leading to a very low carrier mobility for the intermolecular charge transport. Toachieve a higher mobility, the intermolecular ordering should be improved to increase π -πoverlapping between the adjacent organic molecules, which can facilitate carrier transportbetween the molecules. Many methods have been reported to align the molecular chainsin organic semiconductor films for better ordering. For instance, region-regular poly(3-hexylthiophene) (P3HT) could form an ordered lamellar structure by the self-assemblyprocess and exhibits a relatively high mobility of ∼0.1 cm2 V−1 s−1.24

The rapid development of new organic semiconductors has led to the expansive growthof organic electronics, ranging from organic light-emitting diodes (OLEDs),9,25 organicphotovoltaic devices (OPVs)8,26–28 to organic thin-film transistors (OTFTs).29–32 With theemergence of more and more high-mobility organic semiconductors, much effort has beendevoted to the study of OTFTs for practical applications, such as flexible active matrixdisplays,33,34 radio-frequency identification (RFID) tags,35,36 and sensors,37,38 etc., due totheir numerous advantages and huge potential markets.

1.2 OTFT-Based Chemical and Biological Sensors

The unprecedented interest in the study of analytical techniques for the detection of specificchemical and biological species have fuelled the emergence of a large number of sensors.39

A sensor is a device that can measure a specific quantity and convert it into a signal which canbe directly observed or recorded by an external instrument. The applications of chemicalsensors and biological sensors represent a new trend in the environmental control andmedical diagnostics.11 Chemical sensors normally consist of a chemically sensitive layerand a physical transducer and can provide information of its ambient environment, whilebiological sensors typically incorporate with biological molecules, such as DNA, antibody,bacteria, etc., as the biological elements for sensing analytes.40 Among the reported sensingapproaches, field-effect transistors (FETs) have attracted increasing interest for its highsensitivity, small size, and versatility. Compared with many other types of transistors basedon silicon, graphene, carbon nanotube, or oxide semiconductors, OTFTs can be more

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conveniently prepared by the solution process, which is a cost-effective method for massproduction. More importantly, some recently reported OTFT-based sensors have exhibitedhigh sensitivity, excellent biocompatibility, and flexibility. Therefore, OTFTs are the idealcandidates for high-performance and disposable sensors.41

OTFTs are three-terminal electronic devices, including the source, drain, and gateelectrodes, with a thin film of organic semiconductor as the pivotal active channel betweenthe source and drain electrodes.42 The magnitude of the channel current flows through theorganic semiconductor and could be effectively modulated by the source-drain and gatevoltages.40 The transistor could be switched between two states, that is, the “ON” state withhigh channel current and the “OFF” state with very low channel current. To date, severaldifferent organic semiconductors, such as poly(3,4-ethylenedioxythiophene):poly(styrenesulfonic acid) (PEDOT:PSS),43 PANI,43 poly(3-alkylthiophene) (P3AT),44 and pentacene,45

have been widely used in OTFT-based sensing applications. These sensors have beensuccessfully fabricated by facile approaches, including spin-coating,46 screen printing,47

thermal evaporation, and inkjet printing,48 etc., which can largely reduce the complexity ofthe manufacture process and the cost.

OTFTs can be divided into two primary categories, namely, organic field-effect transis-tors (OFETs) and organic electrochemical transistors (OECTs), according to the operationprinciples and the device structures.40 The basic components of an OFET include a thinorganic semiconducting film as the active channel, an insulating layer as the gate dielectric,and three electrode (source, drain, and gate). The operation of the OFET is based on themodulation of carrier density in the channel by the electric field from the gate electrode.Consequently, the channel current could be changed for several orders of magnitude bythe gate voltage, indicating that the device is a type of amplifier or a transducer that canconvert a voltage/potential signal into a current response.49 In a typical OFET-based sensor,the active layer is either exposed to liquids or gases containing the analytes of interest, asshown in Fig. 2a.50 On the other hand, an OECT is also built-up by the combination ofthree electrodes (source, drain, and gate) similar to that of OFETs; however, an electrolytemedium instead of solid-state insulating layer is incorporated between the gate electrodeand the organic semiconductor channel, as shown in Fig. 2b. The conductivity of the activelayer is modulated by electrochemical doping/de-doping processes at the semiconductor-electrolyte interface, which usually involve with the migration of mobile ions into or out of

Figure 2. Schematic structures of (a) an OFET-based sensor and (b) an OECT-based sensor (Colorfigure available online).

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Figure 3. The numbers of papers published on OTFT-based chemical and biological sensors. Someworks on gas sensors and chemical vapor sensors are not included in this table. Data collected fromWeb of Science R©, on Mar 3, 2013 (Color figure available online).

the semiconducting layer.51 Since the doping/de-doping processes are reversible, the semi-conductor channel could be switched between the different doping levels by gate voltages.52

OTFT-based sensors have exhibited high sensitivities, low detection limits, and high selec-tivity, and can be easily miniaturized and integrated into portable electronic devices withsolution process.31,53 Therefore, the devices have attracted more and more attention in thepast few years. Figure 3 shows the representative papers on OTFT-based sensors, covering abroad range of applications, including environmental monitoring,54 food safety detection,55

artificial skin,56 biological warfare agents,40 medical diagnostics,57,58 and drug delivery59

in which some devices are now excitingly stepping close to commercial applications.OFET-based sensors are able to determine the target analytes due to the physical

or chemical changes in the semiconducting layers induced by the targets.32 The OFETperformance is significantly influenced by the properties of the grain boundaries and di-electric/insulator interface. Therefore, the properties and functionalities of OFET-basedsensors can be tuned via surface engineering.60,61 Compared with the OFET-based sen-sor, OECT sensors show some different properties. First, OECTs operate at much lowerworking voltages (about 1 V), which effectively prohibit the risk for hydrolysis during op-eration. More importantly, OECT-based sensors can operate in aqueous electrolytes that areessential for real-time chemical and biological sensing applications.62 In addition, OECTshave very simple structures, in which channels and gate electrodes could be fabricatedseparately.63 The above merits make OECTs to be easily integrated with complex systems,like microfluidic channels, or to be fabricates as sensor arrays for high throughput sensing.

Recently, a concept-new type OFETs, electrolyte-gated organic field-effect transistors(EG-OFETs), have emerged as a potential platform for effective sensing applications aswell.64–66 Being different from the conventional OFETs, an EG-OFET has the organicsemiconductor channel and gate electrode that are separated by an electrolyte layer.67–68

Even when a small gate voltage is applied (∼1 V), a high charge density in the channel of theEG-OFET can be induced due to the high capacitance (about tens of μC/cm2) of the electricdouble layer (EDL) at the electrolyte/ organic semiconductor interface.70–73 Although some

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reports unveiled the underlying operation mechanism of an EG-OFET being similar to thatof an OECT due to the electrochemical reaction of the active layer in EG-OFETs,74,75 westill classify the EG-OFETs as a type of OFET in the paper according to the terminology.Further work is needed to clarify the working principle of EG-OFETs. If the modulationof the channel current by the gate voltage in the device is mainly due to electrochemicaldoping instead of field-effect doping in the active layer, it is better to be regarded as atype of OECT. Many EG-OFET-based sensors have already been developed, such as DNAsensors76 and pH sensors.77 More importantly, all printed EG-OFET devices can be easilyintegrated with microfluidic systems, implying broad potential applications in the future.78

In this review, we will present the state-of-art research in OTFT-based sensors surged inthe <T2, 3>past 5–10 years (see Tables 1 and 2). The first part of this paper is devoted to abrief introduction on the fundamentals of organic semiconductors and OTFTs, then sensorsbased on OTFTs grouped by the active layer materials and functionalities are discussedin details. Older literatures that are of paramount importance in shaping this field are alsoreviewed. The paper is mainly focusing on the chemical and biological sensors. Therefore,most of the previously inspiring works on OTFT-based gas sensors, chemical vapor sensors,optical sensors, etc., will not be reviewed in this paper. Intended readers can refer to relatedliteratures.37,79,80

2. Mechanism of OTFTs

2.1 OFETs

OFETs have the source, drain, and gate electrodes, similar to their inorganic counterparts.33

Being absent of a depletion region, an OFET normally operates in an accumulation regionand the low off current in the device is dominated by the inherent conductivity of the activelayer. When a gate voltage is applied, the channel current flowing through the organicsemiconductor film can be modulated via field-effect doping process. The channel currentof the OFET is given by:29

IDS = W

LμCi

(Vg − Vt − VDS

2

)VDS VDS� Vg − Vt

IDS = W

2LμCi

(Vg − Vt

)2VDS > Vg − Vt,

in which IDS is the channel current between source and drain, W and L are the width andlength of the channel, respectively,μis the mobility of the carrier, Ci is the capacitance ofthe gate insulator, Vg is the gate voltage, VDS is the applied source-drain voltage, and Vt

is the threshold voltage of the device. Therefore, the channel current of an OFET-basedsensor could be primarily modulated by the two parameters in the above equations, namelythreshold voltage Vt and carrier mobility μ, which both play significant roles in sensingapplications. The threshold voltage of the device is normally changed by the doping effectupon the exposure to the target analytes81,82 while the carrier mobility is usually influencedby the analyte molecules diffused into the semiconductor grain boundaries.83,84

2.2 OECTs

An OECT normally operates in a liquid or solid electrolyte at low voltages. When a gatevoltage is applied, cations in the electrolyte are injected into the semiconducting layer todope or de-dope the semiconductor, resulting in a change of channel current. Similar tothe OFET, the channel current in an OECT at a low source-drain voltage is proportional

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Table 1Summary of OFET-based chemical and biological sensors

Active layer Sensor type AnalyteDetection

limit (range) References

P3HT Ammonia sensor Ammonia 10–100 ppM 93Humidity sensor Humidity 20%–80% 70pH sensor pH 3.4–5.6 94

pH 6.6–9.5 95pH 4–10 81pH 2–10 96pH 3.5–5.5 97pH 4–10 98

Ion sensor K+ 33 mM 95Na+, K+, Ca2+ 0.001% 81

Na+, K+ 0.5 mg/mL 94Na+ 0.001% 99

Glucose sensor Glucose 10−5 M 98DNA sensor DNA — 100

DNA — 76Cell sensor Cell — 99

Cell — 1011-pentanol sensor 1-Pentanol — 102Vanillin sensor Vanillin — 102Dopamine sensor Dopamine 1 ppM 103Biotin sensor Streptavidin 10 nM 104

Biotin 10−2 ppb 61Anesthetic sensor Halothane 1%–5% 61

PDTT Ion sensor SO42− 1 mM 50

Poly-DPOT Alcohols sensor 1-Hexanol 10–20 ppm 105Chemical vapor

sensorn-Heptane, Ethanol,

1-Butanol,1-Hexanol,

Acetone, and2-Propanol

— 106

Poly-DDT n-Heptane, Ethanol,1-Butanol,1-Hexanol,

Acetone, and2-Propanol

— 106

PEDOT:PSS Glucose sensor Glucose 1.1 mM 107PTAA pH sensor pH 2–10 108Pentacene Humidity Sensor Humidity 0%–75% 45

Humidity 0%–75% 110Alcohols Sensor 1-pentanol — 111Trimethylamine

sensorTrimethylamine 0.1 ppm 112

Ammonia sensor Ammonia 0.5 ppm 113pH sensors pH 4–10 114

(Continued on next page)

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Table 1Summary of OFET-based chemical and biological sensors (Continued)


limit (range) References

pH 2.5–7 115pH 2.5–7 116

DNA sensor DNA — 83,117,118,124DNA 1 μg/mL 119DNA 650 ng/mL 84DNA 10 nM 120DNA 0.1 nM 58DNA 50 pM 121,122DNA 10 pM 123

Antibody-antigensensor

BSA/antiBSA 500 nM 125

SDS sensor Sodium dodecylsulfate

1 μM 126

Biotin sensor Biotin 200 nM 126DDFTTF pH sensor pH 3–11 82

TNT sensor TNT 40 ppm 127Glucose sensor Glucose 10 ppm 82MPA sensor MPA 100 ppb 82Cysteine sensor Cysteine 100 ppb 82,128

