The amino acid L-glutamate (glutamate) is the major excitatory neurotransmitter in the mammalian central nervous system and as such underlies not only normal, but also many abnormal behaviors apparent in neurological and psychiatric disorders [1
]. Therefore, a tool for measuring glutamatergic transmission in a behaviorally relevant manner will greatly aid our understanding of these processes.
A variety of sampling methods for the measurement of extracellular brain chemicals, including glutamate, are available. One commonly used method, microdialysis coupled with high performance liquid chromatography, allows for the selective measurement of many different neuromodulators. Unfortunately, even advanced microdialysis techniques do not offer the temporal resolution required for sophisticated behavioral studies [6
]. Behavior, especially motivated behavior, can change within seconds of stimuli presentation [7
], and the 5-10 min temporal resolution of microdialysis [6
] time-averages these fast changes [7
]. Electrochemical sensors used with voltammetric recording techniques offer an alternative method for measurement of electroactive neurotransmitters, such as dopamine (DA), with improved temporal and spatial resolution [10
]. The non-electroactive nature of glutamate poses difficulties to its sensitive and selective measurement with such techniques. Fortunately, implantable biosensors, analytical tools consisting of both a biochemical recognition element and a physical transducer, circumvent these obstacles.
Amperometric electroenzymatic methods for the near real-time detection of glutamate have been developed using platinum electrodes modified with glutamate oxidase (GluOx) [11
]. GluOx is a flavoenzyme that catalyzes the oxidative deamination of glutamate in the presence of water and oxygen with the formation of α-ketoglutarate, ammonia and hydrogen peroxide (H2
]. Electrooxidation of the enzymatically generated H2
allows for effective glutamate detection [11
]. Unfortunately, efficient oxidation of H2
requires a high anodic potential at which electroactive interferents, such as DA and ascorbic acid (AA), are also oxidized and thereby contribute an undesired amperometric current signal [15
]. Several approaches have been taken to eliminate electroactive interference, such as immobilization of ascorbate oxidase [16
], coating with permselective polymers [15
], self-referencing [18
] and co-immobilization of peroxidase with a redox polymer [19
In addition to electroactive interference exclusion and temporal resolution, precise spatial resolution is also important to permit measurement of glutamate from discrete brain regions in vivo
. Glutamate can play differing roles in behavior based on the specific brain region or even subregion in which it is released [20
], therefore an optimal biosensor for glutamate would be able to make glutamate recordings from a population of cells within a single brain subregion. Previously, we and others described effective platinum wire-based electrodes for amperometric detection of glutamate [13
]. However, the signal-to-noise characteristics and required exposed surface of these electrodes make them sub-optimal for spatially precise glutamate measurements in vivo
. More recently, we adapted an over-oxidized polypyrrole coating approach to commercially available ceramic MEAs developed by Gerhardt and colleagues (Quanteon, LLC). Here, we describe the fabrication of significantly smaller, silicon wafer-based microelectrode array (MEA) probes coated with both Nafion and polypyrrole (PPy), which reduce signal from the interferents, AA and DA, respectively, to below baseline noise levels while maintaining the fast response time necessary for temporally precise measurements of glutamate in vivo
. These silicon wafer-based glutamate biosensors have been tested extensively in vitro
and have been applied to measurement of cortical electrical stimulation- and behaviorally-evoked glutamate release in vivo
2. Experimental Section
Nafion (5 wt.% solution in lower aliphatic alcohols/H2O mix), bovine serum albumin (BSA, min 96%), glutaraldehyde (25% in water), pyrrole (98%), L-glutamic acid, L-ascorbic acid, 3-hydroxy-tyramine (dopamine) were purchased from Aldrich Chemical Co. (Milwaukee, WI, USA). GluOx from Streptomyces Sp. X119-6, with a rated activity of 24.9 units per mg protein (U mg-1, Lowry's method), produced by Yamasa Corporation (Chiba, Japan), was purchased from Associates of Cape Cod, Inc. (Seikagaku America, MA, USA). Phosphate buffered saline (PBS) was composed of 50 mM Na2HPO4 with 100 mM NaCl (pH 7.4). Ultrapure water generated using a Millipore Milli-Q Water System was used for preparation of all solutions used in this work.
Electrochemical preparation of the sensors was performed using a Versatile Multichannel Potentiostat (model VMP3) equipped with the ‘p’ low current option and low current N' stat box (Bio-Logic USA, LLC, Knoxville, TN, USA). In vitro and in vivo experiments were conducted with a multichannel FAST-16 potentiostat (Quanteon, LLC, Lexington, KY, USA). Electropolymerization of PPy was conducted using a standard three-electrode system, consisting of a platinum wire auxiliary electrode, a glass encased Ag/AgCl in 3M NaCl solution reference electrode (Bioanalytical Systems, Inc., West Lafayette, IN, USA), and a platinum working electrode on our MEA probes. In vitro and in vivo measurements were conducted using a two-electrode system, with reference electrodes consisting of a glass-enclosed Ag/AgCl wire in 3 M NaCl solution (Bioanalytical Systems, Inc., West Lafayette, IN, USA) or a 200 μm diameter Ag/AgCl wire, respectively. All potentials are reported versus the Ag/AgCl reference electrode.
