In Vivo Analysis with Electrochemical Sensors and Biosensors

Review pubs.acs.org/ac

In Vivo Analysis with Electrochemical Sensors and Biosensors Tongfang Xiao,†,‡ Fei Wu,†,‡ Jie Hao,†,‡ Meining Zhang,†,‡ Ping Yu,†,‡ and Lanqun Mao*,†,‡ †

Beijing National Laboratory for Molecular Sciences, CAS Key Laboratory of Analytical Chemistry for Living Biosystems and Photochemistry, Institute of Chemistry, Chinese Academy of Sciences, Beijing 100190, China ‡ University of Chinese Academy of Sciences, Beijing 100049, China

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CONTENTS

Electrochemical Sensors Differential Pulse Voltammetry Fast Scan Cyclic Voltammetry Amperometry Based on Unique Redox Property of Neurochemicals Ascorbate DA O2 NO Hydrogen Peroxide Hydrogen Sulfide Electrochemical Biosensors Oxidase-Based Electrochemical Biosensors Dehydrogenase-Based Electrochemical Biosensors Multienzyme-Based Electrochemical Biosensors Aptamer-Based Biosensor for ATP Conclusion Author Information Corresponding Author ORCID Notes Biographies Acknowledgments References

scavenger in cells due to its electron donating property. With sufficient evidence came the recognition of its crucial role in regulating DA- and glutamate-mediated signal transduction in addition to neuroprotection.5,6 Besides, a huge amount of biologically important small ions (H+, K+, Na+, Ca2+, Mg2+, etc.), energy suppliers (glucose, ATP, etc.), amino acids, peptides, lipids, proteins, and nucleic acids (DNA and RNA) are actively involved in cell growth, replication, response, communication, and organization in the neuronal network. Dynamics of neurochemicals in terms of their interactions, transformations, and distributions ultimately define the unique neuron and brain structures and functions and digging into the chemical nature of intricate neuronal activities makes up the research body in neurochemistry. A 3.8 billion-year evolution of the CNS has brought an exceptional complexity to its chemical communities, making neurochemistry so distinguished from pure chemistry in a flask or other biochemistry and difficult to understand by means of traditional chemical methodologies. Fortunately, scientists have made tremendous progresses in neurochemistry in an accelerating pace since its infancy decades ago, which is propelled by new micro- or nanosized concepts and techniques, in particular concerning qualitative and quantitative monitoring of the (near) real-time change of neurochemicals in the brain of a living animal (termed in vivo analysis).7 There have been several excellent reviews and books summarizing all developed methods for in vivo monitoring of neurochemicals.1−3,8−28 This Review will be narrowed down on the electrochemical sensors and biosensors for in vivo analysis, with an emphasis on the latest advancements in this field that have not been reviewed elsewhere.

A B B D D E F F F G G G I J J K K K K K K L L

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hemistry constitutes the essence of life, with the central nervous system (CNS) not being an exception. As the most sophisticated and mysterious biological system, the CNS arises from a complex combination of chemical processes involving neurotransmitters and neuromodulators at different biological levels ranging from single vesicles, single cell/ synapse, brain slices, brain regions to the entire brain. Typically, neurotransmitters are pooled inside vesicles in neurons and released by exocytosis for synaptic transmission that regulates all physiological or pathologic processes in the CNS.1,2 For instance, dopamine (DA) is a well-known neurotransmitter essential for the establishment of reward-based behaviors, memory, and addiction in mammals, and depletion of dopamine is a cause of several neural diseases including Parkinson’s disease.3 5-Hydroxytryptamine (5-HT) is another important neurotransmitter closely related to cognition, mood, emotions and memory formation in the brain, and shown to exert therapeutic effects on Parkinson’s disease and Alzheimer’s disease in recent studies.4 Neuromodulators do not directly participate in but modulate neurotransmissions. One example is the ascorbate as the famous small antioxidant and free-radical © XXXX American Chemical Society

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ELECTROCHEMICAL SENSORS Many of the neurochemicals such as DA, ascorbate, oxygen (O2), 5-HT are electroactive and can be potentially detected electrochemically. However, the complexity of the CNS gives rise to great challenges to electrochemical analysis of these neurochemicals with the following criteria to be considered: (1) high selectivity to detect the neurochemicals of interest out of an ocean of interfering species in the CNS, (2) good stability to ensure short- and long-term sensing, (3) antiprotein-fouling ability, and (4) capability to detect basal levels. Furthermore, extensive overlap of peak potentials for the redox reactions of neurochemicals at the electrodes, especially the carbon fiber electrodes (CFE) widely used by neurochemists, normally Special Issue: Fundamental and Applied Reviews in Analytical Chemistry 2017 Published: November 7, 2016 A

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accuracy and extension to in vivo monitoring of other neurochemicals besides DA, which will be highlighted below. Yang et al. increased the sensitivity for in vivo detection of DA by modifying a metal microelectrode with carbon nanotubes (CNTs). CNTs were successfully grown on niobium substrate under chemical vapor deposition for the first time. CNTs coated on Nb wires were shorter, denser, and more aligned than CNTs grown on other substrates. The prepared CNT-coated niobium (Nb) microelectrodes exhibited higher sensitivity and lower ΔEp value as compared with those of CNTs grown on CFEs or other metal wires by FSCV. The detection limit for DA at the CNT-Nb microelectrodes was approximately 2-fold lower than that at bare CFEs. The ratios of oxidative current for ascorbate and DOPAC to that of DA at the CNT-Nb microelectrodes were significantly smaller than those at CFEs (Table 1), demonstrating that the CNT-Nb microelectrodes exhibited higher selectivity for DA over ascorbate and DOPAC than CFEs did.43

invalidates conventional electrochemical sensors for in vivo applications. So far, a few strategies have been developed to differentiate the peak potentials corresponding to various neurochemicals during in vivo analysis, as summarized below. Differential Pulse Voltammetry. Differential pulse voltammetry (DPV) is a combination of linear sweep voltammetry and square wave techniques. The applied signal is a small amplitude square wave (∼25 mV) at a constant frequency superimposed on a slow linear potential ramp.3 The charging currents are strongly discriminated and the ratio of faradaic to charging current is large thus making DPV a highly sensitive voltammetric technique.29 The differential current of DPV is a symmetric voltammetric peak, of which the intensity is proportional to the concentration of the analyte. The oxidation potentials of two compounds separated by more than 100 mV can be simultaneously measured without interference from each other, improving the sensing selectivity toward a specific substance. Lane et al. first used DPV to in vivo record two separate oxidation peaks, which were attributed to the oxidation of ascorbate and catecholamine in caudate nucleus of rats at a platinum electrode treated with aqueous iodide.30 Gonon et al. used DPV to record two peaks in neostriatum of rats with electrochemically treated CFE. Combined with the pharmacological investigations, they ascribed the peak at the negative potential to the oxidation of ascorbate and the one at the positive potential to the oxidation of 3,4-dihydroxyphenlacetic acid (DOPAC) other than DA.31 Crespi et al. also used DPV method with electrochemically treated CFEs to perform in vivo measurements and recorded four well-resolved oxidation peaks. They attributed the peaks to the oxidation of ascorbate, catcholamine, 5-hydroxyindoleacetic acid (5-HIAA), and homovanillic acid (HVA).32 Because of its high sensitivity and selectivity, DPV has been used for the simultaneous detection of multiple neurotransmitters in vivo.33−40 Chai et al. used DPV to selectively monitor Cu2+ in vivo by a two-channel ratiometric electrochemical sensor. They designed a Cu-free derivative of superoxide dismutase (SOD), i.e., E2Zn2SOD, and used it as the biomolecular recognition element based on the specific interaction between E2Zn2SOD and Cu2+. The accuracy of the sensor was further improved with an inner reference channel to provide a built-in correction, ensuring its function reliability in the complicated brain environment.41 More recently, another two-channel electrochemical ratiometric sensor was developed by Zhao et al. for local pH determination in the live brain of rats. They synthesized N-(6aminopyridin-2-yl) ferrocene to indicate pH change, in which pyridine unit was used as a recognition element for protons and ferrocene unit as an electroactive probe. As pH altered, the anodic and cathodic potentials of the sensor shifted accordingly with high selectivity, linearity, and sensitivity. With this sensor, they observed different pH changes under global cerebral ischemia in different brain regions of rats.42 Fast Scan Cyclic Voltammetry. To improve the temporal resolution of the electrochemical methods to meet the requirements of fast dynamics of the neurotransmitter release, fast scan cyclic voltammetry (FSCV) was developed and used to monitor the neurotransmitters in vivo with high sensitivity, selectivity, as well as high temporal resolution. For more details on FSCV, please refer to the previous reviews contributed by Wightman et al.3 Recent several years have witnessed some progresses on FSCV in terms of improvement of sensitivity and