Cysteine 10 ppb 127TNB sensor TNB 300 ppb 82

TNB 100 ppb 128DNA sensor DNA 1 nM 129

C12FTTF Humidity sensor Humidity 0%–75% 110C6TFT Humidity 0%–85% 110α6T Salt sensor KCl, NaCl, KBr 0.1 mM 77

pH sensor pH 2–7 77Glucose Sensor Glucose 1 mM 130Lactic acid sensor Lactic acid 0.3 mM 130

DHα6T Lactic acid — 130Chemical vapor

sensor1-Pentanol, — 131

Octanenitrile — 131DHα4T 1-Pentanol — 132

1-Pentanol 100 ppm 131F8T2 Antibody-antigen

sensorAvidin/BSA — 133

NTCDA Humidity Sensor Humidity — 134Humidity 0–83% 135

CuPc Hydrogen sulfidesensor

Hydrogen sulfide 100 ppm 136

Lactic acid sensor Lactic acid 10 μM 130Pyruvic acid sensor Pyruvic acid — 130

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Table 2Summary of OECT-based chemical and biological sensors


limit(range) References

Polypyrrole pH sensor pH 3–11 140NADH sensor NADH — 139Penicillin sensor Penicillin 0.1 mM 140

Polyaniline Humidity sensor Humidity 0%–100% 142Humidity 25%–80% 63

pH sensor pH 1–6 143Ion sensor Ru(NH3)6

3+/2+ — 143Fe(CN)6

3-/4− — 143Metal ions — 144

SO2 — 145Glucose sensor Glucose 2 mM 146

Glucose 10 mM 147,148,150Glucose 2 μM 151–153,155Glucose — 158

H2O2 Sensor H2O2 <1 ppm 156H2O2 <0.5 mM 157

Metabolic sensor Urea 10 mM 147,148,150,163Lipid 3 mg/mL 163

Hemoglobin 0.1 g% 163Triglycerides 10 mM 147,148,150

p-aminophenol 100 nM 161Alkaline

phosphatase<1 nM 162

PEDOT:PSS Humidity sensor Humidity 40%–80% 63Ions sensor Ca2+ 10−4 M 174

K+ — 175K+, Ca2+ 10−4 M 176

Ag+ 10−5 M 176Metal ions 10−6 M 177

CTAB 10−4 M 181Polyelectrolytes

SensorPolymer loadednanoparticles

10−5g/L 182

Glucose 0.1 mM 183Glucose 1 μM 86Glucose 1 μM 184,185,190Glucose 100 nM 186Glucose 5 nM 191Glucose — 189

Lactate sensor Lactate — 189Lactate <0.3 mM 193

H2O2 Sensor H2O2 <10−6 M 87H2O2 5 nM 191H2O2 ppm range 183

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Table 2Summary of OECT-based chemical and biological sensors (Continued)


limit(range) References

H2O2 1–100 mM 197,198Ethanol sensor Ethanol 0–30% 197,198DNA sensor DNA 10 pM 202Cell sensor KYSE30 cancer

cell lines— 195

HepG2 cancer cells — 88Caco-2 cell lines — 197,198

Antibody-antigensensor

Antigen 1 pg/mL 204

Bacteria sensor Bacteria 102 cfu/mL 206Dopamine sensor Dopamine 5 nM 207Liposome sensor liposome- based

nanoparticles10−7 mg/mL 209

PEDOT:PSS coatedcotton fiber

Saline sensor NaCl 10−4 M 180

PEDOT Glucose sensor Glucose 10 μM 187Antibody-antigen

sensorAntigen/ Antibody 10−10 g/mL 203

DNA sensor DNA 8 × 10−8

g/mL201

PEDOT:TOS Cell-based sensor Epithelial cell line — 199PEDOT:TOS–PEG Cell-based sensor HeLa cell — 200Poly(3-

methylthiophene)pH sensor pH

pH1–90–12

211212

Ion sensor IrCl62− 10−15 mol 211Polycarbazole Ion sensor Cu2+ 2.5 ×

10−6 M213

to the carrier density in the channel when an effective gate voltage is applied. So the channelcurrent is approximately given by:51,85,86

IDS = qμp0tW

LVP

(VP − V eff

g + VDS

2

)VDS |VDS | �

∣∣VP − V effg

∣∣VP = qp0t/Ci

V effg = VG + Voffset,

where Vp is the pinch-off voltage; Vgeff is the effective gate voltage; q is the electric charge;

μ is the hole mobility; p0 is the initial hole density in the active layer without any appliedgate voltage; t is the thickness of the active layer; W and L are the channel width andlength, respectively. Ci is the effective gate capacitance of the transistor, and Voffset is theoffset voltage determined by the potential drops at the gate/electrolyte interface and theelectrolyte/channel interface;87 accordingly, the response of the device is dominated by

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the diffusion process of ions in the active layer under gate voltage. To further elucidatethe underlying transient behavior of an OECT, the device can be described with twofundamental circuits, that is, electronic circuit and ionic circuit.85 Electronic transport isdetermined by the mobility and the density of the carriers in the channel, while ionictransport is closely related to the ionic diffusion in both the electrolyte and the active layer.It has been found that the geometric size of the OECT is essential to the transient propertiesof the device.88,89 Some small devices with short channel length and small distance betweengate and channel can show fast response time down to a few milliseconds,89 which willenable fast detection of analytes in sensing applications.

3. Organic Semiconductors in OFET-Based Sensors

3.1 Poly(3-hexylthiophene) (P3HT)

P3HT as a representative semiconducting polymer with high carrier mobility has been in-tensively investigated for the applications in organic electronic devices, including OFETs 90

and organic photovoltaic (OPV) cells.91 P3HT consists of head to tail (HT), head to head(HH), and tail to tail (TT) couplings. The percentage of HT coupling in polymer chains rep-resents the degree of regioregularity of P3HT, which is critical to the electronic propertiesof the material. High-performance OFETs based on high-regioregular P3HT were success-fully fabricated by solution process and showed the mobilities higher than 0.1 cm2/Vs andhigh on/off ratios (>105).92 The high mobility could be attributed to the high crystallinityand the ordered morphology of the P3HT films resulted from the well-defined moleculararchitecture. It is notable that P3HT is an excellent organic material for a wide rangeof sensing applications by virtue of its high carrier mobility, solution processability, andbiocompatibility.

3.1.1 P3HT in Chemical Sensors.3.1.1.1 pH and Ions Sensors. Due to its simple preparation process and commercial

availability, P3HT has been widely investigated as the active transducing material in thefabrication of pH sensors. Bartic et al.96 successfully demonstrated an OFET-based pHsensor in which the semiconducting layer is not directly in contact with aqueous solutions.A thin layer of silicon nitride was deposited on the surface of P3HT layer and exposed toan electrolyte as a proton sensitive surface. Stable pH-dependent responses were observedin the pH range from 2 to 10. The sensing mechanism was attributed to the changesof proton-induced potential drops at the solution/ silicon nitride interface. However, theworking voltage (10 V) of the device was rather high, making the device unsuitable for realphysiological sensing. This problem was overcome by the introduction of a thin tantalumoxide (Ta2O5) layer on the P3HT film98 (Fig. 4). Reproducible pH-dependent response wasachieved when small voltage (<1 V) was applied.

Scarpa et al.81 also demonstrated an ion sensitive organic field effect transistor(ISOFET) based on regioregular P3HT, which operated in aqueous electrolytes at lowvoltages. The device was stable in liquid solutions and exhibited detectable responses toNa+, K+, and Ca2+ ions with the concentrations down to 0.001%. They considered that thedevice is suitable for in vitro bio-sensing applications.

Sensitivity is a key parameter to sensors in practical applications. Ritjareonwattu et al.97

demonstrated a performance-enhanced P3HT-based OFETs for pH sensing, in which aLangmuir-Blodgett film of arachidic acid was incorporated. A thin layer of polymethyl-methacrylate (PMMA) was deposited on top of the P3HT film as the gate insulator. They

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Figure 4. (a) Schematic structure of the P3HT-based OFET sensor. (b) Performance of the P3HTOFET characterized at different pH values. Reproduced from Bartic et al.98 with permission fromAIP.

found that the sensitivity of the device to H+ ions was significantly improved in the pHrange from 3.5 to 5.5.

The selectivity of the P3HT OFETs can be improved by doing surface modification onthe gate dielectrics. Ji et al.95 successfully built up an OFET-based ion sensor that can beused to selectively detect H+ and K+ ions. A thin layer of high-k gate dielectric tantalumpentoxide (Ta2O5) was deposited on top of the P3HT layer and used as the H+ ion selectivelayer, which enabled pH detection in the range from 6.7 to 9.5. Then valinomycin ionophorewas loaded on the gate dielectric owning to its high selectivity to K+ ions. The device canbe used to selectively detect K+ ion down to 33 mM.

3.1.1.2 Other Chemical Sensors. P3HT OTFT-based chemical sensors have been usedfor effectively monitoring target analytes in the atmosphere. Jeong et al.93 proposed anammonia sensor based on a P3HT OFET platform with a detection limit down to 10 ppm.The channel current dropped rapidly when the device was exposed to ammonia gas (NH3).The device operates based on a direct interaction between the organic transducing layerand the analyte molecules without the introduction of additional sensing elements, offeringan effective way for NH3 detection in the environment. Moreover, analytes in liquid, suchas 1-pentanol and vanillin,102 can also be detected using P3HT OTFTs. The devices alwaysshowed decreased channel currents resulted from the diffusion of the analytes into the grainboundaries, which in turn reduced the carrier mobility by charge trapping effects in thechannel.

Since the physical properties of P3HT are sensitive to humidity, OFETs based onP3HT can be used as humidity sensors. Said et al.70 carefully investigated the effects ofhumidity on the device performance. The devices were prepared by coating regioregularP3HT on the substrate, followed by the deposition of poly(styrene sulfonic acid) layer asthe gate insulator. The transfer characteristics of the devices were strongly dependent on thehumidity level of the environment. They concluded that the hole mobility in P3HT decreasedwith the increase of environmental humidity due to the diffusion of water molecules intothe P3HT film.

3.1.2 P3HT in Biological Sensors.3.1.2.1 DNA Sensor. Fast-response, sensitive, and low-cost methods for DNA detec-

tion are in great demand for many applications, such as the clinical analysis. In 2009, we100

demonstrated a label-free DNA sensor based on P3HT OFETs. Thiol-modified probe DNAstrands were immobilized on the gold source/drain electrodes via chemisorption, followedby the hybridization of the target DNA strands. Finally, P3HT film was spin coated on the

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channel area to cover the DNA-modified electrodes. The channel currents were decreasedupon the immobilization and hybridization of DNA molecules on the electrodes, leading tothe successful differentiation of ssDNA and dsDNA. The sensing mechanism was attributedto the variation of the contact resistance of device, which was caused by the decreased workfunction of gold electrodes upon the introduction of DNA molecules.

Recently, Kergoat et al.76 fabricated DNA sensors based on the configuration of water-gated organic field-effect transistor with low operation voltages (below 1 V). DNA probeswere covalently grafted on the semiconducting layer via a simple coupling reaction betweenthe functional moieties of the semiconductor and the amino groups of DNA strands. A clearchange of the channel current was observed upon the immobilization or hybridization ofDNA molecules, which was resulted from the steric hindrance of the immobilized DNAchains that prohibit ion diffusion into the bulk of the semiconductor.

3.1.2.2 Cell Sensors. Due to its excellent biocompatibility, an OFET is potential forthe sensing applications integrated with cells. Scarpa et al.101 fabricated P3HT-based OFETfor biological sensing. They demonstrated that fibroblast cells could be successfully culturedon the surface of P3HT films treated with protein and oxygen plasma. They consideredthat a substantial hurdle for the realizations of low-cost and mass-produced sensors in lifescience were overcome, which opened new possibilities of biological sensing with organicelectronic devices.

3.1.2.3 Other Biological Sensors. Rapid and accurate determination of glucose levelin physiological fluids is of critical importance to the diagnosis of diabetes. Highly sensitiveglucose sensor based on OFET was realized by the incorporation of glucose oxidase (GOx)on the dielectric of the organic transistor.98 To facilitate the immobilization of GOx, thegate dielectric Ta2O5 surface was treated with cyanopropyltrichlorosilane. The drain currentshifted to a lower value upon a glucose addition, which could be attributed to the pH variationnear the device as a result of gluconic acid produced by the electro-oxidation of glucose.The device showed a good linear response to glucose concentration in the range of 10−5 to10−2 M.

