enzyme-modified microelectrodes for in vivo neurochemical measurements

405

Review

Enzyme-Modified Microelectrodes for In Vivo Neurochemical Measurements Paul Pantano’ and Werner G. KuhP” +

++ Department of Chemistry, University of California, Riverside, CA 92521, USA.

Received: March 28, 1994 Final version: June 6, 1994

Department of Chemistry, Ecole Normale Superieure, 24 rue Lhomond, F-75231 Paris Cedex 05, France

Abstract Various approaches to the production of miniaturized chemical sensors for in vivo applications are reviewed with special regard to the measurement of chemical transients in the mammalian brain. The basic tenants of biosensor principles are discussed in terms of the temporal response of enzyme-modified electrodes and selection of redox enzymes. The effects of electron transfer mediators and differential measurements to gain selectivity are also examined. A broad range of immobilization methods including membrane entrapment, physical and chemical adsorption, cross-linking agents, polymeric entrainment, electropolymerized films and covalent derivatization are examined. The use of polymer layers and redox polymers to gain selectivity are discussed. The discussion is limited to microelectrode surfaces ( < 100 pm diameter) and various electrode types are examined, including bare and platinized carbon fibers, platinum microelectrodes, conducting organic salts and carbon paste electrodes.

Keywords: Microelectrodes, Enzyme electrodes, Neurochemical measurements.

1. Introduction

Many of today’s most challenging analytical measurements are carried out only in the same environment in which the analyte of interest normally exists. In most cases, the limitations of these in situ measurements stem from the spatial and temporal constraints imposed by the sample environment itself. For example, consider the microenvironment of the mammalian brain and the dynamic chemical events that must be dealt with in order to monitor the stimulated release and uptake of a specific neurotransmitter in vivo.

Neurotransmitters, the diffusible chemical messengers which relay electrical signals from one nerve cell to another, are in rapid transition in the extracellular fluid (ECF) during the course of neurotransmission. This limits their residence time in the ECF, and therefore the distance over which they can freely diffuse (typically 1-25 pm). Consequently, the sample available for analysis is extremely small (on the order of a few hundred attomoles), is localized to a very small volume (typically a few picoliters) and is available only for a short time (approximately a few hundred milliseconds). In addition to these features, it is important to be able to localize such a measurement since the distribution of nerve terminals in the brain is extremely heterogeneous.

The extracellular fluid is an extremely complex environment since all chemicals entering and leaving a cell must pass through this space (e.g., nutrients, metabolites, waste products, hormones, drugs, toxins, neurotransmitters and numerous elemental ions). The ECF is an extremely hostile matrix because it contains surfactants (e.g., lipids), electrode poisons (e.g., proteins), electrocatalysts (e.g., glutathione and ascorbate), and immunological agents that will react to a foreign entity (e.g., an implantable probe) [l]. In addition, the ECF is an extremely sensitive matrix; specifically, one must consider the physiological concentration range within which the removal of a chemical species from this matrix would not disturb any of the biochemical processes that occur there [2].

Therefore, it is not surprising that there are relatively few analytical techniques that have the requisite sensitivity,

selectivity, probe size, and more importantly, temporal resolution, to make physiologically relevant measurements in this environment. In fact, the only analytical techniques that have found widespread use for the measurement of neurotransmitter dynamics in vivo involve the use of rapid electrochemical techniques and microelectrode sensors (electrodes of micro- meter dimensions) [3, 41.

Currently, one of the most important and active areas of research within the neuroscience community is the mechanism by which the brain releases amino acid neurotransmitters. Unfortunately, the direct voltammetric detection of excitatory and inhibitory amino acid neurotransmitters is not possible because unlike the catecholamines, the amino acids (and the majority of other neurotransmitters and metabolic intermedi- ates) are not electroactive at analytically useful potentials. This work is intended to review the generation and utilization of miniaturized amperometric sensors that will allow the measurement of these nonelectroactive neurochemicals in vivo under physiologically relevant conditions.

