Stability and Sensitivity Enhanced Electrochemical In Vivo Superoxide

Jun 25, 2012 - Kyubin Shim , Jeonghun Kim , Mohammed Shahabuddin , Yusuke ... A. K. M. Kafi , M. M. Yusoff , Mohammad Choucair , Maxwell J. Crossley...
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Stability and Sensitivity Enhanced Electrochemical In Vivo Superoxide Microbiosensor Based on Covalently Co-immobilized Lipid and Cytochrome c Md. Aminur Rahman,‡ Anitha Kothalam,† Eun Sang Choe,§ Mi-Sook Won,∥ and Yoon-Bo Shim*,† †

Department of Chemistry and Institute of BioPhysio Sensor Technology, Pusan National University, Pusan 609-735, South Korea Graduate School of Analytical Science and Technology (GRAST), Chungnam National University, Daejeon 305-764, South Korea § Department of Biological Sciences, Pusan National University, Pusan 609-735, South Korea ∥ Korea Basic Science Institute, Pusan 609-735, South Korea ‡

ABSTRACT: Enhanced stability and sensitivity of a superoxide anion radical (O2•−) microbiosensor were achieved through the sequential immobilization of lipid and cytochrome c (Cyt c) covalently bonded onto a conducting polymer layer that showed a clear quasi-reversible direct electron transfer (DET) process. The formal potential and the apparent standard rate constant were determined to be −0.24 V and 0.62 ± 0.05 s−1, respectively. The detection of O2•− was attained through the catalytic activity of the haem group of Cyt c stabilized by coimmobilized lipid molecules (1,2-dipalmitoylsn-glycero-3-phosphoethanolamine-n-dodecanylamine (DGPD)). The linear dynamic range and the detection limit of the O2•− analysis were determined to be 0.2−6.0 nM and 30.0 ± 0.9 pM, respectively. The in vivo microbiosensor implanted into rat brain successfully determined the extracellular level of O2•− produced by acute and repeated injections of cocaine. The present O2•− microbiosensor could be an effective tool for monitoring the change in extracellular O2•− levels in response to stimulant drug exposure. uperoxide anion radical (O2•−), the primary species of the so-called reactive oxygen species (ROS), is generated in significant quantities as a result of normal intracellular metabolism.1,2 Although, the concentration of O2•− in normal physiological condition is remarkably low (approximately nanomolar concentrations) as it undergoes disproportionation by an enzymatic reaction of superoxide dismutase (SOD),3 its concentration increases (up to approximately micromolar concentrations) because of brain injury, stresses, and ischemic-reperfusion.4 The O2•− has been implicated in a variety of pathological conditions, including rheumatoid arthritis, neurodegenerative disorders, aging problem, inflammatory diseases, atherosclerosis, and cancers.5 O2•− is also of considerable importance in the determination of cellular damage and thus, a specific and sensitive method for durable and reliable measurements would facilitate the investigations on the pathology and physiology of diseases. Thus, real-time monitoring of the O2•− in living cells and tissues require a sensitive and stable detection technique. As a sensitive and stable detection technique for O2•−, electrochemical biosensors6−12 have attracted growing attention because of their simple instrumentation, high sensitivity, and possibility to be miniaturized. In most cases, a metalloprotein, cytochrome c (Cyt c) has been used for the fabrication of electrochemical O2•− biosensors.8−12 Its direct

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© 2012 American Chemical Society

electron transfer (DET) process was widely investigated among the redox proteins at platinum, gold, nickel, mercury, and ptype silicon electrodes.13−17 However, the DET process of Cyt c at platinum, gold, nickel, mercury, and p-type silicon electrodes had been shown to be slow due to adsorptioninduced denaturation. It is difficult to observe the DET process of Cyt c at a bare electrode, due to the inadequate orientation and denaturation. Thus, the bare electrode surfaces were modified with electron promoters such as conducting polymers (CP)9,18 and self-assembled monolayer (SAM) of organic molecules12,17 to observe the DET of Cyt c. However, electrode modified with CP or SAM does not provide any biological environment for the immobilized Cyt c. Therefore, the activity of the Cyt c might be decreased when immobilized onto the CP or SAM. The lipid based model biomembrane systems, which are quite similar to biomembranes in biological cells, can act as a suitable medium for the immobilization of redox proteins. Furthermore, lipid layers can significantly reduce interference by effectively excluding hydrophilic electroactive materials from the detecting surface. To mimic lipid-based model biomembranes, functional coupling of lipid constituents with solid Received: April 25, 2012 Accepted: June 25, 2012 Published: June 25, 2012 6654