A label-free dopamine sensor based on the EG-OFET architecture has also been demon-strated, as shown in Fig. 5.103 The device exhibited a dopamine-dependent performancedown to pM level, several orders of magnitude better than the amperometric sensors basedon the same principle. The sensing mechanism of the P3HT OFET-based dopamine sensoris attributed to the specific and covalent bonding between the boronic acid moiety anddopamine molecules occurred on the gate electrode surface, which in turn modulates boththe work function of electrode and the capacitance of double layer. Therefore, suitablesurface engineering in the EG-OFET is a promising approach for realizing highly sensitivebiosensors.

Biomaterials have been integrated into OFETs to achieve high-performance biosensors.Torsi’s group61 recently reported the functional biointerlayer organic field-effect transistors(FBI-OFETs) based on P3HT for effective biosensing applications. Functional biointerlay-ers, including a phospholipid (PL) bilayer, a purple membrane (PM) film, and a streptavidin(SA) protein layer, were coated between the gate dielectric and the organic semiconductorof three devices. They found that both the biological functionality and the electronic proper-ties were fully retained. The FBI-OFETs integrated with PL bilayer demonstrated a reliableresponse upon the exposure to anesthetic doses (1–5%) that unveiled the stimuli-inducedchanges in the lipid layer. The FBI-OFETs embedded with a SA layer demonstrated its supe-rior capability for the label-free biotin detection down to 10 parts-per-trillion concentrationlevel, comparable to the performance of the state-of-the-art fluorescent assay techniques.Later, the group104 introduced the EG-OFET-based sensors integrated with a biotinylated

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Figure 5. (a) Schematic picture of the EG-OFET device for label-free dopamine (DA) sensing. (b)Transfer characteristics of the device in different DA concentrations. (c) Potential drops at the organicsemiconductor/electrolyte and gate/electrolyte interfaces. Reproduced from Casalini et al.103 withpermission from Elsevier (Color figure available online).

phospholipid bilayer (Fig. 6). The bi-layer structure anchored on the surface of the organicsemiconducting layer was directly in contact with the aqueous solutions. They found thatthe device allowed rapid label-free detection of biotin in the range of 10 nM—1 μM.Owing to the specific bonding between biotinylated PLs and analytes, the biotin sensorbased on the EG-OFET platform also exhibited high selectivity. All of the above resultsindicate that P3HT OFETs can be easily integrated with biomaterials for various biologicalapplications.

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Figure 6. (a) Scheme of the fabrication of phospholipid modified EG-OFET. (b) Transfer character-istics of a BIOEGOFET with biotinylated PLs exposed to PBS (open symbols) and streptavidin (fullsymbols) solutions. (c) Transfer characteristics of a biotin-freeBIOEGOFET exposed to PBS (opencircles) and streptavidin solutions (full triangles). (d) Transfer characteristics of a BIOEGOFET withbiotin exposed to PBS (open circles) and bovine serum albumin (BSA) (full diamonds) solutions.Reproduced from Magliulo et al.104 with permission from Wiley (Color figure available online).

3.2 Other Semiconducting Polymers

Many other conjugated polymers have been used in OFET-based chemical and biologicalsensors. Maddalena et al.50 reported an ion-sensitive OFET sensor based on poly(4,4′-didecylbithiophene-co-2,5-thieno[2,3-b]thiophene) (PDTT). The modified sulfate bindingprotein (SBP) was covalently coupled to the device and used for the determination of sulfateion (SO4

2−) with a detection limit down to 1 mM.A PEDOT:PSS-based OFET was firstly reported by Liu et al. as an enzymatic sen-

sor.107 Glucose oxidase enzyme (GOx) was immobilized on the surface of PEDOT:PSSconducting polymer film. As the device was protected with a cellulose acetate membrane,PEDOT:PSS matrix and GOx could not be dissolved by the aqueous solution in the detectionof glucose. Under optimized testing conditions, the device was able to detect glucose downto millimolar concentration in a response time less than 20 s. The sensing mechanism was

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due to the de-doping process in PEDOT:PSS induced by the electro-oxidation reaction ofglucose.

Spijkman et al.108 demonstrated a dual-gate OFET-based pH sensor using semiconduc-tor polytriarylamine (PTAA) as the active layer. The semiconductor layer was protected bya hydrophobic barrier layer consists of a stack of polyisobutylmethacrylate (PIBMA) andthe Teflon derivative AF-1600. The channel current shifted to a lower value when the pHvalues varied from 2 to 10. They also found that the sensitivity of the device was stronglyinfluenced by the capacitance ratio between the top and bottom gates in the dual-gateOFETs.

Besides the crystallinity of the organic semiconductor, the inherent molecular struc-ture of the semiconductor also plays an important role on the performance of OFETs. Torsiet al.106 carefully examined the OFET sensors based on the alkyl and alkoxy-substitutedpolymers (Poly-DPOT and Poly-DDT). These two different transistors showed stable re-sponse to various analytes, such as alcohols, alkanes, and ketones. The sensing mechanismwas attributed to the weak interactions between the side chains of polymer and the functionalgroups of analytes. The OFET transistors with Poly-DDT exhibited responses influencedby the alkyl chain length of the analytes, while the devices with Poly-DPOT exhibitedresponses influenced by the dipole moment of the analytes.

3.3 Pentacene

Pentacene is a highly conjugated small molecule consisting of five linearly-fused benzenerings. This p-type organic semiconductor can generate excitons upon the absorption ofvisible light or ultra-violet, which makes this material unstable in air for slow oxidization.Pentacene is a representative and very important small molecule organic semiconductorowing to its outstanding properties, such as high mobility, low-cost, and ordered crystalstructure, etc. Pentacene has been intensively investigated for the applications in OTFTsand organic circuits. However, the pentacene-based devices have some shortcomings in theapplications as chemical and biological sensors since pentacene is not stable in aqueoussolutions. Therefore, the pentacene devices should be packaged when they are placed inelectrolytes or can only be operated in dry conditions.109

3.3.1 Pentacene in Chemical Sensors. Pentacene itself is sensitive to ambient humidity.Zhu et al.45 examined the response of a pentacene-based OFET to the variation of humidity.The results indicated that the sensitivity of the OFET sensor was strongly dependent onthe thickness of the semiconductor film. They also found that the device showed goodlinear responses when the humidity was above 30%. Later, Li et al.110 conducted similarstudies on humidity effects to the electrical performance of pentacene OFETs. They alsoobserved that the device performance degraded with increasing humidity level (Figs. 7aand 7b). The sensing mechanism was attributed to the changed hole mobility resultedfrom the interactions between water molecules and pentacene grain boundaries. Actuallythere are two competing factors involved in the above sensors.111 Carrier density may beincreased due to the doping of water molecules in pentacene films and charge mobility canbe decreased due to charge trapping and scattering at the grain boundaries.

Pentacene-based OFET was also a feasible platform for pH sensing. Loi et al.114

presented a flexible pH-sensitive OFET prepared on flexible Mylar sheet, which actedas both the substrate and the gate dielectric in the device configuration. One side of theMylar sheet was patterned with Au source/drain electrodes and covered by pentacene film bythermal evaporation, while the other side was exposed to the aqueous electrolyte containing

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Figure 7. (a) Schematic of a pentacene-based OFET. (b) Humidity dependence of the two types ofOTFT devices. The circles are for the OTFT with a 500 Å thick pentacene layer; the triangles are forthe OTFT with a 1000 Å thick pentacene layer. Reproduced from Zhu et al.45 with permission fromAIP. (c) Structure of a pH-sensitive OFET based on pentacene. (d) IDS responses of the device to thevariation of pH. Reproduced from Loi et al.114 with permission from AIP.

Ag/AgCl electrode as the gate (Fig. 7c). The device showed observable responses to thechanges in pH (Fig. 7d). The transfer curve of the device shifted to higher gate voltage withthe decrease of pH due to the change of effective gate voltage applied on the transistor.

Recently, Caboni et al.115,116 developed a pH sensor based on dual-gate OFET usingpentacene as the semiconducting layer. The device was fabricated on the flexible film (My-lar). They found that the device showed a pH sensitivity when the floating gate surface wasfunctionalized with thioaminic groups. Since the thioaminic groups protonate proportion-ally to the concentration of H+ ions in the electrolyte, the device was sensitive to the pHvalue ranging from 2.5 to 7.

3.3.2 Pentacene in Biological Sensors.3.3.2.1 DNA Sensors. DNA sensors have attracted much interest due to the great scien-

tific and economic importance. They have significant applications in gene expression mon-itoring, viral and bacterial identification, biowarfare and bioterrorism agents detecting, andclinical medicine. Pentacene OFETs have been successfully used as DNA sensors based ondifferent sensing mechanisms. Zhang et al.117 reported the label-free DNA sensor based onpentacene OFETs for the first time (Fig. 8a). The DNA molecules were immobilized on thesurface of semiconductor film via physical adsorption. The threshold voltage of the devicetested in dry condition shifted to a more positive value upon the immobilization of DNA.

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Figure 8. (a) Transfer curves of a DNA sensitive pentacene OFET. Reproduced from Zhang and Sub-ramanian117 with permission from Elsevier. (b) Schematic structure of the charge-modulated OFET(OCMFET). (c) Sensing mechanism of the OCMFET-based DNA sensor. (d) Transfer characteristicsof the OCMFET before and after DNA functionalization and hybridization. Reproduced from Laiet al.58 with permission from Wiley (Color figure available online).

More interestingly, the single-stranded DNA (ssDNA) and double-stranded DNA (dsDNA)could be successfully differentiated by recording the shifts values. The authors suggestedthat this phenomenon was induced by the electron-withdrawing nature of DNA. They alsofound that the pentacene film morphology was an important property that greatly influencedthe sensitivity of the device.83,117 Enhanced-performance was achieved by increasing theroughness of the pentacene film or reducing the film thickness. Later, Zhang et al.118 inte-grated the devices with microfluidic channels to realize the disposable sensors for portablediagnosis. The performance of the OFETs was characterized after they were dried. Thesensing mechanism of the sensors was also attributed to the electron-withdrawing nature ofDNA. DNA molecules adsorbed on the pentacene film can extract electrons from pentacene,which in turn change the doping level in the active layer.

Stoliar et al.84 demonstrated a label-free DNA sensor based on an OFET using the ultra-thin pentacene film. The DNA molecules were self-assembled on the surface of pentacenefilm without any external binding agents. The pinch-off voltage of the devices tested in dryconditions shifted across a wide range of DNA concentration, which could be ascribed tothe increased positive charges within the pentacene film induced by DNA adsorption. Thedetection limit was down to 650 ng mL−1.

It is inconvenient to firmly immobilize the bio-molecules onto the surface of pentacene.In order to solve this problem, Liu et al.124 introduced different electric biases during the

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DNA immobilization process to improve the immobilization efficiency of DNA moleculeson the surface of pentacene film. The highest immobilization efficiency was achieved atan optimized process condition (30 min immobilization at +50 V bias). However, thistechnique has not been proved to be useful in DNA sensors since the high bias voltage maychange the properties of the pentacene film.

The above reports paved the way for the development of label-free transducers for prac-tical applications. However, as pentacene is sensitive to external stimuli, such as moistureand ions, the direct immobilization of DNA molecules on the pentacene film may decreasethe stability and electrical properties of the active layer. Lai et al.58 fabricated a high-performance DNA sensor operable in aqueous solution with ultralow operation voltages,as shown in Figs. 8b–8d. This device allows for the detection of DNA without exposing theorganic semiconductor pentacene to the aqueous solutions containing target DNAs. Thedevice was able to selectively detect the target DNA molecules with a concentration as lowas 0.1 nM. Since the sensing layer is completely distinct from the active pentacene film, thesensing mechanism is reliant on the shift of the threshold voltage of the device, instead ofthe electronic property changes of pentacene. Therefore, the stability issue of the pentacenefilm was overcome in this proposed structure. More recently, the same group120 reported anovel DNA sensor based on the similar architecture. The device showed good stability insolution and could detect DNA molecules down to 10 nM.