2. Discussion

2.1. Miniaturized Chemical Sensors for In Vivo Applications

The pioneering work of Adams has lead to the development of in vivo monitoring techniques for catecholamines and other electroactive neurochemical species [ 5 ] . Electrodes designed for monitoring transient neurochemical events in vivo must not only possess a rapid response time but must be miniaturized so that spatial resolution is possible. For example, miniaturization of electrodes has allowed electroanalytical methods to examine neurochemical and neurophysiological events in extracellular space as well as inside single cells [6]. There are many advantages that accrue from the miniaturization of a chemical sensor. First, decreasing the amount of sample required for the analysis is advantageous since the consumption of analyte can adversely effect the organism and the biological system under investigation. Miniaturization improves biocompatibility by minimizing

Electroanulysis 1995, I , No. 5 0 VCH Verlugsgesellschuft mbH. 0-69469 Weinheim. 1995 1040-039 719510505-405 $ 5 .OO+ .25/0

406 P. Pantano, W.G. Kuhr

immunological responses (i.e., the amount of tissue encapsula- tion or the degree of an inflammatory reaction over the time course of the measurement). It permits measurements in discrete and inaccessible locations and/or small volumes, and it allows the determination of spatial concentration profiles (i.e., to provide the spatial resolution and accessibility to measure processes occurring in dense heterogeneous nuclei).

Wightman and others have developed a Vsnge of microelectrode sensors constructed out of carbon fibt;;, platinum, and gold [7]. Since the characteristic electrode area--to-diffusion layer thickness ratio is small, the behavior observed at these microelectrodes can be significantly different from that observed at conventionally sized macroscopic electrodes. For example, the resulting small currents (in the pico- to nanoampere range) make possible the use of faster scan rates in cyclic voltammetry experiments. When these sensors are implanted into living tissue, the electrodes will sample the fluid in the “pool” immediately surrounding the electrode tip [8]. In combination with the inherently small size of these probes, this phenomena leads to the exceptional spatial reisohtion observed with these probes.

Inherent problems with the majority of in vivo probes is their fragility and the drift of the response [9]. This is especially true at bare gold or platinum electrodes, where strong adsorption of proteins and other biomaterials can virtually eliminate electron transfer sites within minutes of implantation within tissue. Carbon fibers are the most common microelectrodes used today since they are strong, stable and facilitate various electrode modification procedures. Carbon is less prone to surface changes over the long periods of time which are required for the majority of in vivo experiments. While a considerable portion of the sensors’ initial response can be lost following implantation into a biological sample, carbon electrodes will reach and maintain a steady-state fairly quickly following implantation [lo].

2.2. Biosensor Principles

A biosensor can be described as an analytical detector (typically an electrode or a fiber optic transducer) whose selectivity is enhanced by immobilizing a sensntive and selective biological element (typically an enzyme) within close proximity of the sensor. Amperometric biosensors measure the current produced during the oxidation or reduction of a product or reactant from an enzyme-catalyzed reaction, usually at a constant potential (i.e., an externally applied potential drives the electrochemical reaction and the faradaic current is measured). Since amperometric enzyme electrodes consume a specific substrate via the enzymatic reaction (which alters the concentration profile of the substrate in the active layer), they usually display an expanded linear response range and larger apparent Michaelis-Menten constant (K,) than their potentio- metric enzyme electrode counterparts (where an electrochemical equifibrium is established at the sensor interface and changes in the electrochemical potential are measured).

While the efficacy of enzyme electrodes has been demonstrated in many analytical applications, they do suffer from a number of practical limitations which has prevented their widespread usage. The advantages of immobilized enzyme electrodes include the ease of analyte determination in complex mixtures with minimal pretreatment, the use of small sample volumes, and the ability for inexpensive and repeated use. However, the usefulness of the immobilized enzyme electrode depends on factors such as the method of immobilization, the

chemical and physical conditions of use (pH, temperature, ionic strength of the sample, long term stability of the biocomponent, interfering species, etc.), the activity and stability of the enzyme once immobilized, the stability of the electrochemical sensor, and the response time and the storage conditions required u11.