2.3. Electrode Fabrication and Polymer Modification
The MEA probes were fabricated at the Nanoelectronics Research Facility at UCLA. A 1 μm thick layer of silicon dioxide was grown thermally on a thin (150 μm) silicon substrate (Figure 1A
). The thermal oxide is a high quality dielectric film that electrically isolates the substrate from the metal layer subsequently deposited. Electron-beam evaporation was used to deposit 1000 Å of platinum on a 200 Å chromium adhesion layer. The metal was patterned by photolithography and lift-off to define the bonding pads, connections, and electrode sites (Figure 1B
). Next, plasma enhanced chemical vapor deposition (PECVD) was used to deposit a 1 μm layer of silicon dioxide (Figure 1C
). This second dielectric layer chemically isolates the connections from solution during electrochemical testing. After patterning of the oxide layer with a conventional photolithographic technique, the contact pads and electrode sites were plasma etched by reactive ion etching (RIE) (Figure 1D
). A third photolithographic treatment was performed to pattern the outline of the probes. RIE was then used to etch through the first and second dielectric layers, and deep reactive ion etching (DRIE) by the Bosch process was used to etch through the silicon substrate (Figure 1E
After the MEA probes were individually released from the wafer they were packaged and chemically cleaned to prepare the electrode surfaces for chemical modification with polymers and enzyme. Packaging involved soldering 28-gauge wire to the platinum bonding pads at the top of the MEA. Each MEA was cleaned with a 1:4 H2O2:H2SO4 solution. The tip of the MEA was lowered into the cleaning solution for 3 min and then rinsed with stirred purified H2O for 3 min; this process was repeated 3 times. Following cleaning, the electrodes were dried with argon. Each electrode was coated with PPy and Nafion. PPy was electrodeposited by holding the voltage constant at 0.85 V for 2.5–5 min until a total charge density of 20 mC/cm2 was reached in a 200 mM argon-purged solution of pyrrole in PBS at pH 7.4. The polymer Nafion was deposited on the sites by rapid dip-coating of the probe tips in the Nafion solution and oven-casting at 180 °C for 4 min, followed by 4 min cooling in ambient air. This process was repeated 3 times. After the polymer treatments, enzyme immobilization was accomplished by chemical crosslinking using a solution consisting of GluOx (2 wt%), BSA (2 wt%) and glutaraldehyde (0.125%). A ∼1 μL drop of the solution was formed on a syringe tip and fixed in place under a microscope. The probe was attached to a micromanipulator (Sutter Instruments) and positioned vertically relative to the enzyme solution droplet. With the aid of the microscope, the MEA was lowered into the enzyme droplet to either coat only the bottom 2, or all 4, electrodes. This was repeated 4 times with each application consisting of 2-3 dips. MEAs coated with PPy/Nafion and GluOx are referred to as MEA glutamate biosensors. The MEAs were sealed in a container with desiccant and stored at 4°C.
2.4. Electrode Characterization and Data Analysis
MEA biosensors prepared for glutamate detection were calibrated in vitro to test for sensitivity, selectivity and response time to glutamate. In vitro testing was carried out using constant potential amperometry with the FAST-16 electrochemistry system. A constant potential of 0.7 V was applied to the working electrodes against a Ag/AgCl reference electrode in 40 mL of stirred PBS at pH 7.4 and 37 °C within a Faraday cage. Data were collected at 80 kHz and averaged over 1 s intervals. After the current detected at the electrodes equilibrated to baseline (approx. 30 min), three 40 μL aliquots of glutamate (20 mM) were added to the beaker to reach a final glutamate concentration of 20, 40 and 60 μM glutamate. Additionally, aliquots of the potential interferents, AA (250μM final concentration) and DA (5-10μM final concentration), were added to the beaker in most tests to determine selectivity for glutamate. In some cases, lower concentrations of glutamate were added to the beaker (5-10 μM final concentration) to more accurately determine glutamate sensitivity. A calibration factor based on analysis of these data was calculated for each electrode on the MEAs to be used for in vivo experiments. In order to assess the sensitivity and response time to peroxide at sites uncoated with enzyme aliquots of H2O2 were also added to the beaker.
Estimations of MEA response time to glutamate were also made in vitro
using a custom-made flow cell chamber modeled after Lu et al.