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Figure 1. Schematic illustration of a poly-TTCA/DGPD/Cyt c based O2•− microbiosensor and the design of an Au microelectrode (100 μm).

surfaces has been studied and has become popular.19 To enhance the stability of lipid layers, the solid surfaces have been modified with the conducting polymers for bonding the lipid, and the electrochemical behavior of adsorbed Cyt c on the layers has been studied.9,10,18 However, the immobilization of Cyt c through the physical adsorption is not suitable and cannot make a stable and sensitive biosensor for the in vivo detection of O2•−. Thus, a stable immobilization of the Cyt c and lipid should be followed for the fabrication of a stable and sensitive in vivo biosensor. The coimmobilization of a lipid and Cyt c covalently bonded onto a CP film can make a stable O2•− biosensor probe because lipid and Cyt c can be immobilized through chemical bonding. Although, some electrochemical biosensors for the in vivo O2•− detection are reported,7,20 covalently coimmobilized lipid and Cyt c based in vivo O2•− microbiosensor has not yet been reported. In the present study, we have fabricated an in vivo O2•− microbiosensor by covalently coimmobilizing an amino lipid and Cyt c onto a conducting polymer (poly-5,2:5,2terthiophene-3-carboxylic acid21 (poly-TTCA/lipid/Cyt c) for monitoring the change in extracellular levels of O2•− in the rat brain by repeated administrations of an abused drug, cocaine. To construct the microbiosensor, we first made the poly-TTCA film on a needle type gold (Au) microelectrode,22 followed by coimmobilization of lipid and Cyt c on the Poly-TTCA film through covalent bonding. The experimental conditions of the O2•− detection were optimized, and the microbiosensor was calibrated before in vivo experiments. Finally, it was implanted into the rat brain for monitoring the concentration of O2•− release in the extracellular space stimulated by repeated cocaine administrations.

previous paper.23 Briefly, Cyt c from horse heart was converted to the fully oxidized form by the addition of excess K3Fe(CN)6 and then purified by ion-exchange chromatography on Whatman CM-32, eluted with 0.5 M NaCl + 0.01 M phosphate buffer solution at pH 7.0. Tetrabutylammonium perchlorate (electrochemical grade) was obtained from Fluka (U.S.A) and purified according the general method, followed by drying under vacuum at 10−3 Torr. A terthiophene monomer bearing a carboxylic group (5,2:5,2-terthiophene-3-carboxylic acid, TTCA) was synthesized according to a previous report.21 Glucose, ascorbic acid, uric acid, L-glutamate, L-cysteine, Lglutamine, dopamine, dihydroxyphenylacetic acid, choline, salicylic acid, hydrogen peroxide were obtained from Sigma Co. Tris-HCl buffer was used through out the experiments. Other chemicals were of analytical reagent grade. All aqueous solutions were prepared in doubly distilled water, which was obtained from a Milli-Q water purifying system (18 MΩ.cm). Instruments. Poly-TTCA/DGPD/Cyt c, Ag/AgCl (in saturated KCl), and a Pt wire were used as working, reference, and counter electrodes, respectively. Cyclic voltammograms (CVs) were recorded using a potentiostat/galvanostat, Kosentech model KST-P2 (South Korea). A quartz crystal microbalance (QCM) experiment was performed using a SEIKO EG&G model QCA 917 and a PAR model 263A potentiostat/galvanostat (U.S.A.). An Au working electrode (area: 0.196 cm2; 9 MHz; AT-cut quartz crystal) was used for the QCM experiment. Electron spectroscopy for chemical analysis (ESCA) was performed using a VG Scientific ESCALAB 250 XPS spectrometer with a monochromated Al Kα source with charge compensation (at KBSI, Busan). Impedance spectra were recorded with an EG&G PAR 273A Potentiostat/Galvanostat and a lock-in amplifier (PAR EG&G, Model 5210) linked to a personal computer. The frequency was scanned from 100 kHz to 10 Hz at the open circuit voltage, acquiring 5 points per decade. The amplitude of a sinusoidal voltage of 10 mV was used. Preparation of the Poly-TTCA/DGPD/Cyt c Microbiosensor. An Au microelectrode was made according to a previous report22 with a few slight modifications. Briefly, an Auwire (Au 99.999%, 100 μm in diameter, Johnson Matthey Inc., West Chester, PA, U.S.A.) of about 3 mm long connected with a copper wire was introduced into a micropipet and advanced