Table 1. Average ΔEp, Current Density, and Limit of Detection for 1 μM DA at CNT-Grown Microelectrodes and CFEsa electrode CNT-Nb CNT-Ta CNT-CF CFME

ΔEp (V)

current density (pA/μm2)

± ± ± ±

197 ± 16 82 ± 10 100 ± 25 135 ± 24

0.73 0.87 0.81 0.67

0.03 0.01 0.03 0.01

LOD (nM) 11 91 46 19

± ± ± ±

1 27 10 4

a All n = 5; errors are standard error of the mean. Reproduced from Yang, C.; Jacobs, C. B.; Nguyen, M. D.; Ganesana, M.; Zestos, A. G.; Ivanov, I. N.; Puretzky, A. A.; Rouleau, C. M.; Geohegan, D. B. Venton, B. J. Anal. Chem. 2016, 88, 645−652 (ref 43). Copyright 2016 American Chemical Society.

Keithley et al. observed that the increase of scan rates essentially resulted in the enhancement of in vivo sensitivity toward DA without increasing quantization error. They found that signal-to-noise ratios were increased approximately by 4 folds upon increasing the scan rate from 400 to 2400 V/s with the 1.0 V waveform.44 Mathematical analysis methods were used to exactly calibrate the DA data obtained by FSCV as a complement for the modification methods for in vivo use.45,46 Wightman et al. first employed principal component regression (PCR) as a useful method for the detection of DA and pH changes measured with FSCV. As demonstrated, the PCR data reduction method could accurately reproduce the data set, and PCR proves an effective method for predicting concentration.47 In the recent work, they furthermore improved the multivarate prediction of neurochemicals detected with FSCV using PCR for in vivo analysis.48 Roberts et al. developed a model that utilized information contained in the background charging current to predict electrode sensitivity to DA, ascorbate, hydrogen peroxide (H2O2), and pH shifts during in situ electrochemical experiments. They obtained a good correlation between predicted sensitivity and values obtained by the traditional postcalibration method, across all analytes, and proved this approach useful for in vivo analysis.49 To solve the problem associated with electrode fouling in FSCV, Singh et al. investigated the effects of different coating materials, i.e., Nafion and base-hydrolyzed cellulose acetate (BCA), on the sensitivity, selectivity, and biofouling of CFEs with monoamine neurotransmitter as the target. After careful B

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Figure 1. PEDOT:Nafion electrodes resist (A) synthetic fouling (40 g/L BSA in pH 7.4 aCSF) and (B) in vivo biofouling. The DA detection waveform (− 0.4 V to +1.3 V) was continuously applied for the duration of synthetic and in vivo biofouling experiments. Asterisks indicate statistical significance when compared to uncoated CFMEs. Error bars are SEM (n = 3−4). (C) Electron micrograph of two representative uncoated CFEs removed after 30 min implantations in the nucleus accumbens or prefrontal cortex show large accumulations of biomaterial on the surface of the electrode. (D) Electron micrograph of a LD PEDOT:Nafion-coated CFE implanted in the nucleus accumbens for 6 h shows decreased accumulation of biomaterial when compared to uncoated CFEs. Statistical significance is marked with asterisks (*P < 0.05, **P < 0.01), and selectivity of treated electrodes is compared to uncoated CFEs. Reproduced from Vreeland, R. F.; Atcherley, C. W.; Russell, W. S.; Xie, J. Y.; Lu, D.; Laude, N. D.; Porreca, F.; Heien, M. L. Anal. Chem. 2015, 87, 2600−2607 (ref 51). Copyright 2015 American Chemical Society.

simplified the procedures for electrode calibration for in vivo detection of DA with FSCV. In addition to its application for in vivo detection of DA and 5-HT, FSCV was recently used for in vivo monitoring of other physiologically important species.54−58 Somber et al. used FSCV to in vivo record H2O2 change with a single, uncoated CFE.59 To overcome the kinetic limitation, the carbon surface was electrochemically conditioned on the anodic scan and H2O2 was irreversibly oxidized on the cathodic scan at ∼+ 1.2 V. They verified the identity of the signal by monitoring the selective enzymatic degradation of H2O2 in the presence of catalase and simultaneously monitored H2 O2 and DA fluctuations in intact striatal tissue under basal conditions and in response to the initiation of oxidative stress. They assessed the effect of acute increases in local H2O2 concentration on both electrically evoked DA release and basal DA levels.60 Venton et al. recently developed a new method for the measurement of adenosine by combining FSCV and specific oxidation properties of adenosine.61 The method exhibited good selectivity and excellent spatial and temporal resolution and was used to monitor adenosine change on a subsecond time scale. With this method, they found that the implantation of a probe into brain tissue caused a transient adenosine response, which could be neuroprotective.62,63 Very recently, they improved the selectivity of the method by using a modified sawhorse waveform to maximize the time for adenosine oxidation and to manipulate the shapes of cyclic voltammograms of analytes that oxidize at the switching potential.64 By using this method, mechanically evoked adenosine was identified with PCA and changes in the ratio of ATP to

comparison, they came to the conclusion that electrode responses to different coating processes and fouling were complex, with some coatings better suited for sensitivity and selectivity (i.e., Nafion), whereas others were better at preventing fouling (i.e., BCA or fibronectin).50 Vreeland et al. described a biocompatible microelectrode coated by PEDOT:Nafion composite with an antifouling characteristic for detection of DA in vivo. The coated electrodes showed increased sensitivity and selectivity toward DA. Moreover, the sensitivity of the PEDOT:Nafion-coated CFEs to DA over the in vivo experiments lost only 9 ± 5% of the precalibration sensitivity, as typically shown in Figure 1.51 Chandra et al. demonstrated a strategy to minimize electrode fouling by hydrogenation of conical-tip carbon electrodes for in vivo DA detection. They found that the hydrogenation could produce hydrophobic sp3 carbon surface that deters adsorption of amphiphilic lipids, proteins, and peptides. During DA detection in vivo with the hydrogenated carbon electrodes, over 70% of the DA oxidation current remained after the first 30 min of a 60 min experiment, and at least 50% remained over the next half-period.52 Very recently, Liu et al. proposed another strategy to exactly calibrate the DA concentration through solving the electrode calibration problem caused by electrode fouling by pretreatment of CFEs with bovine serum albumin (BSA).53 They found that the pretreatment of the CFEs with BSA essentially minimized the further interference from the proteins in the live brain of rats. They also found that the results obtained for the electrode precalibration in artificial cerebral spinal fluid (aCSF) containing BSA were almost the same as those for the electrode postcalibration after in vivo experiments, which largely C