3.3.2.2. Antibody-Antigen Sensors. Recently, Khan et al.125 demonstrated a label-freeantibody/antigen sensor based on a pentacene OFET. In the device fabrication, eitherperfluor-1-3-dimethyl cyclohexam (PFDMCH) or cyclized perfluoro polymer (CYTOP)was deposited on the pentacene film by plasma-enhanced chemical vapor deposition (PE-CVD) and/or spin-coating to passivate the active layer, which significantly improved thestability of the device in aqueous buffer media. Bovine serum albumin (BSA) was thenmodified on top of the passivation layer, which led to the highly selective detection ofantiBSA with a detection limit down to 500 nM. This work provides a bright futurefor the OFETs as highly sensitive, disposable, and label-free biosensors for physiologicalanalysis. In addition, other biological elements, including biotin126 could also be effectivelydetermined by the pentacene-based OFET sensors.

3.4 5,5′-bis-(7-dodecyl-9H-fluoren-2-yl)-2,2′-bithiophene (DDFTTF)

DDFTTF is a robust p-channel semiconductor for OFETs and exhibits superior stabilityin harsh conditions such as aqueous environments.127 Bao’s group exhaustively investi-gated the role of DDFTTF in the OFET-based sensors.82,128,129 Roberts et al.82 presented aDDFTTF-based OFET sensors exhibited remarkable stability in aqueous solutions (Fig. 9).To avoid the solution degradation that resulted from high operating voltage, thin cross-linked polymer poly(4-vinylphenol) (PVP) film was used as the gate dielectric layer, whichexhibited low operation voltages below 1 V. When source-drain and gate voltages were ap-plied, the channel current of the device was sensitive to the analytes in solutions, includingtrinitrobenzene (TNB), glucose, cysteine, and methylphosphonic acid (MPA) at the concen-trations as low as ppb. In addition, the device exhibited a reproducible response in the pHrange of 3–11. The sensing mechanism is believed to be analogous to that of the chemicalvapor sensing systems, in which analytes diffuse into the semiconductor–dielectric inter-face and change the charge mobility in the semiconducting layer. Although the selectivityof the device was not researched, the novel OFET configuration still holds great potentialfor bio-sensing applications. Recently, they128 replaced the silicone substrate with flexible

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Figure 9. (a) Schematic diagram showing a DDFTTF-based OFET in flow cell for aqueous-phasesensing applications. (b) Source-drain current (IDS) response of the DDFTTF OFET to pH. (c-f) IDS response to different analytes: (c) trinitrobenzene (TNB), (d) glucose, (e) cysteine, and (f)methylphosphonic acid (MPA). Reproduced from Roberts et al.82 with permission from PNAS, USA(Color figure available online).

polyethylene terephthalate (PET). The flexible transistor was highly sensitive to the analytesin the solution. Stable and reproducible responses were obtained when 100 ppb cysteine orTNB was added in tested solutions.

Roberts et al.127 also systematically studied the influence of molecular structure andfilm properties on the performance of the OFET sensors based on many small moleculesemiconductors. They found that the performance and the response time of a device inwater were greatly influenced by the thin-film morphology and the molecular structure ofthe organic semiconductor. The sensor was used to detect cysteine down to 10 ppb, anda greater response with a faster response time was achieved when the channel area wasdeposited with high grain density films. Moreover, the device with optimized thin filmmorphology is able to detect 2,4,6-trinitrotoluene (TNT) down to 40 ppm, expanding thesensing applications of OFETs.

In situ label-free DNA sensor based on the DDFTTF OFET was demonstrated by Khanet al.129 for the first time. Cross-linked poly(vinyl phenol) polymer was adopted as the gatedielectric, allowing the stable operation in aqueous buffer solutions at biases below 1 V.The surfaces of the DDFTTF film deposited by thermal evaporation were modified with athin maleic anhydride (MA) polymer layer employing the plasma-enhanced chemical vapordeposition (PE-CVD), which enabled the covalent immobilization of the peptide nucleicacid (PNA) strands that designed to detect the target DNA selectively. The transistorsensor could detect the target DNA down to 1nM concentration, being comparable to theremarkable sensitivity of the fluorescence based techniques.

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Figure 10. Humidity responses of (a) an OTFT with C6FTTF, (b) an OTFT with C12FTTF. Repro-duced from Li et al.110 with permission from AIP.

3.5 Other Small Molecule Organic Semiconductors

Many other small molecule organic semiconductors have been used in OFET-based biosen-sors. Li et al.110 reported humidity-dependent device characteristics when several differentp-channel semiconductor materials, including oligofluorene derivatives (C12FTTF andC6TFT), were used as the active layers of OFETs. As shown in Fig. 10, the sensitivities ofthe devices to humidity level ranged from 0% to approximately 80%. The correspondingcurrent response could be explained by the microscopic process that polar water moleculesdiffuse into the grain boundaries of the organic semiconductors and lead to the decreasedcharge carrier mobilities.

α-sexithiophene(α6T) has been considered to be stable in aqueous environment.127

Buth et al.77 reported the pH-dependent response of the electrolyte-gated OFET (EGOFET)devices using α6T as the organic semiconductor. Since the active layer was in direct contactwith electrolytes, the thin electrical double layer (EDL) at the electrolyte/organic semicon-ductor interface was regarded as the gate dielectric layer, enabling the modulation of theconductivity of the active layer with a low operation voltage even less than 1 V. The pHsensitive behavior of the devices was attributed to the surface charge on the channel, whichwas dependent on the pH value of the electrolyte. Moreover, the devices also exhibited adetectable response to a variety of salts, including KCl, NaCl, and KBr. Someya et al.130

integrated micro-fluid channels with the OFETs based on different organic semiconductors,including pentacene, α6T, and α,ω-dihexyl-α-hexathiophene (DHα6T). The devices exhib-ited stable performance under flowing water. Substantial current responses to the dilutedbiological elements, such as glucose and lactic acid, were successfully observed.

Torsi et al.131 fabricated a number of OFETs based on different alkyl-substituted α-thiophene oligomers and investigated their responses to chemical vapors. They found thatthe devices with adequate number of grain boundaries displayed a large response to 1-pentanol while grain-boundary-deficient devices almost showed no detectable responseto the analyte. These results indicated that the sensing responses to the alcohol vaporwere primarily resulted from the analyte-semiconductor interaction at grain boundaries.Someya et al.132 also explored the role of grain boundaries in the vapor sensing with α,ω-dihexylquarterthiophene (DHα4T) -based OFETs. The morphology of the DHα4T films

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was controlled by the substrate temperature and the channel length (L). The results alsosuggested that the device response was mainly due to the diffusion effect of vapor at thegrain boundaries of the semiconductor film.

Li et al.136 fabricated the copper phthalocyanine (CuPc) -based OFETs for hydro-gen sulfide (H2S) detection. They found that the sensing properties of the devices weredependent on the thickness of insulator layer. CuPc OTFT sensor with a 195 nm-thicksemiconducting layer exhibited the best sensing response to H2S with a detection limitdown to 100 ppm. In addition, highly sensitive lactic acid sensor and pyruvic acid sensorbased on the CuPc OFET have also been reported.130

Besides p-type organic semiconductors, some n-type ones can be used in OFET-based sensors. Torsi et al.135 reported that n-channel OFETs based on 1,4,5,8-naphthalene-tetracarboxylicdianhydride (NTCDA) could be used as humidity sensors due to the interac-tions between NTCDA and water molecules. The OFETs exhibited a better sensitivity and alarger detection range to relative humidity (RH) (from 0 to 83%) than that of NTCDA-basedchemiresistors.

4. Organic Semiconductors in OECT Sensors

4.1 Polypyrrole (PPy)

Polypyrrole (PPy) is a chemical compound formed from a number of connected pyrrolering structures with high conductivity.137 Due to its excellent environmental stability andoutstanding biocompatibility, polypyrroles is widely investigated for the bio-applications.The so-called organic electrochemical transistor (OECT) using an organic material as theactive component of a transistor was first established by Wrighton et al.138 in the early1980s. The first OECT had three Au microelectrodes coated with polypyrrole films, inwhich the source and the drain electrodes were interconnected via the conducting polymer(Fig. 11). The devices were characterized in the aqueous electro-lyte of CH3CN/0.1 M[n-Bu4N]ClO4. The sensing mechanism relies on the de-doping effect of the polypyrrole,resulting in the changes of the electrical conductivity. When a negative gate voltage wasapplied, the device was in “OFF” state. While when the gate voltage was positive, thedevice shifted to “ON” state due the increment of channel current induced by the oxidationof polypyrrole film. This pioneering work initiated the applications of OECTs as chemicaland biological sensors.

Then Matsue et al.139 developed a NADH sensor based on the OECTs devices. TheOECTs were deposited by a thin layer of pyrrole-N-methylpyrrole copolymer containingdiaphorase. The device was at “ON” state when the polymer was in the oxidized form, whilechanged to the “OFF” state when NADH was added. The sensitivity toward NADH wasattributed to the reduction of the polymer catalyzed by the immobilized diaphorase enzyme.Later, Nishizawa et al.140 in the same group also demonstrated a penicillin sensor by usingpH-sensitive OECTs. The OECT sensors were fabricated by the deposition of polypyrroleand penicillinase membrane on the inter-digitated microarray electrodes. In this report, theconductivity of the polymer could be reversely changed within the pH region from 3 to11. The enzyme reaction induced by the penicillin addition could change the pH valuenear the polymer layer, leading to the modulated output current. The device started toshow an obvious response when 0.1mM penicillin was added and reached a saturated pointof 8mM.

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Figure 11. Output curves of the poly(pyrrole) OECT in the aqueous electrolyte of CH3CN/0.1 M[n-Bu4N]ClO4. Inset: the schematic diagram of the device. Reproduced from White et al.138 withpermission from ACS.

4.2 Polyaniline (PANI)

Polyaniline captured the intensive attention of the scientific community since the early1980s, particularly due to its high electrical conductivity, good stability in electrolytes,and good biocompatibility.141 As shown in Fig. 12, polyaniline has three idealized oxi-dation states, including leucoemeraldine, emeraldine, and (per)nigraniline, among whichthe emeraldine state was regarded as the most important one of polyaniline due to itshigh conductivity at room temperature. Another advantage of polyaniline is the solutionprocessability. Polyaniline has been successfully used in chemical and biological sensors,electrochromic coatings and transparent electronic conductors, etc.

Figure 12. Three idealized oxidation states of polyaniline.

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4.2.1 Polyaniline in Chemical Sensors.4.2.1.1 Humidity Sensors. OECTs based on polyaniline have been used as humidity

sensors. Followed the pioneering work in OECTs,138 Wrighton’s group142 used the “solid-state” polyaniline-based transistor as a humidity sensor. Polyaniline was used as the activelayer and the poly(vinyl alcohol)/phosphoric acid acted as the solid-state electrolyte. Thedevice demonstrated a reversible response to the humidity level from 0% to 100% at roomtemperature. The current response of the devices was the combined consequences of themodulation of H+ ionic conductivity in solid electrolyte and the faradic leakage currentacross the gate.

4.2.1.2 Ions Sensors. Direct reactions between the analytes and the active layer willintroduce significant changes in the polymer channel conductance. The electrical conduc-tivity of polyaniline can be changed for several orders of magnitudes before and afterchemical reactions. Based on this sensing mechanism, Wrighton’s group143 reported thefirst OECT-based ion sensors. Polyaniline was electrodeposited onto an Au microelec-trode array. Dramatic changes in the conductivities of polyaniline films were observedupon the exposure to the redox reagents in NaHSO4 solution. They demonstrated that thepolyaniline-based OECT could be switched to “ON” and “OFF” states by adding Fe(CN)6

3−

and Ru(NH3)6, respectively, which paved the way for the rapid detection of redox reagents.Furthermore, the device was proved to be sensitive to the pH value ranged from 1 to 6 insolutions.