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EXPERIMENTAL SECTION Reagents. 1,2-Dipalmitoyl-sn-glycero-3-phosphoethanolamine-n-dodecanylamine (DGPD, molar mass 888.693 g) was purchased from Avanti Polar Lipids Inc., Alabaster, AL, and was used without purification. Horse heart cytochrome c (Cyt c, molar mass 11 kD), dichloromethane (99.8%, anhydrous, sealed under N2 gas), and 1-ethyl-3-(3-(dimethylamino)propyl) carbodiimide (EDC), polyethylenimine (PEI) were purchased from Sigma Co. (USA) and were used as received. The purification process of Cyt c has been described in our 6655

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reference/counter electrode was also inserted into the dorsal striatum through the same hole with a distance of 0.1 mm from the sensor tip. Possible acute inflammatory response caused by the implantation of the electrodes was verified by Nissl staining in which there was no glial staining along the sensor placements (data not shown). During this study, a 30 min time point was selected based on a time course of repeated cocaine injections in which O2•− levels were shown to peak at 30 min relative to control groups.

until it protruded about 1.0 mm from the tip of the micropipet. Then, the tip of the micropipet was heated gently in a butane flame until the glass collapsed and sealed the Au wire. The other end of the micropipet, from which the copper wire was introduced, was fixed with epoxy resin. The total length of the Au microelectrode was 35 mm as shown in Figure 1. The polyTTCA film was grown on the Au microelectrode through electropolymerization of the TTCA monomer in a 0.1 M TBAP/CH2Cl2 solution according to a previous report.22 After electropolymerization, the Poly-TTCA film coated Au microelectrode was washed with CH2Cl2 to remove any remaining monomers from the electrode surface. The Poly-TTCA coated Au microelectrode was immersed for 1 h in a mixed dichloromethane/acetonitrile (CH2Cl2/AN, 8:2) solution containing 10.0 mM of EDC/NHS and 1.0 mM DGPD lipid. After 1 h, DGPD lipid was partly immobilized on the PolyTTCA film through the amide bond formation between carboxylic acid groups of Poly-TTCA and amine groups of DGPD (Poly-TTCA/DGPD). Cyt c was then successively immobilized on the Poly-TTCA/DGPD modified Au microelectrode by incubating the modified electrode for 3 h in a 0.01 M PBS (pH 7.4) solution containing 10.0 mM Cyt c. By this step, Cyt c was covalently immobilized on the remaining empty sites of Poly-TTCA film through the amide bond formation between carboxylic acid groups of Poly-TTCA and amine groups of Cyt c (Poly-TTCA/DGPD/Cyt c). Finally, after blocking the modified electrode with 1.0% BSA solution, the modified microelectrode was coated with a cationic polymer polyethyleneimine (PEI) by dipping the modified electrode in a 1% PEI for three times to enhance the stability of the microbiosensor. The schematic illustration of the modified electrode is shown in Figure 1. Electrochemical Measurements. Cyclic voltammograms were recorded for Poly-TTCA/DGPD/Cyt c modified electrode from +0.10 to −0.70 V vs Ag/AgCl in a 0.1 M PBS of pH 7.4. Chronoamperometric experiments were carried out by applying the potential of −0.35 V. In in vivo experiments, two electrode configurations were used: a Poly-TTCA/DGPD/ Cyt c and an Ag/AgCl wire as working and reference/counter electrodes, respectively. All the microsensors were calibrated before and after the in vivo experiments. All calibrations were performed at 35 ± 1 °C. Animals. Adult male Sprague−Dawley rats (200−250 g) were obtained from Hyo-Chang Science Co. (Daegu, Korea). Rats were individually housed in a controlled envirionment during all experimental treatments. Food and water were provided ad libitum, and rats were maintained on a 12 h light/ dark cycle. On the day of in vivo experiment, cocaine injection was made in the quiet room to minimize stress to rats. All animal use procedures were approved by the Institutional Animal Care and Use Committee and were accomplished in accordance with the provisions of the NIH “Guide for the Care and Use of Laboratory Animals”. After each in vivo measurement, the rats were humanely killed. Cocaine Administration and Surgery. Rats received saline, acute (20 mg/kg), or repeated cocaine (20 mg/kg) injections for seven consecutive days intraperitoneally (i.p.). On the day of the experiment, rats were anesthetized with 8% chloral hydrate (6 mL/kg, i.p.) and placed in a stoelting stereotaxic apparatus 5 min before the time of measurement after final saline or cocaine injections. Under an aseptic condition, a O2•− microbiosensor was implanted unilaterally into the central part of the dorsal striatum. The micro-