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Figure 2. (A) Typical cyclic voltammograms (CVs) obtained at the heat-treated SWNT-modified GC electrodes in 0.10 M phosphate buffer (pH 7.0) in the absence (dotted lines) and presence (solid lines) of 0.5 mM ascorbate. Scan rate, 50 mV s−1. Reproduced from Zhang, M.; Liu, K.; Gong, K.; Su, L.; Chen, Y.; Mao, L. Anal. Chem. 2005, 77, 6234−6242 (ref 72). Copyright 2005 American Chemical Society. (B) Schematic diagram of the OECS for continuous monitoring of ascorbate in the CNS with the heat-treated SWNT-modified GC electrode as the detector. Reproduced from Zhang, M.; Yu, P.; Mao, L. Acc. Chem. Res. 2012, 45, 533−543 (ref 6). Copyright 2012 American Chemical Society. (C) DPVs at the CNT-modified CFE in the striatum of the anesthetized rat before (black) and after (red) exogenous infusion of AOx (39.3 U mL1) into the striatum. Reproduced from Zhang, M.; Liu, K.; Xiang, L.; Lin, Y.; Su, L.; Mao, L. Anal. Chem. 2007, 79, 6559−6565 (ref 70). Copyright 2007 American Chemical Society. (D) Current responses recorded for the dialysate continuously sampled from the rat striatum with aCSF and aCSF containing 47.2 U mL−1 of AOx as the perfusion solutions. The electrode was polarized at +30 mV. Flow rate, 3 μL/min. Reproduced from Zhang, M.; Liu, K.; Gong, K.; Su, L.; Chen, Y.; Mao, L. Anal. Chem. 2005, 77, 6234−6242 (ref 72). Copyright 2005 American Chemical Society.

electrode preactivation conditions is required to achieve the optimum electrode performance. From an electrochemical perspective, redox transformation of ascorbate on carbon electrodes undergoes an inner-sphere electron transfer and is sensitive to the nature of electrode surface, which can be fine-tuned by physical or chemical modifications. Early efforts have demonstrated that ascorbate oxidation could be promoted by preheated or electrochemically pretreated carbon electrodes featuring diverse surface characteristics, such as morphology, cleanliness, electronic structure, functionality, hydrophilicity/hydrophobicity, etc.68,69 In line with this point of view, Zhang et al. examined ascorbate oxidation at a series of carbonaceous electrodes including glassy carbon (GC), graphite, single-walled CNTs (SWNTs), and heat-treated SWNTs and observed, for the first time, that ascorbate oxidation was dramatically accelerated at the surface of heat-treated SWNTs (−0.05 V vs Ag/AgCl), as shown in Figure 2 A.70,71 On the basis of this finding, they could realize selective determination of ascorbate in vivo with the CNTconfined CFEs. Fast electrochemical kinetics of ascorbate oxidation was elucidated to be favored by the distinct electronic structures and excellent electrode reactivity of CNTs. Moving forward from the CNT-modified electrodes, Zhang et al. developed an online electrochemical system (OECS) for ascorbate sensing in combination with in vivo microdialysis (Figure 2B).72 The CNT-based electrochemical detector in the OECS showed a high selectivity toward ascorbate, as illustrated by a clear current decrease to the baseline when AOx was added into the microdialysate sampled from the live brain of rats perfused with aCSF (Figure 2D) and a satisfactory stability as well as reproducibility during continuous monitoring of cerebral ascorbate. Meanwhile, the detector exhibited remarkable tolerance against O2 and pH fluctuation induced by

adenosine were observed after manipulating ATP metabolism with POM-1 in slices. Amperometry Based on Unique Redox Property of Neurochemicals. In addition to the in vivo electrochemical sensors mentioned above that are mainly based on the uses of electrochemical techniques to achieve the selectivity and sensitivity, the electrochemical sensors could also be developed for in vivo applications by fully exploiting the electrochemical redox nature of neurochemicals themselves and rationally designing the electrode/electrolyte interfacial structure, as demonstrated below. Ascorbate. Ascorbate is the anionic form of ascorbic acid, a water-soluble, six-carbon sugar acid with two dissociable protons (pKa 4.04 and 11.34), under physiological conditions. Electrochemically, a two-electron and one-proton oxidation (−0.23 V vs Ag/AgCl) converts ascorbate into dehydroascorbic acid, which is rapidly hydrolyzed (1.2 × 103 s−1) into 2,3diketogulanic acid, an electroinactive product readily adsorbed on electrode surfaces.65 In consequence, electrode fouling by 2,3-diketogulanic acid becomes a big hurdle for in vivo selective sensing of ascorbate because it leads to elevated potential for ascorbate oxidation (∼+0.30 V vs Ag/AgCl), which overlaps with the potential window for the oxidation of catecholamine and uric acid. In a very early study by Brazell et al., ascorbate oxidase (AOx) was exogenously infused to distinguish the oxidative current of ascorbate from unintended currents recorded in vivo.66 Gonon et al. later reported an electrochemical pretreatment of CFEs in their attempt to selectively monitor cerebral ascorbate.67 However, the effect of such an electrochemical treatment on electrode selectivity toward ascorbate may vary with carbon fibers from different sources. In fact, a time-consuming trial-and-error procedure in search of D

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Figure 3. SEM images of the (A) VACNT-CF, (B) Pt/CF, and (C) Pt/VACNT-CF. Scale bar, 10 μm. Inset in part C, enlarged SEM image of Pt/ VACNT-CF. (D) Amperometric response for the hippocampus O2 recorded in anesthetized rats during global ischemia/reperfusion. Reproduced from Xiang, L.; Yu, P.; Zhang, M.; Hao, J.; Wang, Y.; Zhu, L.; Dai, L.; Mao, L. Anal. Chem. 2014, 86, 5017−5023 (ref 89). Copyright 2014 American Chemical Society.

in vivo ascorbate sensing. As a step further, a Mg2+ sensing unit built on an ITO glass substrate coated by a toluidine blue O (polyTBO) film was incorporated into the microfluidic chip sensor to achieve in vivo selective detection of ascorbate and Mg2+ avoiding electrode cross-talk. The dual-component sensor was proved highly selective, stable, reproducible, and durable for continuous and simultaneous monitoring of ascorbate and Mg2+ in cerebral systems.78 Facilitated electron transfer for ascorbate oxidation at CNTs has also driven the pursuit of electrochemical sensors for in vivo real-time monitor of ascorbate with CNT-modified CFEs. Using vertically aligned CNT-sheathed carbon fibers (VACNTCFs) as highly reproducible pristine microelectrodes for in vivo monitoring of ascorbate, Xiang et al. observed that a significant acceleration in ascorbate oxidation on electrodes electrochemically pretreated in NaOH solution, suggesting that the electron transfer of ascorbate may be facilitated at the oxidized and opened tips of CNTs.79 DA. In addition to in vivo FSCV, an amperometric sensor based on inherent electrochemical properties of DA was developed to achieve desired in vivo sensitivity and selectivity. Xiang et al. selectively determined DA free from the interference of ascorbate and DOPAC with a GC electrode, on which laccase was cross-linked into a multiwalled carbon nanotube (MWNT) layer.80 Specifically, DA was irreversibly oxidized into dopamine-o-quinone (DOQ) under the catalysis of laccase with O2 being reduced to water. DOQ then underwent a pH-dependent deprotonation and an intramolecular Michael addition to give 5,6-dihydroxyindoline, followed by a disproportionation into 5,6-dihydroxyindoline quinone. The final quinone product could generate a current response at a negative potential on MWNTs, therefore separating the oxidation of DA from those of ascorbate and