2.3. Temporal Response of Enzyme-Modified Electrodes

Several investigators have modeled the response of amperometric enzyme sensors mathematically with respect to properties of the immobilized enzyme layer or membrane, the geometry of the electrochemical probe, the kinetics of the enzymatic reaction, etc. [12-151. The overall sensor response depends upon the detector (the rate of electron transfer), the rate of enzymatic catalysis, and the mass transport processes in operation. In other words, the sensor response is controlled by mass transport when the enzyme reaction proceeds faster than the rate of mass transport. Traditionally, immobilization procedures involve the deposition of a large quantity of enzyme in a semipermeable matrix so that the response of the analyte would be mass-transport limited (and therefore independent of the surface concentration of enzyme). This method of enzyme immobilization increases the long-term Stability of the biosensor, increases the linear dynamic range of the determination, and decreases the susceptibility to low concentrations of interfering chemical species. If an enzyme electrode was actually operated under kinetically controlled conditions, the current-to-concentration relationship would be nonlinear with a useful range of less than one order of magnitude [16].

The response time of an enzyme-modified electrode has been shown to be dependent upon the diffusion of substrate, the concentration of the substrate, the concentration of the enzyme in the layer, the pH of the solution, the temperature of the solution, the thickness of the enzyme layer, the porosity of the enzyme layer or membrane (or the diffusion coefficient of the substrate/products through the enzyme layer relative to that in the bulk solution) and the geometry of the enzyme probe [17]. The establishment of a steady-state response is essentially determined by the diffusion of the substrate through the enzyme layer for an enzyme with high catalytic activity (in contrast to an enzyme sensor displaying low catalytic activity, where the response time of the overall sensor is considerably longer) [ 181. It should be noted that the response values in this review are not absolute since it is very difficult to compare the response times reported by different investigators since the temporal response is dependent upon many different experimental parameters (e.g., concentration of the electroactive analyte, the geometry and area of the electrode, the flow rate and volume in a flow injection apparatus, stirring rate, or rotation rate of a rotating disk electrode, the choice of steady-state vs. transient methods of current measurements, etc.) [19, 201.

2.4. Redox Enzymes used with Biosensors

Enzymes can be classified into many different groups but clearly the most important class are the oxidoreductases which catalyze the oxidation of compounds using oxygen or NAD’. Oxidoreductases have traditionally been divided into two categories: the NAD(P)+/NAD(P)H dependent dehydrogenases and the flavoprotein oxidases. Each class of enzyme has its set of advantages and limitations for use in biosensors

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Enzyme-Modified Microelectrodes for In Vivo Neurochemical Measurements 407

(e.g., O2 dependence of oxidases; high overpotential of NADH oxidation for dehydrogenases).

If it were possible to designate a single enzyme as being ideally suited for use as a biosensor, glucose oxidase (GOD; oxygen-l- oxidoreductase; E.C. 1.1.3.4 isolated from Asperigillus niger) would top the list owing to its high specificity for an important analyte, its high turnover rate, and its high stability [21]. Without question, it is the most widely used and thoroughly investigated enzyme in the biosensor field; an amperometric sensor is typically used to measure either the consumption of oxygen or the production of hydrogen peroxide [22]. However, low and/or variable oxygen tension can lead to fluctuations in the response of these enzyme electrodes. If there were an ideal alternative to the oxidoreductases, this enzyme would possess a tightly bound redox center, exhibit a high turnover number, and require no additional soluble cofactors. One recently appearing alternative involves the use of quinoproteins.