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RESULTS AND DISCUSSION Preparation and Electrochemical Characterization of the Poly-TTCA Film. The Poly-TTCA film was formed on an Au microelectrode through the electropolymerization of a TTCA monomer in a 0.1 M TBAP/CH2Cl2 solution according to a previous report22 with a slight modification. Briefly, cyclic voltammograms (CV) recorded for an Au microelectrode in a 0.1 M TBAP/CH2Cl2 solution containing 1.0 mM TTCA monomer while the potential was cycled between 0.0 and +1.6 V versus Ag/AgCl for five times at 0.1 V/s. The CV exhibited one oxidation peak at around 1.3 V due to the monomer oxidation to form the polymer. There was a polymer reduction peak at around 1.1 V. The currents of these peaks increased as the cycle numbers increased, demonstrating that the polymer film immediately forms after the oxidation of the TTCA monomer at +1.3 V. The thickness of the poly-TTCA film increased as the cycle number increased and after five cycles it was estimated to be 200 nm. Characterization of Poly-TTCA/DGPD/Cyt c Microbiosensor. The covalent immobilization of DGPD lipid and Cyt c onto the poly-TTCA film was characterized using QCM, ESCA, and electrochemical impedance experiments. During the immobilization of DGPD onto the poly-TTCA film (Figure 2a,

Figure 2. (a) Quartz crystal microbalance analyses during the immobilization of (1) DGPD for 2 h, and (2) the sequential immobilizations of DGPD for 1 h on the poly-TTCA film and (3) Cyt c on the poly-TTCA/DGPD film. (b) ESCA analyses of (1) polyTTCA, (2) poly-TTCA/DGPD, and (3) poly-TTCA/DGPD/Cyt c surfaces. (c) The Bode and (d) Nyquist plots for a (1) poly-TTCA, (2) poly-TTCA/DGPD, and (3) poly-TTCA/DGPD/Cyt c modified probes in a 0.1 M PBS at pH 7.0. The impedance measurements were carried out at open circuit voltage. A simple equivalent circuit is shown in the inset of Figure 2d. 6656

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Ret is the electron transfer resistance, W is the Warburg element, and Qdl is the charge of constant phase element (CPE). By fitting the experimental data, Ret values can be determined for various modification steps. However, Ret can also be obtained from the diameter of the semicircle part at higher frequencies in the Nyquist plot, which controls the electron transfer kinetics of the redox probe at the electrode interface. Compared to poly-TTCA (curve a), the Ret of polyTTCA/DGPD increased from ∼965 to ∼2900 Ω (curve b), indicating successful immobilization of DGPD onto the polyTTCA film. The Ret value decreased from ∼2900 to ∼1956 Ω after the immobilization of Cyt c due to the presence of a haem group. The above QCM, ESCA, and EIS results clearly indicate that the DGPD and Cyt c covalently coimmobilized on the poly-TTCA film. Direct Electron Transfer of the Covalently Immobilized Cyt c. Figure 3a shows CVs recorded for the covalently