Based on the same concept, Dabke et al.144 reported a highly sensitive poly-anilineOECTs for the detection of metal ions, including K+, Na+, Ba2+, Sr2+, and Ca2+. Thedevice was deposited with a polyaniline film containing 18-crown-6 ethers, which actedas a specific ion-binding cavity for K+. The ion sensor was capable of detecting K+ downto 10−7 M. This ion-sensitive response was triggered by the conformation changes inthe presence of K+. Polyaniline-based OECTs also showed response to SO2 dissolved inaqueous solutions. Gaponik et al.145 demonstrated that the drain current of the polyanilineOECTs could be effectively modulated by oxidation reaction of the dissolved SO2.

4.2.2 Polyaniline in Biosensors.4.2.2.1 Glucose and Hydrogen Peroxide Sensors. Contractor et al.146 first demon-

strated the proof-of-principle work of glucose biosensor using polyaniline as the activelayer. This OECT sensor was based on the three-layer configuration, in which a polyani-line layer was coated on the platinum source and drain electrodes and then another thinpolyaniline-enzyme film was deposited on the top. The devices showed a stable and re-producible linear response down to 10 mM glucose. As the conductivity of polyanilinewas inherently dependent on the pH value of the electrolyte, the sensing mechanism wasattributed to the conductivity changes induced by the pH variations near the active layer,which was caused by the glucose oxidation catalyzed by the enzyme glucose oxidase.

Then Contractor et al.147 presented microsensor arrays of polyaniline-based OECTsfor the detection of multi-analytes. Glucose oxidase immobilization was accomplished bythe electro-polymerization process. The array showed excellent linear response toward theglucose range of 10–50 mM (Fig. 13a), which was represented by the change in the con-ductivity of the polymer film due to enzyme-catalyzed reaction. Later, the same group148

developed a new way for the fabrication of micro-OECTs using polycarbonate membrane.The effect of channel length on the device was carefully investigated. They also demon-strated that the morphology of the electropolymerized polyaniline can significantly affectthe performance of the OECTs. The device using more disordered polyaniline microtubulesexhibited a larger modulation of conductance and a higher sensitivity. Based on this work,

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Figure 13. (a) Conductance change of a polyaniline film vs. the concentration of glucose. Reproducedfrom Sangodkar et al.147 with permission from the American Chemical Society. (b) Conductancechange of the PSS-PANI microtubule-based sensor as a function of glucose concentration. Reproducedfrom Kanungo et al.150 with permission from the American Chemical Society. (c) Typical plots ofthe corresponding switch response to different glucose concentrations. Reproduced from Bartlett andBirkin152 with permission from the American Chemical Society. (d) The current response to 10 mMglucose for the OECT switch using the GOx-(OsCOOH)6 complex. Reproduced from Battagliniet al.158 with permission from the American Chemical Society.

they built the enhanced microtubular biosensors and arrays based on the polyaniline mi-crotubes.149 Compared with the macro sized sensor, the response of this type microtubularsensor for glucose detection is much higher (more than 1000 times). The detection limitfor glucose sensing was down to 10 mM. This technique demonstrated the possibility forthe development of disposable, inexpensive sensors arrays. In another work reported by thesame group,150 microtubular sensors based on poly(styrene sulfonate)-polyaniline (PSS-PANI) composites were fabricated and exhibited an improved responsibility for glucosedetection due to the improved enzyme immobilization in the channel during polymerization,as shown in Fig. 13b.

Bartlett et al.151–153,155,157,158 extensively studied various polyaniline-based OECTs forhydrogen peroxide and glucose sensing. Bartlett’s group151 demonstrated that the enzymeswitch fabricated on the screen-printed carbon microband electrodes was sensitive to glu-cose for the first time. The switch was constructed by the encapsulation of glucose oxidasein poly(1,2-diaminobenzene) films deposited on the top of polyaniline layers. Acting as

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a redox mediator, tetrathiafulvalene (TTF) was used to shuttle the electrons between theenzyme and the conjugated polymer. Upon the addition of glucose, the drain current wasincreased rapidly to a stable value, as shown in Fig. 13c.

Later, they152 systematically studied the glucose sensing mechanism of the micro-electrochemical enzyme transistor. The conductance change was caused by the directoxidation or reduction of the polymer, rather than the pH value change. The detection limitto glucose could be extended down to 2 μM when the device structure was optimizedby changing the active layer thickness and the channel length. They also investigated thestability of the device under the exposure to hydrogen peroxide. The results indicatedthat hydrogen peroxide had little effect on the conductivity of polyaniline. Furthermore,they153 reported the optimized microelectrochemical enzyme transistor that could be usedto detect the low glucose concentrations in the range of 2–100 μM. Compared with theresults obtained from conventional amperometric enzyme electrodes, the detection limitwas improved by a factor of 40.

One significant problem of polyaniline is the loss of conductivity when pH isabove 5, making this material undesirable for many applications.154 The problem wasfirst solved by Bartlett’s group.155 They prepared composite films of polyaniline withpoly(styrenesu1fonate) as the counter-ion and extended the working range of the polyani-line film to the neutral conditions successfully. Excellent conductivity and stability of theOECTs were achieved in the solutions with pH of 7.

Another strategy for improving the conductivity of polyaniline at neutral pH wasproposed by Battaglini et al.156 They demonstrated that the modification of PANI withpropane sultone could dramatically improve the conductivity of the polymer near neutralpH, without sacrificing its superior properties displayed at lower pH. The OECTs basedon this modified polyaniline material could detect hydrogen peroxide down to 1 ppm atneutral pH electrolyte with fast response when horseradish peroxidase or microperoxidasewas combined with polyaniline.

In 1998, Bartlett et al.157 modified horseradish peroxidase on OECTs and demonstratedthe possibility of detecting hydrogen peroxide at pH 5 with a detection limit of 0.5 mM.The device was slowly switched from “ON” to “OFF” states with the injection of hydrogenperoxide in a solution. The conductivity decrease of polyaniline film in the presence ofhydrogen peroxide was attributed to the oxidation of polyaniline active layer by directelectrochemical communication between the enzyme and the conducting polymer.

The above-mentioned polyaniline-based OECTs were normally involved with the ad-dition of redox mediator, which may impede the applications for in vivo detection andmeasurement. To solve the problem, Bartlett et al.158 covalently attached the redox me-diator to glucose oxidase directly. The redox mediator used here was the pyridine-basedosmium complexes, with a carboxylate or aldehyde group attached. The osmium-modifiedglucose oxidase GOx(OsCOOH)6.7 proved to be effective in mediating the redox reaction,as shown in Fig. 13d.

4.2.2.2 Metabolic Sensors. NADH plays a significant role in the chemical processinvolved with energy generation. People use NADH supplements as medicine, for the treat-ment of high bloodpressure, high cholesterol, as well as for the Alzheimer diseases.159

Therefore, rapid determination of the NADH level in the human body is of critical impor-tance. Bartlett et al.160 fabricated the first highly sensitive NADH sensors by using OECTsbased on polyaniline-poly(vinylsulfonate) composite active layers. The sensors showed agood linear response to NADH in the range of 1 to 4 mM in neutral aqueous solutions, whichwas attributed to the oxidation of NADH directly catalyzed by polyaniline (Fig. 14a). Highersensitivity was expected when the device was coupled with appropriate redox enzyme.

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Figure 14. (a) The switching rate of the PANI-PVS based devices as a function of the concentrationof NADH. Reproduced from Bartlett et al.160 with permission from RSC. (b) IDS vs. time in solutionswith different concentrations of paminophenol. 1–7: blank, 100, 150, 200, 300, 450 and 900 nM.Reproduced from Astier and Bartlett161 with permission from Elsevier. (c) The schematic of the twocell arrangement used to measure alkaline phosphatase concentration. (d) IDS vs. time for differentconcentrations (1, 2, 3 and 4 nM) of alkaline phosphatase, VDS = 30 mV. Reproduced from Astierand Bartlett162 with permission from Elsevier. (e) Simultaneous analysis of glucose, urea, and lipid inthe mixed solution using a sensor array. Reproduced from Sangodkar et al.147 with permission fromthe American Chemical Society (Color figure available online).

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P-aminophenol detection is the basis for electrochemical immunoassays. Bartlett’sgroup161 built a novel two-compartment-cell system to realize the detection of p-aminophenol, as shown in Fig. 14b. The polyaniline-based OECT was separated fromthe biological buffer solution, with a salt bridge integrated between the two cells. Thesystem provided the possibility for the p-aminophenol detection in nanomolar range, inde-pendent of the pH value. Based on the same system, they162 successfully detected alkalinephosphatase down to a very low concentration (<1 nM) as well, as shown in Figs. 14c and14d.

Contractor et al.163 reported that polyaniline-based OECTs with an appropriate en-zyme incorporated in the conducting polymer matrix could be used to detect urea, lipid,and hemoglobin. Since the conductivity of the polyaniline was pH-depended, the currentmodulation of the sensors could be ascribed to the pH changes near polyaniline. Polyaniline-based microsensor array was also investigated.146–148,150 Three types of enzyme, includingglucose oxidase, urease, and lipase, were separately immobilized on the specific electrodes.They demonstrated that the sensor array was able to detect the three analytes simultane-ously from a mixed solution, as shown in Fig. 14e. Following their previous work, theyfabricated the OECTs based on polyaniline microtubules, which could be used to detecturea and triglycerides down to 10 mM. Higher sensitivity could be achieved when geomet-ric parameters were optimized and more disordered polyaniline microtubules were used.Later, they150 reported that the microtubules sensor arrays using poly(styrene sulfonate)-polyaniline composites also showed the capability for urea and triglycerides sensing withdetection limits down to several mM.

4.3 Poly(3,4-ethylenedi oxythiophene) (PEDOT)

One of the most commonly studied conducting polymers is p-doped poly(3,4-ethylenedioxythiophene) (PEDOT). This kind conducting polymer exhibits high conductivityand good stability in a wide pH range and is commercially available,164 all ofwhich make it the prime candidate for numerous applications, including sensors,128,165

OTFTs,166,167memories,168,169 displays,170,171 etc. As the pristine PEDOT is insoluble, high-molecular-weight counter ions are always added to achieve the solution-processable dis-persions of PEDOT. The conductivities of PEDOT films ranged from 1 to 1000 S/cm,depending on the used counter ions and the doping level. The most popularly used one ispoly(styrene-sulfonate) (PSS) -doped PEDOT, that is, PEDOT:PSS. PEDOT doped withanother counter ion p-toluenesulfonate (TOS) also showed excellent properties for chemicaland biological sensing applications.172

4.3.1 PEDOT in Chemical Sensors.4.3.1.1 Humidity Sensor. Nilsson et al.63 reported flexible OECT-based humidity sen-

sors that can be prepared by printing the commercially available PEDOT:PSS on fine papersor thin polyester foils. The PEDOT:PSS films were patterned as the source, drain, and gateelectrodes as well as the active channel layer. Then a Nafion layer was used as the solid-stateelectrolyte to cover the channel area and the gate electrode. Nafion is a proton conductingmaterial popularly used in sensors.173 Since the conductivity of Nafion was strictly depen-dent on the humidity level of the ambient environment, the variation in humidity induced achange of Nafion conductivity up to several orders of magnitudes, which in turn resulted ina change of the doping level in PEDOT:PSS under a gate voltage. They demonstrated thatthe channel current of the OECT experienced an exponential decrease (approximately twoorders of magnitudes) when the ambient humidity increased from 40% to 80%.

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4.3.1.2 Ions Sensor. Rapid tracing or determination of ion concentrations is impor-tant in many applications, such as clinical analysis and environmental monitoring. Highlysensitive and selective OECT-based ion sensors represent a novel platform for the devel-opment of disposable, inexpensive and well-performed devices. A highly Ca2+-selectivePEDOT:PSS sensor based on an OECT has been firstly reported by Berggren et al.174 Uponthe application of small gate voltage (Vg = 0.15V), the OECT could be able to detectCa2+ down to 10−4 M. The high selectivity toward Ca2+ detection could be attributed tothe ionophore-based solvent polymeric membrane coated on the top of the PEDOT:PSSchannel, which was composed of 2-nitrophenyl octyl ether, poly(vinyl chloride), potassiumtetrakis(4-chlorophenyl)borate, and N,N,N,N′-tetracyclohexyl-3-oxapentanediamide.