cerebral ischemia/reperfusion. These advantages together enabled the CNT-based OECS to assess the neuroprotection effect of ascorbate against brain ischemia. With the same system, Liu et al. observed different changes of ascorbate concentrations with respect to investigated brain regions of rats during the acute periods of two-vessel occlusion (2-VO) ischemia and left middle cerebral artery occlusion ischemia models, which might be the synergic consequence of anoxia depolarization, glutamate-ascorbate heteroexchange, cell necrosis, and overproduction of reactive oxygen species.73,74 In a recent study with the CNT-based OECS, they detected an attenuation of ascorbate concentration increase in hippocampus following 2-VO ischemia by immediate intravenous administration of ascorbate or glutathione within 10 min of ischemia and thus demonstrated the neuroprotection effect through antioxidant-induced oxidative stress alleviation in vivo.75 Similar finding with the CNT-based OECS was provided by Li et al., who observed that the increase of ascorbate concentration in the olfactory bulb of rats during the acute period of olfactory dysfunction induced by intraperitoneal injection of 3-methylindole (3-MI) was effectively alleviated by intravenously administrated ascorbate or glutathione within 10 min after injection of 3-MI. Involvement of ascorbate in early stages of 3-MI induced olfactory dysfunction was suggested as a useful clue to understanding the molecular mechanism of olfactory disorders.76 The CNT-based OECS was recently integrated into a microfluidic chip-based OECS. Compared with conventional thin-layer flow cell detectors, the microfluidic chip-based OECS bears the advantages of small dead volume, low cost, and flexibility for in vivo multicomponent sensing.77 Gao et al. fabricated a single-channel microfluidic chip that utilized indium−tin oxide glass modified with CNTs as the probe for E

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lophthalocyanines and metal oxides, to lower the overpotential for redox transformations of NO are good options.93,94 Rivot et al. managed to monitor NO in vivo within the dorsal horn of the lumbar spinal cord of decerebrated-spinalized rats with a NO sensor based on nickel porphyrine and Nafion.95,96 Park et al. deposited Pt−Fe (III) oxide nanocomposites on Pt electrodes to electrochemically monitor NO with less interference from carbon monoxide (CO).97 They used NO microsensor to simultaneously monitor the real-time change in cerebral NO fluctuation in the rat brain cortex following electrical stimulation.98 Lee et al. reported a NO sensor, in which single crystalline RuO2 nanorods were grown on a single CFE via thermal annealing. The RuO2 nanorod-modified CFE showed high catalytic activity toward NO oxidation, with its amperometric response to NO concentration being ∼40-fold higher than that of bare CFE. In this manner, they were able to monitor real-time NO fluctuation upon the administration of Larginine in the rat brain.99 Another solution is the permselective membrane that prevents transport of interfering species to electrode surfaces. Lowry et al. developed a Pt sensor coated by thermally annealed Nafion for real-time monitoring of NO at a potential of +0.9 V (vs SCE).100 Major interferences from nitrite and ascorbate were prohibited on the probe, of which the selectivity was further improved by cocoating Nafion and o-phenylenediamine (o-PD). The resultant microelectrodes exhibited very high selectivity ratios for NO against ascorbate (>30 000:1) and nitrite (>2 000:1). Using the double-layer-coated microelectrodes, Laranjinha et al. studied the relationship between NO and ascorbate and its functional effect on glutamatergic activation.101 Fluorinated Xerogel is also an attractive permselective membrane to the fabrication of NO microsensors by a simple and reliable dip-coating procedure.102,103 Very recently, Suh et al. reported a dual-functional microsensor with fast response for in vivo monitoring of NO/CO. Fluorinated xerogel was coated on the sensing probe to prevent common physiological interferences from approaching the electrode. With this sensor, they simultaneously monitored NO and CO changes in the rat brain during seizure conditions induced by 4-aminopyridine in cortex, and the results demonstrated that NO and CO changes were related to different periods of corresponding seizure local field potential signal changes.104 Later again, they integrated a potassium selective microelectrode into the NO sensor for simultaneous, real-time monitoring of NO and K+ changes in the rat brain during spontaneous neocortical epileptic seizure.105 Hydrogen Peroxide. Being one of the reactive oxygen species and a diffusible messenger like NO, H2O2 occupies a crucial position in intracellular signaling, brain function, and many neurodegeneration disorders like Alzheimer’s disease and Parkinson’s disease.106,107 Thus, a specific and less technically demanding method for durable measurement of H2O2 in vivo would be of great biological implications to fundamental neurochemistry. Although H2O2 is electrochemically detectable in principle, high overpotentials for the oxidation and reduction of H2O2 have laid a huge obstacle in front of in vivo electrochemical measurements of H2O2.108 Even worse, most of the catalytic sensors employing noble metals (such as Pt) reported so far could not be applied for in vivo H2O2 sensing due to the interference from dissolved O2, ascorbate, and catecholamine.109 A smart strategy was proposed by Lowry et al.,

DOPAC. Moreover, Lin et al. immobilized laccase into a fusedsilica capillary to form a magnetic microreactor integrated in an OECS coupled to in vivo microdialysis for selective detection of DA in the microdialysate sampled from the brain of a freely moving rat.81 Linear current response toward DA concentration was obtained from 1 to 20 μM together with an excellent sensitivity of 3.97 nA/μM. O2. Being essential to almost all living biosystems, O2 reduction is closely linked to energy production/consumption and oxidative stress in vivo, hence the concentration level of O2 is a useful indicator of neuron activities.82 On electrode surfaces, O2 reduction may be realized through different pathways, depending on the nature of electrode materials. To be specific, O2 can be completely reduced to water on the platinum (Pt) surface through a four-electron process, but on other materials like carbon and gold O2 reduction is split into two two-electron processes, generating the intermediating H2O2.83−85 For example, Bolger et al. utilized the carbon paste electrode for in vivo monitoring of O2.86 As the electrode was polarized at a potential of −650 mV (vs SCE), they observed the two-electron reduction of O2 with H2O2 as the intermediate. In this way, they studied the alleviation of hypoglycaemia by hippocampal O2 and glucose metabolism. Different reduction outcomes are ascribed to the faster reduction reaction kinetics on platinum as compared to that on carbon or gold. Therefore, most of in vivo O2 probes were made with Pt. However, implanting a Pt microwire directly into the brain is not feasible, as it must be sealed by insulating coatings and turned into a microdisk electrode with a dramatically increased diameter (typically in the millimeter range), such as the Clark-type O2 electrode.87 To solve this problem, Pt nanoparticles were electrodeposited on the microelectrode to enhance catalytic O2 reduction.88 Recently, Xiang et al. reported the fabrication of CFEs sheathed by platinized vertically aligned CNTs (VACNT).89 VACNTs were grown on CFEs by pyrolysis of iron phthalocyanine (FePc) to produce VACNT-coated CFEs. Pt nanoparticles were electrochemically deposited on the surface of VACNT-CFEs. The resultant Pt/VACNT-CFEs exhibited higher sensitivity and better stability for in vivo O2 sensing as compared to traditional Pt/CFEs. With these electrodes, an electrochemical sensor was constructed to monitor O 2 fluctuation in live brain of rats in the early stage of global cerebral ischemia/reperfusion, exposed to O2 and N2 for a short time or hindfeet-pinched (Figure 3). NO. It is now widely accepted that nitric oxide (NO) plays an important role in vasodilation, blood pressure, immune response, and neural communication in the CNS as the gasotransmitter.90 Furchgott and Ignarro were awarded the Nobel Prize in 1998 for the discovery of endothelial derived relaxation factor (EDRF) function of NO. Up to date, quantum cascade laser, electrochemical sensors, electron spin resonance (ESR), and fluorescent probes have been developed to measure NO in various physiological and pathological processes. Among these, electrochemical NO sensors are more beneficial to in vivo application because of the high spatial/temporal resolution.91 On one hand, NO can be reduced to N2O22− at negative potentials but a reductive NO sensor suffers the interference from O2 reduction. On the other hand, in vivo electrochemical NO sensors based on NO oxidation also risk the interference from nitrite oxidation at similar potential.92 To overcome the challenge, electrocatalysts, such as metalloporphyrins, metalF