response 1), the frequency gradually decreased as the immobilization time passed on, indicating the successful covalent immobilization of DGPD onto the poly-TTCA film. The immobilization of the DGPD onto the poly-TTCA film was completed within 2 h at room temperature. After 2 h, the frequency change (Δf) was ∼224 Hz. For the microbiosensor fabrication, DGPD and Cyt c both was covalently bonded onto the poly-TTCA film, therefore, DGPD was immobilized for 1 h to have a partial coverage. After 1 h, the frequency change (Δf) was about 193 Hz (response 2) corresponding to a mass change (Δm) of 212 ± 6 ng, which was determined using an equation reported previously.24 The corresponding surface coverage of DGPD was 2.4 × 10−10 mol·cm−2. During the covalent sequential immobilization of Cyt c on the poly-TTCA film (response 3), the frequency change (Δf) was ∼430 Hz, which corresponds to a mass change (Δm) of 473 ± 12 ng. The corresponding surface coverage of Cyt c was 3.8 × 10−11 mol·cm−2. The above results suggest DGPD and Cyt c were covalently coimmobilized on the poly-TTCA film. To confirm the sequential immobilization of DGPD and Cyt c, ESCA analysis was carried out. Figure 2b shows the ESCA survey spectra obtain for poly-TTCA, poly-TTCA/DGPD, and polyTTCA/DGPD/Cyt c modified surfaces. The survey spectrum for the poly-TTCA surface exhibited two peaks at 284.4 and 531.5 eV corresponded to C−H, C−S, or C−C and CO bonds.25 In the case of the poly-TTCA/DGPD, the C1s peak at 284.4 eV shifted to 286.5 eV, which corresponds to C−N bonds.25 A N1s peak at 400.5 eV corresponds to NH2 groups was observed, which was absent in the spectrum of poly-TTCA surface. The presence of C−N bond and NH2 groups at 286.5 and 400.5 eV, respectively, clearly indicate that DGPD covalently immobilize onto the Poly-TTCA film through the formation of the amide bond between COOH groups of polyTTCA and amine groups of DGPD. Additionally, a P2p peak was observed at 136.6 eV indicates the presence of phosphate groups in DGPD. In the case of the poly-TTCA/DGPD/Cyt c, the C1s peak at 284.4 eV shifted to a higher energy of 286.3 eV, which corresponds to C−N bonds.25 A N1s peak at 400.5 eV corresponding NH2 groups was also observed. In addition, a Fe2p peak at about 717.5 eV was observed in the survey spectrum for poly-TTCA/DGPD/Cyt c surface. The presence of C−N, NH2, and Fe2p peaks at 286.5, 400.5, and 717.5 eV, respectively, and the C1s peak shifted to a higher energy from 284.4 to 286.3 eV clearly indicate that the Cyt c also covalently coimmobilized onto the Poly-TTCA film through the formation of the amide bond between COOH groups of Poly-TTCA and amine groups of Cyt c. Electrochemical impedance spectroscopy (EIS)26 was also used to further investigate the impedance changes of the electrode surfaces before and after coimmobilization of DGPD and Cyt c. Figure 2c shows the Bode plots obtained for polyTTCA, poly-TTCA/DGPD, and poly-TTCA/DGPD/Cytc modified electrodes in a 0.01 M PBS solution at pH 7.5. The bode plot indicates that the impedance value of the poly-TTCA modified electrode increased after the immobilization of DGPD on the poly-TTCA film. After the coimmobilization of Cyt c on the poly-TTCA film, the impedance value decreased due to the presence of a haem group in Cyt c. The Nyquist plot as shown in Figure 2d shows clearly that the charge transfer resistance increased due to the immobilization of DGPD onto the polyTTCA film. A general equivalent circuit can be utilized to model the impedance spectra of the present system as shown in the inset of Figure 2d, where Rs represents solution resistance,

Figure 3. (a) CVs recorded for (1) poly-TTCA/DGPD and (2) polyTTCA/DGPD in 0.1 M PBS at pH 7.0. (b) Scan rate dependency of the redox peak currents. (c) CVs recorded for a poly-TTCA/DGPD/ Cyt c modified electrode (1) without or (2−4) with various concentrations of O2•− (2) 0.5, (3) 1.0, and (4) 2.0 nM. (d) The schematic view of chemical and electrochemical reactions involved in the recognition and transduction steps.