Glucose dehydrogenase (E.C. 1.1.99.17) is one of the recently characterized class of quinoprotein enzymes that possess pyrroloquinoline quinone (PQQ) as its prosthetic group [23- 261. Like the NAD(P)+-dependent glucose dehydrogenase (E.C. 1.1.1.47) and in contract to glucose oxidase, this enzyme is virtually insensitive to oxygen content even though it is still capable of using a wide variety of electron acceptors for the catalytic oxidation of glucose. Unfortunately, only a small number of these enzymes have been thoroughly characterized and it remains to be seen how many possible quinoproteins can be easily purified and obtainable in large quantities [27].

The overriding problem in the development of amperometric flavoprotein oxidase sensors is the ability to electrochemically reoxidize the enzyme in such a way that the oxygen dependence could be avoided. While these enzyme electrodes are useful in well-oxygenated tissues, there are several physiological samples and/or conditions where, despite the low K, of GOD for oxygen, the supply of oxygen may be quite low compared to substrate levels. For example, the local oxygen concentration inside different biological systems will vary widely as a function of metabolic activity; in fact, there are anaerobic zones in certain microbes which would make such a measurement virtually impossible [19, 281. Additionally, in more complex systems such as the in vivo measurement of neurotransmitter dynamics, oxygen is closely linked to the release, uptake, synthesis, and vesicular transport of neurotransmitters [29] (especially in studies involving hypoxic tissues) [30]. In these cases, not only will the upper limit of the linear range of detection be reduced, but the concentration of oxygen will limit the rate of the enzymatic reaction. Two solutions to this problem involve the use of an electron transfer mediator and/or the employment of a differential measurement.

2.5. Electron Transfer Mediators

The use of electron transfer mediators has significantly improved the scope and performance of amperometric enzyme electrode sensors. Most of the recent research has attempted to replace oxygen by using nonphysiological electron transfer mediators immobilized on the electrode surface or within the enzyme layer (e.g., adsorption, codeposition with a polymer coating, or covalent attachment) [31, 321. Provided that it does not react with oxygen, the electron transfer mediator substitutes for oxygen in the enzymatic reaction, and the rate at which the mediator is produced is measured amperometrically at a suitable electrode.

It is not easy to obtain a practical electron transfer mediator

for use in an in vivo setting since a number of criteria have to be fulfilled. An ideal mediator should have a low molecular mass (which allows it to shuttle electrons from the redox center of the enzyme to the surface of the electrode), exhibit reversible electron transfer kinetics at the electroactive surface and rapid homogeneous kinetics for the enzymatic reaction in solution, allow the regeneration of the oxidized form of the mediator to be accomplished at a low overvoltage, be pH and temperature independent, be stable in both its oxidized and reduced forms, and be easily immobilized onto the electrode surface. Unfortu- nately, most of the mediators which meet these initial criteria are toxic in biological tissues [ 3 3 ] . In addition, care must be exercised to ensure that the problem of electrochemical interference is not enhanced by the use of a mediator; oxidized ferrocenes, for example, are reduced by ascorbate and GOD [211.

2.6. Differential Measurements

In common with general analytical practices, certain interferences (e.g., fluctuations in pH, temperature, and ionic strength) can be minimized and/or compensated for by the judicious use of “blanks”. Although differential measurements do not prevent interferences, they are commonly employed to compensate for their effects [34]. In general, the method consists of subtracting the signal of a control electrode from the signal of the active working electrode. The precision of the method is dependent on how well matched the control and working electrodes are. Control electrodes are constructed in an identical fashion to the working electrode with the exception that the enzyme has been eliminated or deactivated via a heat treatment. There are problems in both cases. First, if the enzyme is not added to the control electrode, one cannot expect a protein covered electrode to behave similarly to a bare electrode. Second, while heat treating an enzyme does deactivate it and ensures that both electrodes have protein on them, heat treating an electrode can dramatically change the electron transfer properties of the electrode surface [35-381.