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Figure 4. (A) Scheme of microbiosensor implanted in rat brain and structure of PDA/PB/CNT/CFE. (B) Amperometric response recorded at the PDA/PB/CNT/CFE in the cortex of anesthetized rats during local microinfusion of 1 mM MCS and 10 mM MCS as indicated. Reproduced from Li, R.; Liu, X.; Qiu, W.; Zhang, M. Anal. Chem. 2016, 88, 7769−7776 (ref 115). Copyright 2016 American Chemical Society.

exhibited highly efficient ECL, which was strongly quenched by Cu2+ through an excited-state electron transfer or energy transfer mechanism and recovered by H2S. Accordingly, the [Ru(bpy)2(bpy-DPA)]2+/TPA composite complexed with Cu2+ was confined within a Nafion film on the GC electrode sealed in a sample vial to analyze H2S volatilized from the cerebral microdialysate.127 It was an impressive attempt since the basal level of H2S in the microdialysate from the cortex of adult male Sprague−Dawley rats was determined to be around 2.3 ± 0.9 μM (n = 4).

who immobilized catalase on one of the paired Pt electrodes to degrade H2O2 into water and O2 on one side.110 When used in vivo, the paired electrodes were implanted in close proximity with respective current simultaneously recorded. Information on concentration was thus extracted by monitoring the current difference. Similar to natural enzymes, Prussian blue (PB) can catalyze H2O2 reduction as an “artificial enzyme” for selective H2O2 sensing;111,112 however, interactions between Fe3+ and OH− at neutral pH cause collapse of the framework of PB and loss of the electrocatalytic activity.113,114 To solve this problem, Li et al. stabilized PB within polydopamine (PDA) under the physiological condition and enabled the PB−PDA-based sensor to monitor in vivo H2O2 changes during local microinfusion of H2O2 and mercaptosuccinate (MCS) (Figure 4).115 Hydrogen Sulfide. In the CNS, hydrogen sulfide (H2S) not only acts as a gasotransmitter and a neuroprotectant via antioxidant, anti-inflammatory, and antiapoptotic effects but also exerts potential therapeutic effects on several CNS diseases including Alzheimer’s disease, Parkinson’s disease, ischemic stroke, and traumatic brain injury.116,117 Therefore, increasing research interest has been attracted to the in vivo sensing of H2S. H2S is a volatile weak acid existing in forms of HS− (80%), H2S (20%) and S2− (trace) under physiological conditions. H2S can be electrochemically oxidized into sulfur but at a very high potential (e.g., +1.5 V at Pt electrode).118,119 Although several metallic electrodes (Ni, V2O5, etc.) can significantly lower the oxidative potential, they generally require an acidic or alkaline media that is distant from in vivo environment.120,121 Extreme pH is also necessitated by potentiometric methods based on silver/silver sulfide ion selective electrodes or ionophores.122,123 In vivo selective measurement of H2S could be carried out with a commercial Clark-type O2 microsensor.124,125 In this case, dissolved H2S passes through a H2S-permeable membrane into the alkaline solution (pH < 9), ensuring complete dissociation of H2S into HS−. Oxidation of HS− with an oxidizing reagent (i.e., ferricyanide) results in the formation of sulfur and ferrocyanide, generating a current signal corresponding to H2S concentration.126 Electrochemiluminescence (ECL) has lately become a powerful tool for H2S sensing in combination of in vivo microdialysis. In the recent work by Yue et al., a synthetic composite of a Ru(II)-bipyridine complex bearing a di(2picolyl)amine (DPA) moiety and tri-n-propylamine (TPA)

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ELECTROCHEMICAL BIOSENSORS Oxidase-Based Electrochemical Biosensors. Electrochemical biosensors relying on catalysis by oxidases have been designed toward direct electron transfer (DET) or mediated electron transfer (MET) at electrode-oxidase interface (Figure 5). DET-based biosensors are technically simple to

Figure 5. Schematic illustration of typical electronic transfer processes for the first generation (A), second generation (B), third generation (C) oxidase-based biosensors, and dehydrogenase-based biosensors (D).

fabricate, but they have frequently encountered inefficient electron transfer between electrode surface and the FAD active site deeply buried inside the large insulating protein body,128,129 not to mention that only properly orientated oxidases on electrode are capable of DET, resulting in much lower current response than expected.130 Another drawback of DET-based oxidase biosensors is the O2-dependency since O2 acts as the G

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showed a steady performance when pH fluctuated from 6.5 to 8.5, and displayed a linear sensitivity and good selectivity.151 The self-referencing technique was employed by Gerhardt et al. to achieve desired selectivity of oxidase-based biosensors utilizing two microelectrodes on a multisite array.152,153 One recording site was coated with Nafion film containing Lglutamate oxidase and bovine serum albumin (BSA) crosslinked by glutaraldehyde, while the other was coated with same Nafion film and BSA cross-linked with glutaraldehyde without L-glutamate oxidase. Glutamate was thus selectively measured by the difference between currents recorded at both recording sites. This method was successfully used for in vivo monitoring of the clearance of locally applied glutamate and release of glutamate in the prefrontal cortex of anesthetized rats. O2 can be replaced by artificial mediators to improve the sensitivity and selectivity of oxidase-based biosensors. In this way, polarization potentials of the biosensors were mostly dependent on the redox potentials of artificial mediators used and thus the measurements could be carried out at much lower potentials, as compared to the first-generation oxidase-based biosensors mentioned above. So far, a number of organic compounds, such as ferrocene derivatives, ferricyanide, conducting organic salts, quinone compounds, transitionmetal complexes have been used as electron transfer mediators for oxidases.154−157 For instance, Boutelle et al. used TTFTCNQ as the mediator for glucose oxidase in preparing a glucose biosensor for continuously monitoring glucose level in the brain without interference from ascorbate or other electroactive species.158 Even so, it is not always true that artificial mediators are wellsuited for continuous monitoring of neurochemicals in vivo, if their formal potentials fall in the oxidation window of ascorbate, uric acid, catecholamine, and related metabolites in the CNS,159 or if they are not stable on electrode surfaces. One example is ferricyanide (Fe(CN)63−), which has been popular in METbased biosensors because of its good electrochemical property and reactivity with the FAD-containing oxidases.160−162 However, high formal potential of Fe(CN)63−/Fe (CN)64− in solution and easy leaching from electrode surfaces hindered the in vivo applications of Fe(CN)63−-based biosensors.163−165 In this regard, attempts to negatively shift the redox potential of the mediators and improve their surface stability are very essential for developing in vivo MET-based electrochemical biosensors. Zhuang et al. recently developed a ferricyanide-incorporated oxidase-based biosensor capable of selective in vivo measurements by rationally controlling different adsorption ability of Fe(CN)63− and Fe(CN)64− onto the imidazolium-based polymer (Pim) modified electrode surface.132 Strong interactions of Pim with Fe(CN)63− than with Fe(CN)64− resulted in stable surface adsorption of both species and, more importantly, a negative shift of the redox potential of surfaceconfined Fe(CN)63−/Fe (CN)64−. New low-potential mediators were also explored for in vivo purposes. Zhang et al. developed a xanthine oxidase (XOD)based OECS, in which thionine was adsorbed onto CNTs and used to shuttle electrons between XOD and CNT-modified electrodes.131 The biosensor was demonstrated to be highly selective against endogenous species during continuous monitoring of xanthine in the brain. Enzyme immobilization on electrode surfaces is a key factor determining the performance of enzymatic electrochemical biosensors. Vasylieva et al. introduced a new enzyme