Bernards et al.175 integrated a PEDOT:PSS-based OECT with biological recognitionelements to distinguish the monovalent and divalent cations. The biological elements,bilayer lipid membrane (BLM) coupled with ionophore Gramicidin, can form the ion-channels that are only permeable to monovalent cations rather than polyvalent cations oranions. For the OECT with KCl solution, the conductivity of the PEDOT:PSS channelwas dramatically changed when a small gate voltage was applied. However, the channelconductivity cannot be changed by gate voltage when CaCl2 solution was used in the device.The results indicated that the integration of ionophore proteins into the BLMs led to thehigh selectivity for monovalent ion identification. This concept provides a feasible way forsensing different ions with OECTs.

Mousavi et al.176 incorporated OECT sensors with K+, Ca2+, and Ag+-selectiveionophore-based solvent polymeric membrane and realized highly selective and sensi-tive K+, Ca2+, and Ag+ sensors. The detection limits toward K+, Ag+, and Ca2+were 10−4

M, 10−5 M, and 10−4 M, respectively. More interestingly, PEDOT:PSS active layer itselfcould serve as Ag+-selective membrane for the detection of Ag+ due to the spontaneousoxidization of PEDOT:PSS in Ag+ solution.

Later, our group177 systematically elucidated the ion-sensitive mechanism ofPEDOT:PSS-based OECTs (Fig. 15a). The device was characterized in several ion so-lutions, including H+, K+, Na+, Ca2+, and Al3+ ions and could be used to detect metalions down to 10−5 M. We demonstrated that the transfer curves shifted to a lower gatevoltage when the cation concentration of the electrolyte was increased. A universal curvewas obtained by shifting the curves horizontally. The relationship between the gate voltageshift and the ion concentration could be explained by the Nernstian equation (Fig. 15b) dueto the following reaction:

n(PEDOT+ : PSS−) + Mn+ + ne− Reduction−−−−−→←−−−−−Oxidation

nPEDOT0 + Mn+ : nPSS−.

In addition, different gate electrodes, including Ag/AgCl, Pt, and Au, were employedto study the transfer characteristics of the OECTs. Interestingly, we found that the ion-sensitive properties of the device were largely depended on the gate electrodes. The devicewith Pt or Au gate electrodes presented bigger gate voltage shifts to the change of ionconcentrations than the case when a Ag/AgCl electrode was used. The different voltageshifts could be explained by the distinct interfacial properties between the gate electrodeand the electrolyte.

Tarabella et al.178 demonstrated a similar result when different gate electrodeswere used in the PEDOT:PSS-based OECTs. They reported that the device with an Aggate showed a larger modulation of channel current than that of the device with a Pt gateelectrode (Fig. 16a). The distinct responses were attributed to the different properties of the

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Figure 15. (a) Transfer characteristics of a PEDOT:PSS-based OECT measured in KCl solutionswith different concentrations. Inset: schematic of the device. (b) OECT response |�VG| as functionsof the concentrations of metal cations in the solutions of KCl, Ca(NO3)2 and Al2(SO4)3. Reproducedfrom Lin et al.177 with permission from the American Chemical Society (Color figure availableonline).

gate electrodes in electrolytes. The Ag gate electrode reacts with anions and has a Faradaiccurrent under a gate voltage. Therefore, the voltage dropped at the electrolyte/gate interfaceis constant in the device. However, the Pt gate electrode is very stable in electrolytes, sothe voltage drop on the double layer of gate electrode is relatively high under a gate volt-age, which leads to reduced voltage applied on the electrolyte/semiconductor interface andlower modulation of channel current IDS (Fig. 16b). Yaghmazadeh et al.179 simulated theeffect of device architecture and materials parameters on the performance of OECT-basedsensors. To achieve a higher sensitivity, a gate electrode with the area smaller than thatof the channel region was recommended. They found that the sensitivity could also beimproved by maximizing the ratio between channel width and length and increasing themobility of the charge carrier and the capacitance per unit area of the active layer (Figs.16c and 16d)

More recently, Tarabella et al.180 presented an elegant work by testing saline in sweat.The OECT sensors were based on a single natural cotton fiber functionalized with PE-DOT:PSS and a simple Ag wire as the gate electrode. The devices were competent for theeffective detection of NaCl concentration in the range of 10−1–10−4 M. (Figs. 17a and 17b)The sensing mechanism could be attributed to the redox reaction between ions in solutionand the Ag gate electrode. The devices had the working voltages as low as 0.2 V, implyingthe great potential for sensing salts in the physiological conditions. Tarabella et al.181 theninvestigated the role of the specific electrolyte on OECT performance. Planar OECTs basedon PEDOT:PSS were fabricated by a lithographic patterning process, using the micelle-forming cationic surfactant cetyltrimethylammonium bromide (CTAB) as the electrolyte.When the CTAB concentration was lower than its critical micellar concentration (CMC,approximately 10−3 M at 298 K), the channel current modulation by gate voltage was verylow. Only when the CTAB concentration was above the CMC, the channel current wasdramatically changed with gate voltage (Figs. 17c and 17d). These results clearly indicatedthat positively charged CTA+ micelles (existed above CMC) dedoped the PEDOT:PSS layer

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Figure 16. (a) OECT response �I / I0 as functions of gate voltage Vg measured in 0.1 M NaClaqueous solution, using Ag (circles) and Pt (squares) gate electrodes. (b) Potential drops between thegate/electrolyte and the electrolyte /channel interfaces for Ag and Pt gate electrodes. Reproduced fromTarabella et al.178 with permission from AIP. (c) Schematic of net charge distribution in the electrolytewhen a positive gate voltage was applied. (d) Simulated sensitivity of OECT-based enzymatic sensorsvs. the ratio between channel area and gate area (Ach/Ag). Reproduced from Yaghmazadeh et al.179

with permission from Wiley.

more effectively than CTA+ dissociated ions (existed below CMC). This work extendedthe sensing applications of OECTs by demonstrating the detection of micelle formation.

4.3.1.3 Polyelectrolyte Sensor. Iannotta et al.182 reported the application of OECTs inthe determination of polyelectrolyte polymeric shells for the first time. The OECTs basedon PEDOT:PSS were fabricated on glass slides by using an automatic syringe and micro-positioning system. A PDMS vessel was then attached to the glass substrate to create achamber. OECT sensors were used to detect poly (acrylic acid) (PAA) and poly (allylaminehydrochloride) (PAH), which are typically used to functionalize the nanoparticles for drugdelivery. They found that the modulation of the channel current was reduced upon theaddition of negatively charged PAA, while no visible effect on the modulation of channelcurrent was observed upon the addition of positively charged PAH. The distinct responsestoward PAA and PAH could be explained by the interaction of the polymeric chain with theions presented in the CaCl2 solutions. They further demonstrated that gold nanoparticles(NPs) modified with PAA could also reduce the modulation of IDS. Interestingly, thedevices showed a lower detection limit to the gold nanoparticles modified with PAA. Thisresult was attributed to the fact that more than one polymeric chain was attached to single

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Figure 17. (a) Schematics of the OECT based on cotton functionalized with PEDOT. (b) Normalizedtransistor response |(I - I0)/I0| as a function of the salt concentration with different gate voltages. Theinset shows the response in the physiological concentration range of saline in sweat (10–100 mM).Reproduced from Tarabella et al.180 with permission from RSC. (c) Schematic of the OECT deviceindicating the electrical contacts. (d) Channel current response |(I - I0)/I0| as a function of gate voltageVg characterized in different CTAB concentrations, Vd = −0.4 V. Reproduced from Tarabella et al.181

with permission from RSC (Color figure available online).

NPs, which resulted in a higher response than that of the polymer alone. This idea willprovide us the possibilities to discriminate different polymeric shells of the NPs, and morepractically, to determine the quantity of the polymer materials already loaded on the surfaceof nanoparticles.

4.3.2 PEDOT in Biological Sensors.4.3.2.1 Glucose Sensors. Malliaras’s group183–189 investigated the OECTs for glucose

sensing from 2004, which paved the way for a series of applications of the devices. Zhuet al.183 demonstrated the application of OECTs for glucose sensing in aqueous solutions.PEDOT:PSS layer was patterned as the active channel, source, and drain electrodes, and Ptwire was used as the top gate electrode. A dramatic change of the channel current IDS wasobserved upon the addition of glucose in the solutions containing glucose oxidase (GOx),which was attributed to the catalyzed oxidation reaction of glucose, producing H2O2 andD-glucono-1,5-lactone. The by-product H2O2 was oxidized on the Pt gate electrode andinduced Faradic current on the gate. To maintain a charge balance, PEDOT was de-dopedto the neutral state, which consequently reduced the conductivity of the active layer (Fig.18a). Then Macaya et al.184 utilized the PEDOT:PSS-based OECTs for glucose sensingdown to micromolar concentration, which were sensitive enough for the determination ofglucose level in human saliva. The performance of the devices was also influenced by thegate voltage. They suggested that the sensing technology could be used for non invasive

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Figure 18. (a) Sensing response as a function of time in PBS solution. Inset: relative changes of IDS

vs. Vg for two solutions. Reproduced from Zhu et al.183 with permission from RSC. (b) The offsetvoltage of an OECT vs. glucose concentration. Reproduced from Bernards et al.86 with permissionfrom RSC. (c) Sensing response of an OECT as a function of glucose concentration. Inset: Schematicof the OECT device. Reproduced from Yang et al.186 with permission from RSC (d) Photograph of theOECT array integrated with a surface-directed microfluidic system. Reproduced from Yang et al.189

with permission from RSC (Color figure available online).

glucose detection, which could be an important alternative technology for blood glucosesensing.

In order to elucidate the mechanism of sensing, Bernards et al.86 investigated theunderlying device physics of the glucose sensor based on OECTs shown in Fig. 18b.The PEDOT:PSS-based OECT with Pt gate electrode was characterized in PBS solutioncontaining GOx. Due to the de-doping of the organic semiconductor, IDS decreased withthe increasing gate voltage. The channel current was dramatically decreased with theincreased concentration of glucose. The sensing mechanism could be described by the

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Nernst equation:

V effg = Vg + (1 + γ )

kT

2eln[H2O2] + const,

where Vgeff is the effective gate voltage applied on the OECT; γ is the ratio between

the capacitances of the channel and the gate. [H2O2] is the concentration of hydrogenperoxide. H2O2 generated by glucose was oxidized at the Pt gate electrode, resulting in theFaradic current flow near the Pt-gate. The potential drop at the interface of electrolyte/Pt-gate decreased with the increase of glucose concentration according to the equation. So theeffective gate voltage applied on the PEDOT:PSS channel was increased with the increaseof glucose concentration. It is interesting to find that the transfer curves for different glucoseconcentrations can be shifted horizontally to a universal curve. The improved understandingof the sensing mechanism opened the door for the rational design of OECT-based enzymaticsensors.

New glucose sensor based on an all-plastic OECT was successfully realized by Shimet al.185 The device with whole PEDOT:PSS showed adequate response to glucose additionin the range of 1 μM–200 μM, which was potentially useful in the detection of glucose levelsin human saliva. More importantly, the normalized response was dramatically increasedwhen the electrolyte was preloaded with ferrocene, which plays a critical role in facilitatingthe electron transfer between the redox enzyme and the PEDOT:PSS gate electrode.

Cicoira et al.87 also studied the role of geometric parameters based on the planarOECT with PEDOT:PSS as the active layer. Patterned hydrophobic SAM perfluoro octyltrichlorosilane (FOTS) were used to define electrolyte regions. Interestingly, they foundthat the device with smaller gate electrode demonstrated higher sensitivity and lowerbackground signal toward the H2O2 detection. They associated this phenomenon with thepotential drops at the electrolyte/Pt gate interface. The modulation of the channel currentwas improved in the presence of H2O2 due to the Faradaic contribution. But the detectionlimits (both minimum and maximum) to the analyte were insensitive to the channel/gateareas. This report provided a feasible way to enhance sensing performance by optimizingthe geometric size of OECTs.