coimmobilized Cyt c (poly-TTCA/DGPD/Cyt c) (1) and lipid (poly-TTCA/DGPD) (2) modified electrodes in a 0.1 M PBS at pH 7.0. A pair of redox peak was observed at approximately −0.16/−0.32 V vs Ag/AgCl for a poly-TTCA/DGPD/Cyt c modified electrode, which did not appear for a poly-TTCA/ DGPD electrode. The redox peaks might be due to the direct electron transfer of the covalently immobilized Cyt c. The formal potential of the redox reaction was determined to be −0.24 V vs Ag/AgCl. This value of formal potential is more negative than the value obtained in the previous reports using a poly-TTCA/Cyt c monolayer/multilayer modified electrodes using Langmuir-blodgett (LB) technique (+ 0.2 V),9 and Au/ poly- DATT/DGS/CL (+0.3 V) and Au/poly-DATT/DGS/ POPA (+0.29 V) electrodes.18 The significant negative shift observed in the present study might be related to the easy electron transfer of the haem group, due to the short electron transfer distance when Cyt c was directly immobilized on the poly-TTCA film. The negative formal potential of the immobilized Cyt c is comparable to those obtained at mixed 6657

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and electrochemical reactions involved in the recognition and transduction steps is shown in Figure 3d. O2•− was reduced by the surface confined Cyt c (Fe(II)) and it was oxidized to Cyt c (Fe(III)). The reduction current was due to the rereduction of Cyt c (Fe(III)), which was oxidized by the O2•−.33 The catalytic current increased with increasing concentrations of O2•−, indicating that the cathodic peaks solely arose from the catalytic reduction of O2•−. Although, the O2•− detection at this low potential is highly effective for the elimination of some interfering compounds in real samples such as brain, blood, urine, etc., the interference effects were further minimized by coating the sensor surface with polyethyleneimine (PEI) polymer. The specificity of the O2•− microbiosensor was examined by adding various interfering compounds in rat brain homogenates (Figure 4). The O2•− microbiosensor did not

monolayer modified electrodes composed of pyridine (−0.172 V) and nitrile (−0.415 V) functionalities.9 The peak currents were proportional to the scan rates, indicating that the redox reaction associated with a surface confined process of Cyt c (Figure 3b).27 The surface coverage of the immobilized Cyt c was calculated to be 4.0 ± 0.3 × 10−12 mol·cm−2 from the cathodic peak current value.26 This surface coverage value was higher than those obtained for the adsorbed Cyt c on Au/polyDATT/DGS/CL (2.8 × 10−12 mol·cm−2) and Au/polyDATT/DGS/POPA (1.5 × 10−12 mol·cm−2) bilayers,9 and imidazole- (3.3 ± 0.2 × 10−12 mol·cm−2) and pyridineterminated (2.5 ± 0.3 × 10−12 mol·cm−2) thiol monolayers modified electrodes.28 The observed higher surface coverage of Cyt c at the Poly-TTCA film revealed the strong interaction between Cyt c and poly-TTCA in the presence of DGPD lipid. The peak separation between the anodic and cathodic peaks was 0.16 V, indicating that the redox reaction of immobilized Cyt c on the Poly-TTCA film was controlled by the kinetic effect. The peak separations increased with the increasing of the scan rates, indicating that the redox reaction was quasireversible. The electron transfer rate constants (ko) of the immobilized Cyt c at the Poly-TTCA film was 0.62 ± 0.05 s−1 using the method for a surface-bound electrochemical system.29 This value is comparable to that obtained at poly-TTCA/Cyt c multilayers (0.68 s−1) modified electrode using Langmuir− Blodgett (LB) technique.18 In the CV recorded for a poly-TTCA/Cyt c modified electrode, a pair of less reversible redox peak with higher oxidation and reduction currents was observed (figure is not shown). However, electrode modified with only poly-TTCA does not provide any biological environment for the immobilized Cyt c. Therefore, the activity of the Cyt c toward the reduction of O2•− might be decreased when poly-TTCA/ Cyt c based microbiosensor is used for in vivo detection. The lipid molecules can provide biological environment for Cyt c, which improve the physical adherence and communication between Cyt c and biological compounds such as O2•−. Thus, the covalent coimmobilization of Cyt c and a lipid onto a polyTTCA film (poly-TTCA/DGPD/Cyt c) could make a stable O2•− biosensor for in vivo measurements of the O2•− detection. Application of Covalently Immobilized Cyt c Based Modified Electrode to Superoxide Detection. The covalently immobilized Cyt c modified electrode (polyTTCA/DGPD/Cyt c) was applied for the detection of the superoxide anion radical (O2•−). O2•− is a short-lived intermediate resulting from the 1 e− reduction of oxygen, and it is a primary component of so-called reactive oxygen species.30,31 O2•− was generated by the injection of KO2 in dimethyl sulfoxide (DMSO)32 into the phosphate buffer solution. This method was followed because the half-life of O2•− is very short in aqueous solutions. Figure 3c shows the CVs recorded for a poly-TTCA/DGPD/Cyt c modifiedelectrode in phosphate buffer solution (PBS) without or with various concentrations of O2•−. Even, the oxygen reduction potential was observed at the potential of about −0.5 V vs Ag/ AgCl, all CVs were recorded in PBS after purging the nitrogen gas for 30 min. In the absence of O2•− (CV (1) in Figure 3c), the CVs exhibited only the Cyt c redox peak. When various concentrations of O2•− (CVs 2,3,4 in Figure 3c) added into the PBS with stirring, the cathodic peak of Cyt c at −0.32 V increased, whereas the anodic peak at −0.16 V decreased, indicating that the immobilized Cyt c catalyzed the reduction of O2•−. The schematic view of the catalytic cycle of the chemical