natural electron acceptor when electrons from substrates cannot be efficiently transferred to the electrode. MET-based biosensors normally use additional electron transfer mediators to shuttle electrons between electrode and distant FAD active sites, yielding much higher current response than DET-based biosensors do. As a matter of fact, MET-based biosensors have gained more research attention and effort from neurochemists for their in vivo applications in amperometric, real-time neurochemical biosensing and continuous monitoring of neurochemicals in brain microdialysate with an OECS.131−133 Recent years have witnessed a rapid development of MET-based in vivo electrochemical biosensors, as summarized below. The first-generation biosensors utilized O2 as the natural electron acceptor for oxidases and sensing the substrates of interest in terms of measuring H2O2 generated from the oxidase-catalyzed reactions. Relatively high potentials (e.g., +0.70 V vs Ag/AgCl) for H2O2 oxidation readily allowed cooxidation of endogenously existing ascorbate, UA, DA, and other metabolites.134,135 Thus, poor selectivity was the biggest hurdle for in vivo biosensing. To improve the selectivity of the first-generation biosensors, some permselective membranes such as Nafion, cellulose acetate, polypyrrole, poly(phenylenediamine), and polyaniline were exploited.136−140 Chen et al. suppressed the interference in the brain by electropolymerizing phenol to form a protective polyphenol film on the electrode.141 Baker et al. developed glucose oxidase-modified poly(phenylenediamine)-coated electrodes to avoid the interference from ascorbate.142 Recently, Rocchitta et al. described a new approach for simultaneous detection of brain glucose and lactate in a real-time manner by using a Pt/Ir cylinder oxidase-based electrode coated with poly(o-phenylenediamine).143 Combined with telemetric technique, the electrode was used to monitor glucose and lactate in the brain of freely moving rats. Biosensing selectivity was also improved by using horseradish peroxidase (HRP) as a biorecognition unit for H2O2 at a relatively lower potential.144 Garguilo et al. constructed a choline sensor by entrapping choline oxidase (ChOx) and HRP within a cross-linkable redox polymer on CFEs coated with Nafion to eliminate unwanted currents.145 However, oxidation of ascorbate by H2O2 could compete with the catalytic reduction of H2O2. To address this issue, AOx was added to consume ascorbate in the sample.146,147 Lin et al. demonstrated an OECS for simultaneous measurements of glucose and lactate in microdialysates sampled from the brains of freely moving rats.148 Selective electrochemical detection was accomplished by using glucose oxidases and lactate oxidases as the specific sensing units together with PB as the electrocatalyst for H2O2 reduction. Less interference from ascorbate and other electroactive species as well as O2 was observed. The dual-functional oxidase/PB-based biosensors in the OECS risked no crosstalk and exhibited good stability and reproducibility for in vivo analysis. To stabilize PB framework at neutral pH,149 Li et al. recently incorporated PB into polydopamine (PDA) to form a stable polymer that was used for in vivo monitoring of glucose by coupling with glucose oxidase.150 Another glucose biosensor built with glucose oxidase and PB was developed by Salazar et al. by using poly o-phenylenediamine (PoPD) as the PB stabilizer and polyethylenimine (PEI) to fix enzyme conformation through electrostatic interactions. This biosensor H

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Figure 6. (A) Schematic illustration of the OECS for in vivo monitoring of glucose in the cerebral system. (B) Online current−time response for glucose standards with concentrations indicated in the figure recorded on the OECS with a ZIF-70-based biosensor as the detector. Pure aCSF or aCSF containing standards was perfused into the system from pump 1 (P1) at 2 μL/min, and aCSF containing 5 mM NAD+ was perfused from pump 2 (P2) at a rate of 2 μL/min. The biosensor was polarized at 0 V. Inset, schematic illustration of the ZIF-70-based biosensor. (C) Online current−time response recorded for the brain microdialysates of guinea pig on the online detecting systems with the MG/ZIF-70-modified GC electrode (i.e., GDH-free ZIF-70-based sensor, black line) and ZIF-70-based biosensor (red and blue line) as the detector. The brain microdialysate was continuously sampled from pump 1 (P1) and online mixed with pure NAD+ solution (5.0 mM, aSCF) from pump 2 (P2). The perfusion rates for both pumps were 2 μL/min. The electrodes were polarized at 0 V. Reproduced from Ma, W.; Jiang, Q.; Yu, P.; Yang, L.; Mao, L. Anal. Chem. 2013, 85, 7550−7557 (ref 188). Copyright 2013 American Chemical Society.

studies indicated that enzyme loading, other than the electrode geometry, determined the current density of glutamate oxidation and O 2 -dependence. 174 In other words, the biosensors with higher glutamate oxidase loading would give rise to higher current density and lower O2 dependence. They also noted that whether the electrode design suited an application depended on the concentration of glutamate being monitored as well as the range of fluctuations in PO2 in the medium. For example, excessive glutamate and low PO2 in anoxia could undermine the function of the glutamate biosensor. Even though the oxidase-based biosensors with artificial electron transfer mediators are less O2-dependent, O2 influence cannot be completely excluded due to much faster electron transfer between the enzyme and O2. To minimize O2dependence, Ward et al. utilized an outer permselective membrane made of amphiphobic polyurethane that only allowed glucose passage through the hydrophilic segments.175 Chen et al. presented the manufacture of microneedles coated by nanomembranes composed of 3D porous polyvinylidene fluoride (PVDF) and nanosphere Nafion to rule out disturbance from neurochemicals and O2 during in vivo glucose sensing.176 O2-releasing materials balancing local O2 concentration was also developed to fabricate biosensors with improved tolerance of O2. Njagi et al. applied cerium oxides, which could store and release O2, into a glutamate oxidasebased biosensor and demonstrated its capability of sensing glutamate under oxygen-depleted conditions.177 Sardesai et al. took one more step forward and developed the cerium oxidesmodified biosensor into an O2-tolerant, dual-functional, in vivo lactate and glutamate biosensors.178,179 Dehydrogenase-Based Electrochemical Biosensors. Dehydrogenase-based electrochemical biosensors utilize dehydrogenases to catalyze the oxidation of analytes with the aid of cofactors (e.g., NAD+), as illustrated in Figure 5 D. Unlike oxidase-based electrochemical biosensors, dehydrogenase-based