Yang et al.186 integrated the PEDOT:PSS-based OECT with a room temperature ionicliquid as the immobilization medium for enzyme and mediator. The (tridecafluoro-1,1,2,2-tetrahydrooctyl) trichlorosilane (FOTS) monolayer was then deposited on the surface ofthe device. The uncovered PEDOT:PSS area served as the hydrophilic “virtual wells”to effectively confine the electrolyte. A small amount of room temperature ionic liquidsincluding enzyme GOx and the mediator ferrocene [bis(n5-cyclopentandienyl)iron] (Fc)were first placed onto the device, followed by the addition of glucose PBS solution (Fig.18c). The device could be able to detect the glucose level ranged from 10−7to 10−1 M.

Kim et al.187 developed a new OECT-based glucose sensor prepared by vapor phasepolymerization of 3,4-ethylenedioxythiophene (EDOT). For the first time, they coateda fluorescent layer of 6,6′-bis(4-(4-methylpiperazin-1-yl)styryl)-3,3′-bipyridazine (DPP-BPDZ) on the top of the PEDOT layer, which could also be used as the optical sensinglayer for the glucose detection. The devices could detect glucose concentration down to10 μM. The combination of electrochemical and optical approaches used for glucosedetection makes the device a more versatile platform for the sensing application.

Integrating OECTs into microfluidic systems provides a great deal of opportunities formicrofluidic-based chemical and biological sensors, which is a step closer to real appli-cations. Mabeck et al.188 reported an OECT device integrated with a PDMS microfluidic

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channel. The microfluidic channel not only afforded the controlled flow of small samplevolumes in the transistor, but also acted as the gate electrode for the devices. Under the gatevoltage of +1 V, the channel current IDS could be modulated for two orders of magnitudewhen 10 mM tris(hydroxymethyl)aminomethane hydrochloride (Tris-Cl) buffer solutionwas injected through the microfluidic channel at 0.1μL/min.

Then Yang et al.189 demonstrated that the surface-directed microfluidic system inte-grated with OECTs was capable of detecting multiple analytes with high sensitivity. ThePEDOT:PSS channels were modified with different enzymes. The sample solution wasfirst placed on the top reservoir and began to flow along the hydrophilic channels sponta-neously, finally reached the measurement reservoir (Fig. 18d). Glucose and lactate weresuccessfully identified and differentiated from the sample solutions. Based on this strategy,it would be very convenient to extend the system for the detection of multi-type analytessimultaneously.

Recently, our group191 demonstrated a highly sensitive OECT-based glucose sensorby using PEDOT:PSS as the active transducing layer with unprecedented sensitivity, asshown in Fig. 19. Chitosan (CHIT), a kind of polymeric matrix with superior film-formingability and excellent biocompatibility,192 was used to immobilize GOx onto the Pt gate elec-trode. Nanomaterials, such as Pt nanoparticles (Pt-NPs) and multi-wall carbon nanotubes(MWCNTs), were also used to modify the gate electrode, which could greatly improve thesensitivity of the glucose sensors. For the Pt-NPs/Pt gate electrode, the device displayedan obvious response to an addition of 5 n M H2O2. The enhanced sensitivity toward H2O2

could be ascribed to the excellent electrocatalytic activity of the Pt-NPs. More interestingly,the devices with CHIT/GOx/Pt-NPs/Pt gate electrodes exhibited a pronounced response tothe addition of 5 nM glucose, which are three orders of magnitudes better than the deviceswithout the nanoparticles modified on the Pt gate. The improved performance on glucosedetection is the combined result of the outstanding electro-catalytic properties of nanoma-terials and the enlarged surface area for the enzyme immobilization. The fabricated OECTglucose sensors delivered a much better performance in comparison with conventional am-perometric methods. Based on the same principle, other types of highly sensitive enzymaticsensors could be realized by using gate electrodes with suitable surface modification.

4.3.2.2 Cell Sensors. Transistor-based cell sensors have gained considerable interestin the past many years, owing to the advantages of high sensitivity and miniaturization.Silicon-based transistors have already been successfully used for characterizing varioustypes of cells especially neural cells.194 Compared with Si-based transistors, OTFTs arelow-cost, flexible, and more biocompatible. Therefore, the integration of OECTs with cellsis promising in disposable sensing applications for in vitro or in vivo measurements. Ourgroup first reported the cell-based biosensors by using OECTs with PEDOT:PSS as theactive layers.195 We found that the OECTs showed excellent biocompatibility and verystable performance in culture medium for several days. Human esophageal squamousepithelial cancer cell lines (KYSE30) and fibroblast cell lines (HFF1) were successfullycultivated on the surface of PEDOT:PSS channels of the OECTs (Fig. 20). The devicesshowed pronounced horizontal shifts of transfer curves after the detachment of cultivatedcells on the surfaces treated by trypsin, which was attributed to the electrostatic interactionbetween the cells and the active layers of the devices. So the OECTs were sensitive to thesurface charge and the morphology of attached cells. Then we investigated the effect ofretinoic acid (anti-cancer drug) treatment on the activities of the cancer cells cultivated onthe OECTs. The transfer curves of the devices shifted to positive voltages in the first hourafter the treatment, indicating that retinoic acid can change the zeta potential of the cancercells. Then the transfer curves moved to negative gate voltages after one hour treatment

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Figure 19. (a) The schematic of the PEDOT:PSS based OECT glucose sensor. (b) The enzymaticreaction cycle essential for the determination of glucose by PEDOT:PSS-based OECTs. (c) ID as afunction of time for the OECT devices using Pt-NPs/Pt gate electrodes. VG = 0.4 V. H2O2 additions:a–e: 0.005, 0.05, 0.5, 5 and 10 μ M. (d)�Vg

eff vs. log[H2O2] curves for Pt (line I), MWCNT-CHIT/Pt(line II), and Pt-NPs/Pt (line III) gate electrodes. (e) ID as a function of time for the OECT devicesusing CHIT/GOx/Pt-NPs/Pt, VG = 0.4 V; Glucose additions: a–d: 0.005, 0.05, 0.5, and 5 μ M. (f)�Vg

eff vs. glucose concentration Cglucose curves for devices with Pt (line I), MWCNT-CHIT/Pt (lineII), and Pt-NPs/Pt (line III) gate electrodes. Reproduced from Tang et al.191 with permission fromWiley (Color figure available online).

due to the break and the detachment of cells from the PEDOT:PSS being consistent withthe sensing mechanism explained above.

Microarray technology has emerged as the influential tool for the rapid determinationof multi-analytes in the diagnostic studies. Therefore, OECT-based microarray biosensors

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Figure 20. (a) Schematic of an OECT cell sensor using PEDOT:PSS as the active layer. (b) Opticalimage of the trypsin treated cancer cells cultured on a PEDOT:PSS film. (c) Transfer curves of theOECT with attached cancer cells before and after trypsin treatment. Inset: output characteristics ofthe OECT before and after trypsin treatment. Reproduced from Lin et al.195 with permission fromWiley (Color figure available online).

are expected to play an important role in many practical applications. Most recently, ourgroup88 reported the array of micrometer-sized OECTs based on PEDOT:PSS fabricatedby the combination of photolithography and physical delamination, which could avoidthe damage on PEDOT:PSS layers by some organic solvents. The OECTs showed fasterresponse time and better stability in aqueous solutions than the normal millimeter-sizedOECTs. Human hepatoma (HepG2) cancer cell lines were successfully trapped on theOECTs by poly(ethylene glycol) (PEG) microwells around each device, which could beused for single cell analysis in the future (Fig. 21). All of the above works demonstrated anovel platform based on OECTs for sensing cell activities, which may find broad potentialapplications, such as for high throughput drug selection.

As the essential functional interfaces in multi-cellular organisms, barrier tissues areresponsible for the regulation of bio-fluids by controlling the levels of ions, bio-molecules,nutrients, and electrolytes. Therefore, the disruption or malfunction of barrier tissues couldserve as an indicator for disease states;196 however, present technologies for measuring thepermeability of barrier tissues are costly and time-consuming. Jimison et al.197 presentedthe OECTs for sensing barrier tissue integrity (Fig. 22). The Caco-2 cells layer were grown

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Figure 21. (a) Photograph of OECT arrays in a cell formed by PDMS walls on a glass substrate.(b) The response of channel current IDS of a micro OECT when different gate voltage was applied.Reproduced from Zhang et al.88 with permission from Elsevier (Color figure available online).

on a transwell membrane and incorporated into the device prior to the device charac-terization. The barrier properties of cells with tight junction (TJ) could increase the I DS

response time of the OECTs. The disruption of barrier tissue integrity could be induced bythe introduction of H2O2 or ethanol, resulting in a faster ion-transportation process in thetissue cell layer and thus a faster IDS response time of the OECT. They demonstrated that thesample assays showed no obvious change of the permeability up to 50 mM H2O2 exposure,while the barrier properties began to be disrupted after a 20% ethanol (EtOH) exposure.Most recently, the same group198 further demonstrated the capability of OECTs in barriertissue integrity assessment. This work offered a new approach for the assessment of barriertissue by using the OECTs.

PEDOT doped with tosylate(TOS) also offered many possible applications in biologicalsensing, due to the excellent electrical properties, biocompatibility, and simple processing.Berggren’s group199 demonstrated controllable cell-density gradients in an OECT based onPEDOT:TOS. The Madin–Darby canine kidney (MDCK) cells were successful seeded onthe surface of the active layer of the transistor. The cell-density gradients could be preciselycontrolled by the modulation of source and gate voltages. This work provided a feasibleway for the exact control of gradient characteristics of tissues and other cell clusters.

Most recently, Jimison et al.200 proved the superior biocompatibility of PEDOT:TOSand PEDOT:TOS:PEG materials. Polymer composite (PEDOT:TOS–PEG) with differentamount of PEG was coated on glass substrates, then HeLa cells with high activity wereperfectly transformed onto the surface of the active layer. The addition of PEG into thePEDOT:TOS active layer will not degrade the performance of OECTs. Taking advantageof the readily activated functional groups on the PEG chains, specific bioactive speciescould be easily biofunctionalized onto the OECTs to meet the requirements for certainapplications.

4.3.2.3 DNA Sensors. Krishnamoorthy et al. first explored201 the OECT-based label-free DNA sensors using PEDOT as the active layer. They extended the principle of theantibody-antigen sensor reported by Kanungo et al.203 to their work. The OECTs werefabricated by electropolymerization from ethylenedioxythiophene (EDOT) solution in the

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Figure 22. (a) Schematic of an OECT integrated with barrier tissue. (b) Equivalent circuit indicatingthe ionic transport between gate electrode and transistor channel. (c) In situ normalized responses(NR) of the OECT with barrier tissue as a function of time, with 1, 5, 50, and 100 mM H2O2 added.H2O2 was introduced at t = 0. NR = 0 corresponds to a confluent monolayer with full barrierproperties, while NR = 1 corresponds to a cell layer with no barrier properties. (d) NR of OECT withbarrier tissue on the introduction of ethanol (EtOH) with volume concentrations of 10, 20, and 30%.Ethanol was introduced at t = 0. Reproduced from Jimison197 with permission from Wiley (Colorfigure available online).

presence of probe ssDNA. The specific coupling effect of the probe DNA to its complemen-tary DNA could induce the morphological change and thus the resistance increase of thePEDOT channel. Significant response was observed when the OECT devices were exposedto the complementary ssDNA dissolved in PBS solutions, whereas no detectable responsewas displayed when they were exposed to non-complementary ssDNA. A detection limit of80 ng mL−1 and a good linear region from 8 × 10−8 to 1 × 10−5 g/mL to the complementaryssDNA were obtained. In addition, the performance of the DNA sensors could be enhancedby increasing the length of the ssDNA probes.