3. Enzyme Immobilization Methods

The interface of an electrochemical biosensor, which typically contains the enzyme and an electron transfer mediator, must also be compatible with the chemical environmental require- ments for enzymatic activity and not interfere with the electron transfer process [39, 401. This requires careful optimization of the procedures used to provide confinement of the enzyme to the vicinity of the electrode surface. Many different strategies and variations of strategies have been used to immobilize enzymes onto electrode surfaces. In general, this has been accomplished in one of four ways: entrainment within a membrane, physical adsorption, entrapment within a polymeric matrix, or covalent bonding to the electrode surface [16, 41-43].

3.1. Membrane Entrapment

Discrete macroscopic membranes are well-known for their ability to overcome the problems associated with the fouling of electrode surfaces by large molecular species and various interfering electrochemical species. Immobilization of the enzyme via entrapment within a dialysis membrane [44-471 and/or the covalent attachment of the enzyme onto the dialysis


membrane [48-501 is highly advantageous in that it can provide a stable environment for virtually any redox enzyme. The drawback, however, is that it also adds a diffusional barrier into the sensing scheme which is difficult to overcome if a subsecond measurement is necessary. In general, the biggest obstacles in obtaining an improved temporal response are associated with the physical size of the membrane-enzyme layer (which is often hundreds of microns thick). Additionally, since the porosity of the membrane usually changes with time (as entrapped enzyme and/or proteinaceous material from the sample being to adsorb upon the membrane), the diffusivity of the analyte of interest will not be constant and frequent calibrations would be necessary to maintain reliability [5 I]. For example, Mascini and Selleri have shown that, depending on the sample matrix, the response time of an amperometric biosensor can vary 8-fold with respect to the choice of material from which the dialysis membrane is constructed [44]. Similarly, Gunasingham and co- workers showed that by increasing the hydrophobicity of a cellulose acetate dialysis membrane (by a selective acylation reaction at the C-2 position), the improved biocompatibility of the sensor lead to dramatic improvements in the sensor’s sensitivity, selectivity, and stability [48]. Finally, Montagne et al. have demonstrated that enzyme immobilization conditions can be just as critical to the overall performance of the sensor as is the effect of a complex sample matrix. In this work, the response time of a dialysis membrane sensor could be improved by 200% when the immobilization of enzyme was restricted to the inside of the dialysis membrane (as opposed to allowing enzyme to be coupled to both sides) [50].

Considerable improvement in response time is found through miniaturization of the dialysis membrane/electrode assembly. Kuhr and co-workers have constructed a glutamate dehydrogenase (GDH) - modified microelectrode in which a 150-pm- i.d. (9000 MWCO) microdialysis fiber was used to entrap GDH near the tip of a 10-pm-diameter carbon fiber microelectrode [52]. In this work, a fast scan voltammetric measurement was used to monitor the NADH generated following the enzymatic deamination of glutamate and the response titne of these sensors was on the order of 15 s (63% of the steady state signal). One of the reasons for such a fast response with this type of sensor stems from the fact that a great deal of care was taken in the characterization of the physical state of the dialysis membrane throughout the various steps in the immobilization process. This is especially critical since, amongst other considerations, the temporal response of these sensors will be dependent upon the porosity or the molecular weight cut-off (MWCO) of these small diameter dialysis fibers when used in conjunction with a large, diffusible electron transfer mediator like NA:DH.

Using a similar design, Boutelle et al. used an amperometric platinum microelectrode and a 240-pm-diameter dialysis membrane to entrap glutamate oxidase and ascorbate oxidase [53]. This probe had an ex vivo response time of less than 10s and once implanted in the striatum of a rat, the feasibility of using these microdialysis fibers to monitor physiologically induced changes in both glutamate and ascorbate in vivo was demonstrated. The same group described a dialysis membrane- based enzyme sensor where the working, counter and reference electrodes were inserted into the tip of an operational microdialysis system [54, 551. With this probe, one slowly passed a Ringers solution through a short length of dialysis membrane implanted into the tissue of interest (just as one does in a conventional microdialysis experiment). In conjunction with a platinum working electrode that is held at a potential of 650 mV (vs. Ag/AgCl), the enzymatic production of hydrogen peroxide by glutamate oxidase entrapped within the dialysis

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membrane was monitored. This probe was also used for the continuous measurement of glutamate efflux in the brain of a freely moving rat following a 5-min tail pinch; electrochemical interferences from ascorbate were eliminated through the use of ascorbic acid oxidase. While the in vivo response time of this probe was on the order of one minute, the authors have recently reported an ex vivo response time on the order of 20s for a dialysis-electrode that was constructed in the exact same manner with the exception that it was selective for glycerol [55].