Figure 4. Specificity experiment of the O2•− microbiosensor in rat brain homogenates.

show any response for glutamate, salicylic acid, glucose, DOPAC, ascorbic acid, dopamine, choline, oxygen, hydrogen peroxide, and nitric oxide in the rat brain homogenates. On the other hand, it showed stable and reproducible responses for O2•−, indicating that the O2•− microbiosensor was highly specific for the O2•− detection. Several reactive oxygen species such as O2•−, ONOO¯, and NO are formed in the brain due to the interlinked biological process and can be interfered with O2•− detection. However, the detection potential of O2•− was about −0.31 V vs Ag/AgCl, which was remarkably different from the detection potentials of NO (−0.7 V)34 and ONOO¯ (+0.2 V).35 Thus, the present sensor can selectively detect O2•− in the presence of other reactive species. Before the in vivo measurements of O2•− in the rat brain, the poly-TTCA/DGPD//Cyt c modified electrode was calibrated using the amperometric technique. The amperometric detection of O2•− was performed at −0.31 V by adding various concentration of O2•− in a 0.1 M phosphate buffer solution. Figure 5a shows the typical current−time plot obtained for the addition of O2•−. The catalytic O2•− reduction currents rose steeply to a stable value as soon as O2•− added to the PBS. The O2•− reduction response was confirmed using a superoxide scavenger, superoxide dismutase (SOD) as shown in Figure 5b. During the amperometric experiment, addition of SOD decreased the reduction current, whereas the addition of catalase did not change the reduction current. This indicates that the reduction current was due to the reduction of the superoxide radical. Figure 5c shows the calibration plots obtained for a O2•− microbiosensor (1) before and (2) after in vivo measurements. The sensitivity of the calibration plot slightly decreased after the in vivo measurements. The current response showed a linear relationship with the O 2 •− 6658

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Figure 5. (a) Chronoamperometric measurements with O 2 •− microbiosensors by successive additions of O2•− in a PBS solution at pH 7.0. (b) Current−time response obtained for a O 2 •− microbiosensor in a PBS solution upon addition of 5.0 nM O2•− and the subsequent additions of 300 units of catalase and 2 units/mL superoxide dismutase. (c) Calibration plots obtained for a (a) PolyTTCA/DGPD/Cyt c based O2•− microbiosensor (1) before and (2) after in vivo measurements. Applied potential: −0.31 V.

Figure 6. (a) Photograph of a microelectrode and (b) the placement of a microelectrode in a rat’s brain during in vivo experiment. In vivo amperometric responses recorded with (c) covalently immobilized Cyt c-based O2•− microbiosensors and (d) physically adsorbed Cyt c-based O2•− microbiosensor. The applied potential was set at −0.31 V vs Ag/ AgCl.