Recently, our group202 reported a platform for labeling-free DNA detection based onthe OECTs with an unprecedented detection limit to DNA targets. As shown in Fig. 23,the OECTs based on PEDOT:PSS were fabricated on mechanically flexible polyethylene

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Figure 23. (a) Schematic of an OECT-based DNA sensor integrated in a flexible microfluidic systemwith a PET substrate. (b) Photographs of the device bent to both sides. (c) Transfer curves of the devicebefore and after the immobilization and the hybridization of DNA. Inset: the three transfer curveswere merged by horizontal shift. (d) |�VG| vs. the concentration of DNA targets in PBS solutions.Inset: the voltage pulse applied on the gate to facilitate the hybridization of DNA. Reproduced fromLin et al.202 with permission from Wiley (Color figure available online).

terephthalate (PET) substrates and then incorporated into microfluidic channels, whichforwarded this sensing technique to a more practical “lab on a chip” system. Single strandedDNA probes were immobilized on the Au gate electrodes of the OECTs. Owing to thesurface dipole formed by the inherent negative charge of DNA molecules, the work functionsof the Au gate electrodes were decreased after the immobilization or the hybridization ofDNAs, which led to a change of device performance. The devices were very stable underbending stress and no obvious change in the device performance was observed when theywere bent to both sides. The DNA sensors were able to detect complementary target DNAdown to 1 nM. The detection limit was further extended to 10 pM by applying continuousvoltage pulses on the gate electrode during DNA hybridization, which could enhance thehybridization of DNA on the gate. It is reasonable to expect that many other types ofsensors can be realized with the same principle. So the OECT-based sensors integrated inflexible microfluidic systems are suitable for cost-effective, highly sensitive, and disposablechemical and biological sensing.

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4.3.2.4 Antibody/Antigen Sensors. The electrically conductive polymers have beenreported to be biocompatible, which opened the door for the effective immobilization ofbio-molecules in the fabrication of high-performance OECT-based biosensors. OECTs havebeen used for sensing analyte-receptor binding events. Contractor et al.144 proposed thatmorphology (or conformation) change of the active layer could lead to the switching ofdevice performance. The morphology change could be introduced simply by employing theso-called host-guest system, in which one part acts as a detector for the other one. Becauseof the bio-specificity of the binding effects in the system, these kinds of devices coulddisplay extremely high selectivity toward the target analytes. One of the most interestingbinding systems is the antigen-antibody binding, which is essential to effective immunosen-sors. Kanungo et al.203 demonstrated antibody sensors based on this principle. During thefabrication of those devices, different amount of the goat antirabbit IgG antibody wereadded in the PEDOT electropolymerization solution. The sensitivity of the immunosensorwas improved by increasing the load of the antibody in the polymer matrix. A detectionlimit down to 1 × 10−10 g/mL antigen was achieved in a response time of 3 minutes.Furthermore, they compared the response of the device fabricated by physical adsorptionof the goat antirabbit IgG antibody with the one fabricated by the post-polymerizationprocess. The results indicated that physical adsorption of the antibody in the device led toa much weaker response to the antigen.

Kim et al.204 reported a type of OECT-based immunosensor for the detectionof prostate specific antigen/1-antichymotrypsin (PSA-ACT) complex, as shown inFigs. 24a–24c. The PEDOT:PSS active layer was firstly functionalized with 3-aminopropyldiethoxymethylsilane (APTMS), then treated with ProLinker molecules, andfinally immobilized with PSA monoclonal antibody (PSA mAb). The detection limitof the device was down to 100 pg/mL. The sensing mechanism was attributed to thereduced de-doping effect in the PEDOT:PSS channel induced by negative surface charge ofPSA-ACT complex. Moreover, the detection limit of PSA could be dramatically improved(down to 1 pg/mL) by using gold nanoparticles (AuNPs) conjugated with PSA polyclonalantibody (PSA pAb). The electron transfer from the AuNPs-PSA pAb to the PSA-ACTcomplex/PSA mAb was greatly facilitated due to the larger effective surface area of theAuNPs, which resulted in higher sensitivity to the target.

4.3.2.5 Bacteria Sensor. Bacterial pathogens identification is extremely important toour healthcare activities. Conventional techniques for the analysis of bacterial pathogens,for example, polymerase chain reaction or enzyme linked immunosorbent assays, are time-consuming and are always involved with complex electric equipments.205 Therefore, noveltechniques, which are suitable for the rapid detection of bacterial pathogens, are in greatdemand. Our group206 first proposed the disposable bacterial sensor based on OECTs (Figs.24d–f). Enterohemorrhagic Escherichia coli (E. coli) O157:H7, the widespread foodbornepathogen, was detected by the sensors in KCl solutions. PEDOT:PSS active layers ofthe OECTs were firstly treated with oxygen plasma and 3-glycidoxypropyl- trimethoxysi-lane (GPMS) toluene solution followed by the immobilization of the anti-Escherichia coliO157:H7 antibody on the surface. The transfer curves of the OECT were shifted to highergate voltage after the capture of E. coli O157:H7 by the antibody. This result was attributedto the influence of negative charge from the surface of bacteria on the active layer, which issimilar to the electrostatic interaction between cells and OECTs introduced above.195 Wefound that the voltage shift is dependent on the ion concentration of KCl solution. WhenKCl concentration is higher than 1 mM, the negative charges on the bacteria were screenedby electric double layer and thus the device showed little response to the bacteria. Wedemonstrated that the device could detect bacteria concentrations down to 103 cfu mL−1 in

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Figure 24. (a) Schematic of the OECT-based immunosensor utilizing gold nanoparticles for thedetection of PSA–ACT complex. (b) Transfer curves of OECT immunosensor with 10 nm dia.AuNPs–PSA pAb in different concentrations of targets. (c) Calibration curve for the transfer curvesin (b). Reproduced from Kim et al.204 with permission from Elsevier. (d) The schematic of an OECT-based E. coli O157:H7 sensor with an active layer of PEDOT:PSS. (e) Relative gate voltage shifts ofan OECT vs. bacteria concentrations. (f) Schematic diagram of potential drops in the electric doublelayers (EDL) near the channel/electrolyte and electrolyte/gate interfaces, before and after the captureof E. coli O157:H7 on the PEDOT:PSS surface. Reproduced from He et al.206 with permission fromRSC (Color figure available online).

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the optimized condition. This pioneering work opened the door for developing OECT-basedsensors for detecting various types of bacteria.

4.3.2.6 Dopamine Sensor. As one kind of neurotransmitter, dopamine level in a humanbody is very important. For example, Parkinson’s disease is predominately caused bythe dysfunction of the dopaminergic neuron process. Our group207 developed an OECT-based sensor for the detection of dopamine concentration for the first time. Different gateelectrodes, including graphite, Au and Pt electrodes, were adopted for the PEDOT:PSS-based OECTs. We found that the gate electrode material is important to the responseof the device to dopamine. The OECT with a pure Pt gate electrode characterized at0.6 V gate voltage showed the highest sensitivity. The detection limit of the devices todopamine was lower than 5 nM. Due to the electro-oxidation of dopamine at the gateelectrode, the effective gate voltage was increased upon the addition of dopamine, whichled to the horizontal shift of the transfer curve to lower gate voltages. Compared withthe conventional electrochemical method, the detection limit was improved for about twoorders of magnitude. The OECT-based dopamine sensor is potentially useful in the rapiddetermination of dopamine level for clinical disgnostics (Fig. 25).

4.3.2.7 Other Sensors. Lactate plays a significant role in the anaerobic metabolism.Khodagholy et al.193 demonstrated that an OECT-based lactate sensor integrated with roomtemperature ionic liquids (RTILs) as a solid-state electrolyte. The sensing mechanism ofthe OECT is attributed to the reaction of lactate on the surface of PEDOT:PSS activelayer. They demonstrated the normalized response of the transistor as a function of lactateconcentration in the range of 10–100 mM, covering the relevant physiological ranges oflactate presented in saliva and blood.

Recently, liposome-based structures are extensively investigated for the effective drugdelivery.208 Real-time monitoring of liposome may unveil significant information of adrug delivery system. Tarabella et al.209 reported a sensitive OECT for liposome-basednanoparticle detection on a wide dynamic range down to 10−7 mg/mL, which can meet therequirement for typical drug loading and drug delivery conditions. The device integratedwith microfluidics could discriminate successive injection of different liposomes, whichprovided a viable solution to the quality-control assays in the pharmaceutical industry.

4.4 Poly(3-akylthiophene) (P3AT)

As one of the most intensively studied semiconducting polymers, poly(3-akylthiophene)was widely used in organic electronics. The alkyl group can improve the solubility of thepolymer in organic solvents and facilitate the device fabrication by solution process.210

Wrighton et al.211 first reported the OECTs using poly(3-methylthiophene) (P3MT) as theactive layer. The OECTs displayed a high stability in aqueous electrolytes, with the pHranged from 1 to 9. Due to the conductivity change of P3MT induced by the reversibleoxidation reaction, the device could be used to detect the chemical oxidant IrCl62−, with alimit amount down to 10−15 mol.

They212 then exploited the P3MT-based OECTs to detect H2 and O2 gases, and the pHvalue of electrolytes. To catalyze the reaction between H2 and O2, P3MT film was impreg-nated with platinum particles. They demonstrated that the Pt particles did not significantlychange the performance of the OECTs on amplifying the electrical signals. So the role of Ptis purely as a catalyst to allow equilibration of O2, and H2, with the polymer. The P3MT/Ptbased sensors could be used to determine the pH value ranging from 0 to 12.

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Figure 25. (a) The schematic of an OECT-based dopamine sensor with an active layer of PE-DOT:PSS. (b) The electro-oxidation reaction of dopamine occurred near the surface of the gateelectrode. (c) Typical ID vs. time curve of an OECT using the Pt gate electrode upon the additions ofdopamine in PBS (pH 7.2) solution. Dopamine additions, a–f: 0.5, 5, 50, 500, 1000, and 1000 nM.VG = 0.6 V, VDS = −0.1V. (d) The change of the effective gate voltage (�Vg

eff) of the OECT vs. theconcentrations of dopamine [Cdopamine]. Reproduced from Tang et al.207 with permission from Elsevier(Color figure available online).

4.5 Polycarbazole

As a nitrogen-containing aromatic conjugated polymer, polycarbazole have a conductivityof approximately 1.4 × 10−4 S/cm, which is enough for sensing application. Saxena et al.213

demonstrated a Cu2+-selective OECT sensor with polycarbazole as the active layer, makinguse of the fact that un-doped polycarbazole film is highly selective to Cu(II) ions withoutadditional ion-selective membrane. The device could detect Cu2+ down to 2.5 × 10−6 Mand reached the upper saturation region of about 10−4 M. The conductivity change of thepolycarbazole film was due to the conformational changes in the presence of Cu2+. Similar

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measurements on other metal ions were conducted to verify that the high sensitivity wasspecifically for Cu2+ ions.

5. Conclusions and Outlook

Holding advantages in synthetic tailorability, tunable electronic conductivity, simplifiedsolution processable fabrication methods, as well as the excellent biocompatibility, organicsemiconductors have emerged as the viable candidates for the transistor-based sensors.Typically, PEDOT:PSS, poly(3-akylthiophene), pentacene, and polyaniline, etc., are themost common ones for OTFT sensing application. OTFTs, including OECTs and OFETs,have been intensively investigated for lots of applications in chemical or biological sensors.It is worth noting that OECT-based sensors with stable performance in aqueous solutionsare suitable for in situ or even in vivo detections. A large number of sensors, includingions, pH, DNA, cell, glucose, NADH, and dopamine sensors, have already been reportedwith high performance. Some of the sensors exhibited much better sensitivity than typicalelectrochemical detections.

Research on the novel organic semiconducting materials for the OTFT-based sensorsexponentially surged in the past few years; various types of semiconducting polymers withsuperior carrier mobility and excellent stability have been synthesized and used in sensingapplications. The next big leap will be the further improvement of these devices for realapplications, which could be accomplished via two ways. The basic one is the optimiza-tion of the existing OTFT sensors. The three key parameters of the sensors (sensitivity,selectivity, and stability) could be manipulated by optimizing the geometric features of thedevices or adopting new techniques and materials in the devices fabrication. The secondone is the development of novel types of sensors based on OTFTs. The versatility of OTFTsleads to a new era in the design of novel sensors. From the individual perspective of theauthors, OTFT-based sensors hold enormous potential in both academic study as well asdaily practical applications, although some essential issues of the sensors, including de-vice reproducibility, reliability, and mass production techniques still need to be extensivelyinvestigated.

Acknowledgments

This work is financially supported by the Research Grants Council (RGC) of Hong Kong,China (project number: PolyU5322/10E) and Hong Kong Polytechnic University (projectnumber: 1-ZVAW).

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