immobilization method using poly(ethylene glycol) diglycidyl ether (PEGDE) as a mild cross-linking reagent.166 Comparing five different enzyme immobilization methods including glutaraldehyde or PEGDE cross-linking, hydrogel matrix attachment, sol−gel, or derived polypyrrole entrapment, they demonstrated that enzyme immobilization could significantly impact the enzyme activity on the electrode. For example, cross-linking by glutaraldehyde other than PEGDE markedly decreased the apparent substrate affinity of glutamate oxidase.167 However, Burmeister et al. found that cross-linking of glutamate oxidase by glutaraldehyde still retained sufficient enzyme specificity for in vivo measurements of tonic and phasic glutamate levels when particular “wetting” procedures was involved in the immobilization processes.168 Cheng et al. recently demonstrated that rational integration of enzymes within nanoscale metal-organic framework (MOF) matrixes enhanced bioelectrocatalytic activities and stability as well.169 Wang et al. reported a simple one-step formation of bioelectrochemically multifunctional film (BMF), in which oxidase, ferrocene (Fc) as the electron transfer mediator and graphene oxide were homogeneously dispersed in a polymaleimidostyrene and polystyrene matrix on the electrode.170 The as-prepared biosensors were well responsive to glucose concentration and successfully used to monitor the change of glucose level in the microdialysate continuously sampled from the auditory cortex of the rat brain during the salicylate-induced tinnitus model with a high reliability and robustness. O2-dependence remains an inevitable challenge to the oxidase-based electrochemical biosensors for in vivo applications, considering the fact that large fluctuation in brain O2 level may occur in conditions like spreading depression and ischemia.171,172 McMahon et al. compared O2-dependence of glutamate oxidase-based electrochemical biosensors built on Pt cylinder and Pt disk.173 They found that the disk biosensor gave higher current density for glutamate and showed lower O2dependence, as compared with the cylinder design. In-depth I

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cascade to be electrochemically detectable and thus multienzyme-based electrochemical biosensors emerged.193,194 For in vivo monitoring of acetylcholine, acetylcholine oxidase (AChE), and choline oxidase (ChOx) can be utilized to catalyze conversion of acetylcholine into betaine and H2O2, with choline being the intermediator.

electrochemical biosensors are O2-independent and work at relatively lower potentials.180,181 However, concurrent usage of cofactors (e.g., oxidized form of nicotinamide adenine dinucleotide, NAD+) and electrocatalysts to catalyze regeneration of cofactors (e.g., reduced form of nicotinamide adenine dinucleotide, NADH) is required to accomplish bioelectrocatalysis and transduce biorecognition events into electrical current signals in dehydrogenase-based biosensors.182 Hence, much fewer in vivo applications of this type of biosensors have been reported so far because of lower catalytic activity of dehydrogenases and involvement of cofactors as well as difficult integration of all biosensing elements into one device.183 Lin et al. recently developed a dual-functional dehydrogenase-based OECS for in vivo monitoring of glucose and lactate with an external solution containing NAD+ cofactor pumped in and mixed online with the brain microdialysates.184 The OECS exhibited high selectivity toward glucose and lactate sensing and good tolerance against the changes of extracellular O2 and pH in the rat brain following the global cerebral ischemia/ reperfusion. Stepping further, they integrated the OECS into a microfluidic chip for continuous and simultaneous monitoring of glucose, lactate, and ascorbate in the rat brain.185 The trifunctional OECS displayed a good linear response toward concentrations of glucose, lactate, and ascorbate with little crosstalk, high stability, and high selectivity. As shown in Figure 5D, an NAD+-dependent dehydrogenasebased bioelectrocatalytic scheme is accomplished by dehydrogenases, NAD+, and electrocatalyst for NADH oxidation, so the practical fabrication of such a biosensor calls for step-by-step surface modification procedures that are usually technically complicated and time-consuming, especially when miniaturization is considered.186,187 These drawbacks have limited the in vivo applications of dehydrogenase-based electrochemical biosensors. In up-to-date progress, Ma et al. used zeolitic imidazolate framework (ZIF) for the first time as the matrix for coimmobilization of the electrocatalyst methylene green (MG) and glucose dehydrogenase (GDH) in a GDH-based electrochemical biosensor (Figure 6).188 They found that the unique properties of ZIF like diverse porosity and internal functional groups substantially facilitated the adsorption of GDH and MG. This study broadened the horizon for the application of metal−organic frameworks in developing in vivo electrochemical biosensors. Fabrication of dehydrogenase-based electrochemical biosensors was further simplified by Huang et al. by integrating all biosensing elements into a coordination polymer to form “all-in-one” bioelectrochemically active infinite coordination polymer (ICP) nanoparticles deposited on electrode surfaces. In the one-pot synthesis under mild conditions, addition of an aqueous Tb(NO3)3 solution into the buffer containing NAD+, MG, and GDH induced rapid self-assembly and coprecipitation.189 Terbium ions (Tb3+) were chosen as the metal connector because of its flexible coordination and electrochemically inactive nature. NAD+ served as the bidendate ligand to Tb3+ to form a Tb-NAD+ network with all functional molecules encapsulated through π−π stacking and hydrogen bonding. In a follow-up study, Lu et al. combined ICPs and SWNTs to obtain a three-dimensional conductive matrix for highly sensitive biosensing of neurochemicals in vivo.190 Multienzyme-Based Electrochemical Biosensors. In certain circumstances, electroinactive neurochemicals, such as acetylcholine and ATP,191,192 have to go through an enzyme

AChE

ACh + H 2O ⎯⎯⎯⎯⎯→ Ch + CH3COOH ChOx

Ch + O2 + 2H 2O ⎯⎯⎯⎯⎯→ betaine + 2H 2O2 Cat

2H 2O2 ⎯→ ⎯ 2H 2O + O2

In the CNS, the concentration of choline is much higher than that of acetylcholine, and both of them react at the AChEChOx-modified electrode, which brings in new challenges for in vivo selective detection of acetylcholine. In the OECS for continuous acetylcholine sensing developed by Kato et al., choline was enzymatically removed in a small-volume prereactor packed with ChOx and catalase.195 Lin et al. also used a renewable multienzyme microreactor containing immobilized ChOx and catalase to eliminate choline from the OECS, which was shown to be selective, reproducible, and stable for the continuous measurement of acetylcholine in vivo.196 Burmeister et al. designed a ceramic-based multisite microelectrode array for simultaneous measurements of choline and acetylcholine in the CNS, with one recording site modified with ChOx and the other with AChE and ChOx.197 In consequence, the concentration of choline was directly measured at the ChOx-modified recording site, while the concentration of acetylcholine was derived from the current difference between two recording sites. In vivo applications of multienzyme-based electrochemical biosensors were extended to the monitoring of adenosine and ATP. Adenosine is a neuromodulator regulating heart rate, sleep, and breathing. ATP is the energy supplier that also participates in signal transduction.198 Llaudet et al. created a three-enzyme cascade of nucleoise phosphorylase, xanthine oxidase, and adenosine deaminase to convert adenosine into the reporter molecule (i.e., H2O2).199 By doing so, they developed a small, fast-responding, sensitive, and stable electrochemical biosensor for in vivo monitoring of adenosine release during hypoxia of rat hippocampal slices from the CA1 region. They also developed an ATP biosensor by coating an ultrathin biolayer containing glycerol kinase and glycerol-3phosphate oxidase onto a Pt microelectrode, which responded rapidly and exhibited a wide linear response range toward ATP concentration.200 With this biosensor, they discovered that ATP is a mediator in chemosensory transduction in the CNS. Hinzman et al. fabricated an enzyme-linked microelectrode array to selectively measure extracellular adenosine by the selfreferencing technique. LOD was leveled down to ∼0.04 μM, thus validating the feasibility of the multienzyme array in realtime monitoring of the basal extracellular concentration of adenosine in rat cerebral cortex.201 Aptamer-Based Biosensor for ATP. In addition to multienzyme-based electrochemical biosensors, many ATP biosensors based on aptamers have been developed over the past 2 decades, most of them being not suitable for selective ATP analysis in the CNS. Although aptamers mainly recognize the A nucleobase, weaker but nontrivial interactions with the J