]. The MEA was lowered into the plexiglass chamber such that the tip of the probe entered a narrow channel (2 mm diameter × 5 mm depth) through which PBS entered the chamber from below. Using a 60 mL syringe driven by a syringe pump, PBS was infused through the channel at a rate of 4 mL/min. For these experiments, a potential of 0.7 V versus Ag/AgCl was applied across the electrodes, data were collected at 80 kHz and averaged over 0.1 s intervals. A 1 mL sample loop filled with the analyte of interest (glutamate or H2
) was used to inject the analyte into the chamber over a 10-15 s period. A computer-directed pneumatic actuator controlled the switching of the sample injector, and the entire setup was housed in an incubator to maintain the temperature at 37 °C. Concentrations of glutamate and H2
ranging from 10-100 μM were used to determine MEA response time. Additionally the flow cell apparatus was used to confirm lack of DA interference at higher concentrations (20 μM). The FAST-16 system allowed precise marking of the event time for each sample injection. Data were output as current as a function of time and analyzed in Microsoft Excel. The response time for H2
detection at bare platinum sites was used as an estimate of the dead time in the system and subtracted from all measurements of response time at coated sites.
2.5. In Vivo Electrode Characterization and Data Analysis
The MEA glutamate biosensors were tested in 2 in vivo applications. Male Sprague Dawely rats (Charles River) were individually housed on a 12:12 light/dark cycle with ad libitum access to food and water. All experimental procedures and surgeries were conducted in accordance with the Institutional Animal Care and Use Committee and UCLA. Standard stereotaxic surgical techniques under halothane anesthesia were used to unilaterally implant a microbiosensor, pre-calibrated to glutamate (see above) into the nucleus accumbens core (NAc) of the ventral striatum (VS) using the following coordinates according to the atlas of Paxinos and Watson (4th ed.) (AP: +1.7, ML: -1.5, V - 6.0) Additionally, a bipolar stimulating electrode (Plastics One, Roanoke, VA) was unilaterally implanted into medial prefrontal cortex (mPFC) (AP: +3.2, ML: -0.8, V -4.4). A 200 μm diameter Ag/AgCl reference electrode was implanted contralaterally. The entire experiment was conducted inside a Faraday cage. The biosensor was connected to the FAST-16 potentiostat and a potential of 0.7 V versus Ag/AgCl was applied. Amperometric data were collected at 80 kHz and averaged over 0.1 s intervals. The electrode signal was allowed to equilibrate to baseline for approximately 30 min prior to application of 0.5 s stimulation trains of 1 ms duration 500-800 μA biphasic square wave pulses at 500 Hz (Coulbourn Instruments, Whitehall, PA) to the mPFC to elicit current changes detected in the VS at the PPy/Nafion/GluOx-coated electrode. Stimulations were administered 30 s apart. Using the Quanteon FAST-16 recording system, a record of the precise timing of the stimulation was included with the electrochemical dataset. An in vitro calibration factor was used to convert current changes detected at the electrode into glutamate concentration changes (see above). In addition to analysis of the maximal glutamate concentration change induced by each stimulation train, the temporal dynamics of the stimulation-induced glutamate spike were also analyzed (time to peak response and decay time).
For recording in the awake, freely moving animal connection wires from the MEA and reference electrodes were soldered to gold-plated sockets (Ginder Scientific) and the silicon wafer-based MEA was attached with epoxy to a 9-pin miniature connector (Ginder Scientific) such that all the sockets were encased in the connector. The entire assembly was sealed with epoxy to ensure full insulation and allowed to dry for 1 h prior to implantation. A sensor packaged for freely moving animal experiments is shown in Figure 2C
. The two-electrode MEA biosensor, prepared for glutamate detection and pre-calibrated, was unilaterally implanted, along with a contralateral 200 μm-diameter Ag/AgCl reference electrode, into the rat dorsal striatum (AP: +0.7, ML: +2.4, V -7.0) under halothane anesthesia. The connection assembly was anchored to the skull with three stainless steel skull screws along with dental acrylic cement (Bosworth Trim, Stokie, IL). A 48 h recovery period preceded onset of recordings.
Freely moving recordings were conducted in a plexiglass operant chamber (Med Associates, East Fairfield, VT) housed within a sound- and light-resistant shell within a Faraday cage. The recording headstage (Quanteon) consisted of a round miniature connector with 5 connector pins corresponding to the connector housing the implanted biosensor and reference lead. The headstage was tethered to a low torque 12 lead commutator (Airflyte, Bayonne, NJ) mounted on the top of the chamber. Outside of the operant chamber the commutator connected to the FAST-16 potentiostat. This apparatus allowed the animal unrestricted movement within the operant chamber. At test, the headstage was connected to the implanted MEA biosensor in the awake animal and a potential of 0.7 V versus Ag/AgCl was applied. Amperometric data were collected at 80 kHz and averaged over 1 s intervals. The sensor was allowed to equilibrate to baseline for 1 h before experimentation began. After baseline current was attained, the sampling average was lowered to 0.5 s for the majority of the experiment. Mild 1 s tail pinches were administered using stainless steel forceps. All data were plotted as current versus time (GraphPad Prism) and the in vitro calibration factor was used to convert current changes to glutamate concentration changes.