6c). On the other hand, in vivo performance of a O2•− microbiosensor based on physically adsorbed Cyt c on a lipid layer did not show any significant responses for in vivo measurements in saline, acute, and repeated cocaine treated rat brains (Figure 6d), indicating that this kind of microbiosensor is not stable enough for using in in vivo measurements. In fact, the amperometric responses rapidly decreased and approached the value of zero, which are similar to the response obtained for a null sensor. This result indicates that the physically adsorbed Cyt c detached from the lipid layer on the microbiosensor surface during in vivo measurements. For accurately quantifying the basal level of O2•−, the in vivo measurement was performed in the saline treated rat brain. Based on the postcalibration (calibration plot after in vivo measurement), the basal O2•− concentration was determined to be about 1.2 ± 0.06 nM. To ensure the monitoring of O2•− release by acute and repeated cocaine injections, control experiments were carried out by injecting saline (0.9% NaCl, 1 mL/kg i.p.) using a null microbiosensor that was prepared without Cyt c. Although, a Cyt c based superoxide sensor for the in vivo application was reported,37 no attempt was made for determining the basal O2•− concentration. Based on the in vivo measurement, we can conclude that the O2•− microbiosensor based on the covalent coimmobilization of lipid and Cyt c is highly sensitive and stable for detecting even the lower basal O2•− levels in the extracellular space of the dorsal striatum. As shown in Figure 6c, the response by acute cocaine was not significantly altered as compared to saline, suggesting that acute cocaine do not increase the extracellular O2•− level in the dorsal striatum (1.2 ± 0.06 nM). In contrast, repeated cocaine significantly increased the current response indicating that the repeated cocaine significantly altered the O2•− level in the dorsal striatum. The concentration of O2•− by repeated cocaine was determined to be 6.0 ± 0.7 nM. Figure 7 shows the semiquantitative analysis on the O2•− responses produced by saline, acute, and repeated cocaine to more accurately reflects the above results. These data suggest that repeated cocaine has the capability to increase the release of extracellular O2•− and

concentration in the range between 0.2 and 6.0 nM. The relative standard deviations (RSD) at 2.0 nM of O2•− were 4.7% and 5.2% for before and after in vivo measurements. These linear dependencies of the O2•− concentrations yielded two regression equations of ip (nA) = (2.3 ± 0.90) + (15.0 ± 0.30) [C] (nM)) and ip (nA) = (0.65 ± 0.43) + (11.7 ± 0.13) [C] (nM)) with the correlation coefficients of 0.998 and 0.999, respectively. The sensitivity of the microbiosensor was found to be decreased from 15.0 nA/nM to 11.7 nA/nM after in vivo measurements. The detection limits were 30.0 ± 0.9 pM and 45.0 ± 0.7 pM for pre- and post calibration plots, respectively, which were based on three-times measurements for the standard deviation of the blank noise (95% confidence level, k = 3, n = 5). The detection limit was much lower than those obtained from superoxide sensors based on the physically adsorbed Cyt c on cardiolipin (CL) and 1-palmitoyl-2-oleoylsnglycero-3-phosphate (POPA) lipid bilayers9 and multilayered Cyt c on a self-assembled monolayer (SAM).36 To examine the reproducibility of the microbiosensors response, chronoamperometric experiments were performed five times in 0.2 nM O2•− spiked-rat brain homogenates. The current response did not significantly changed when repeated experiments were performed with a single O2•− microbiosensor. The reproducibility expressed in terms of relative standard deviation (RSD) for 5 times measurements was 4.4%, indicating that the present O2•− microbiosensor exhibit appreciable reproducibility and stability. In Vivo Response of the O2•− Microbiosensor. The in vivo performance of the present O2•− microbiosensor based on covalently coimmobilized lipid and Cyt c was examined several times by inserting the microbiosensors separately for each of the saline-, acute-, and repeated-cocaine injected rat brains. The photograph of a microelectrode and the placement of it in a rat’s brain during the in vivo experiment are shown in Figures 6a and 6b. The O2•− microbiosensor worked well in response to saline-, acute and repeated cocaine-treated rat brains (Figure 6659

dx.doi.org/10.1021/ac301086m | Anal. Chem. 2012, 84, 6654−6660

Analytical Chemistry

Article

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Figure 7. Semiquantitative analysis on the O2•− responses obtained for the covalently immobilized Cyt c-based O2•− microbiosensors. * p < 0.05 as compared with the saline and acute cocaine groups.

the O2•− microbiosensor based on covalent coimmobilized lipid and Cyt c is an effective tool for monitoring of the changes of O2•− levels induced by an abused drug such as cocaine. All in vivo experiments were performed in three separate biological replicates, where the variation was