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(bio)sensors onto electrode arrays will facilitate brain mapping study of biological molecules. Development of the nonpolarized electrochemical principles and techniques would be imperative since the traditional electrochemical sensor arrays are limited in resolution due to overlap of diffusion layers. Finally, electrochemical (bio)sensors should be simply and reproducibly fabricated so that they can be easily prepared by nonelectrochemists. In summary, although in vivo recording of chemical signals with electrochemical (bio)sensors remains a challenging area, we still believe that the electrochemical (bio)sensors hold a great promise for in vivo analysis by fully taking advantage of the development of chemistry and electrochemistry as well as biochemistry and by well benefiting the interface of chemistry, electric engineering, materials science, and physiology.

sugar and triphosphate moieties endow aptamers with similar affinities to ATP, ADP, and AMP.202,203 To improve selectivity to ATP sensing, Yu et al. recently reported a dual recognition unit strategy (DRUS) to construct a highly selective and sensitive ATP biosensor.204 Taking the advantage of the recognition ability of aptamers toward A nucleobase and that of polyimidazolium toward phosphates, the biosensor based on DRUS not only showed an ultrahigh sensitivity toward ATP with the LOD down to the subattomole level but also an ultrahigh selectivity against interference from ADP and AMP in sensing extracellular ATP in microdialysates sampled from the cerebral system.

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CONCLUSION In vivo analysis based on electrochemical sensors and biosensors remains a hot topic in the interdisciplinary research fields spanning chemistry, neuroscience, physics, and material science, because it is of great implications for understanding of the molecular essence of brain functions. As reviewed herein, the uses of electrochemical sensors and biosensors have largely enabled in vivo analysis by combining a rationally designed electrode/solution interface with thorough understanding of the demand in brain chemistry. However, it is fair to say that in vivo measurement of physiologically important molecules continues to be a great challenge; high sensitivity and selectivity, good spatial and temporal resolution temporal, and excellent antifouling property are all essential issues to be carefully addressed. For electrochemical sensors, rationally designing the electrode surface to be analyte-specific and accurately controlling electrochemical parameters will potentially enable the sensors to in vivo monitor the neurochemicals in the CNS. In addition, introduction of new electrochemical principles and techniques (e.g., nanopore and bipolar) may open a new window for in vivo analysis with electrochemical sensors. For electrochemical biosensors, oxidase-based biosensors have been well developed for in vivo applications although they are yet confronted with interference from oxygen fluctuation. Dehydrogenase-based biosensors are oxygen-independent but require the uses of cofactors for biosensing. Compared with the oxidase-based biosensors, dehydrogenase-based biosensors have been fairly used for in vivo monitoring of neurochemicals. To enable dehydrogenase-based biosensors for in vivo analysis, several issues such as stable immobilization of cofactor (i.e., NAD+) and improvement of enzyme activity should be properly addressed. The optimization of the enzyme structure by gene/protein engineering would be an effective solution to these problems. In addition to the sensors based on enzymes, the uses of other recognition units would always be a prominent option for vivo analysis. Still, we have to say that there are numerous opportunities to apply electrochemical (bio)sensors as analytical tools for physiological and pathological studies, in particular by combining these electrochemical methods with other techniques. First, combination of electrochemical methods with physiological techniques, such as electrophysiology and/or optogenetics, would become very attractive for physiological studies. However, crosstalk between electrochemical signal and electrophysiological signal will be an inevitable problem. New electrochemical principles will be definitely needed to solve this problem. Second, merging of wireless technique into electrochemical (bio)sensing methods would drive convincible behavior-related studies. Third, integration of electrochemical

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected]. Fax: (86)-10-62559373. ORCID

Lanqun Mao: 0000-0001-8286-9321 Notes

The authors declare no competing financial interest. Biographies Tongfang Xiao is a Ph.D. candidate in the Key Laboratory of Analytical Chemistry for Living Biosystems at Institute of Chemistry, the Chinese Academy of Sciences (ICCAS). He received his B.A. in chemistry from Zhengzhou University in 2013. He then joined ICCAS to pursue his Ph.D. degree under the guidance of Lanqun Mao. His current research focuses on in vivo electrochemical (bio) sensors. Fei Wu is a postdoc in the Key Laboratory of Analytical Chemistry for Living Biosystems at ICCAS. She received her B.A. in pharmaceutical sciences from Nankai University in 2009. She finished her Ph.D. training in chemistry at the University of Utah in 2015. She is now working on bioelectrochemistry and bioelectrocatalysis. Jie Hao is a Ph.D. candidate in the Key Laboratory of Analytical Chemistry for Living Biosystems at ICCAS. She received her B.A. in Chemistry from Hebei University in 2011. Her current research focuses on in vivo electrochemical sensors. Meining Zhang is an associate professor in the department of chemistry at Renmin University of China. She received her B.Sc. and M.Sc. in chemistry from Shaanxi Normal University in 2001 and 2003, respectively. She then finished her Ph.D. training with Lanqun Mao at the Institute of Chemistry at the Chinese Academy of Sciences in 2006. She has been working on the development of multiscale methods for probing neurochemistry including in vivo electrochemistry and electrochemiluminescence. She is a recipient of the “National Excellent Young Scholars” from Natural Science Foundation of China (2015). Ping Yu is a professor of chemistry in the Key Laboratory of Analytical Chemistry for Living Biosystems at ICCAS. She received her Ph.D. in chemistry from ICCAS in 2007. Her ongoing work focuses on nanoelectrochemistry and chem/(bio)sensors. She is a recipient of the “National excellent Young Scholars” from Natural Science Foundation of China (2013). Lanqun Mao is a professor in the Key Laboratory of Analytical Chemistry for Living Biosystems at ICCAS. He has been working on the interface between electrochemistry and neurosciences, with emphasis on the development of new neurotechnologies to understand the chemical signals involved in some physiological processes. He obtained his Ph.D. from East China Normal University in 1998 K

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and then worked in BAS Inc. Japan as a research scientist (1998− 2000) and pursued his postdoctoral studies at the Department of Electronic Chemistry at Tokyo Institute of Technology (2000−2002). He was a recipient of the “Hundred Distinguished Young Scholars” from the Chinese Academy of Sciences (2002) and the “National Distinguished Young Scholars” from the National Natural Science Foundation of China (2006).

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ACKNOWLEDGMENTS

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REFERENCES

This work was financially supported by NSF of China (Grants 21321003, 21435007, and 21210007 for L. Mao, Grants 21322503 and 21475138 for P. Yu), and National Key Research and Development Project of China (Grants 2016YFA0200104 and 2013CB933704), and the Chinese Academy of Sciences.

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DOI: 10.1021/acs.analchem.6b04308 Anal. Chem. XXXX, XXX, XXX−XXX