3.2. Physical and Chemical Adsorption

The simple physical or chemical adsorption of a biocomponent onto an electrode surface (whether or not predictably invoked) is still a frequently employed immobilization metho- dology since it is the simplest technique by which a biomolecule can be attached to a solid support (i.e., no coupling reagents are required) [56-611. The major disadvantages of this method are two-fold. First, there are very few parameters that can be manipulated to control the surface deposition process. For example, if an enzyme is adsorbed too weakly, the long term stability of the enzyme electrode will be compromised. Conversely, if the interaction is too strong, the enzyme might be rendered inactive. In either case, the lifetimes of adsorbed enzymes are generally shorter than the lifetimes observed via other immobilization techniques [34]. The second disadvantage of this method (as well as most methods where an enzyme is covalently attached onto an electrode) is that the enzyme is placed into the electrochemical double-layer (where the quaternary and even the tertiary structure of the enzyme may be considerably different than that found in homogeneous solution). In addition, when the enzyme is localized in this interfacial region, it can be subject to dramatic variations in pH and ionic strength such that its reaction kinetics will be very different than those observed in homogeneous solution [ 131.

Nonetheless, one of the earliest, if not the first, demonstra- tions of an enzyme-modified microelectrode that was suitable for insertion into brain tissue was constructed by Silver who simply dipped a 0.1-pm-diameter, platinum coated, glass micropipette electrode into an aqueous solution of GOD [62]. In this work, physiological concentrations of glucose were analyzed at the cellular level by the amperometric detection of enzyme-generated hydrogen peroxide at an applied potential of 600mV (vs. Ag/AgCl). While there were difficulties in the quantitation of the glucose concentration (stemming from the high and/or fluctuating in vitro concentrations of oxygen and/or ascorbate), these twenty year old enzyme electrodes had remarkably fast response times (2-3s for 97% response and > 30 s when a protective nitrocellulose membrane was additionally employed).

3.3. Cross-Linking Agents

Dramatic improvements in the long term stability of adsorbed enzyme-modified electrodes have been achieved through the use of bifunctional cross-linking agents, most notably, gluteraldehyde [63-671. These agents diminish the desorption rate of enzyme from the electrode surface through the formation of enzyme-to-enzyme and enzyme-to-electrode covalent linkages. In practice, bovine serum albumin (BSA) is utilized in conjunction with gluteraldehyde in order to create a “protein- membrane” where the BSA is believed to relive steric interactions throughout the three-dimensional cross-linked matrix [60, 68-72]. However, care must be exercised with

Enzyme-Modified Microelectrodes for In Vivo Neurochemical Measurements 409

these methods since increasing the thickness of the membrane layer will slow the response time of the adsorbed enzyme monolayer (or submonolayer) [ 12, 15, 731. Unfortunately, the control of this reaction is not trivial since the degree of cross- linking depends on many factors (e.g., pH, ionic strength, temperature, and reaction time). For example, while Watanabe and co-workers were able to increase the loading of enzyme by a factor of 150, the corresponding change in sensitivity was observed to increase by only a factor of 30 which suggested that as much as 80% of the enzymatic layer was denatured during the cross-linking process [74].

Numerous variations of this reaction scheme have been used with enzyme-modified microelectrodes to produce a thin layer of immobilized enzyme and a fast temporal response. Tamiya et al. produced a fast responding (ca. < 10 s) glucose-sensitive microelectrode by dipping a 12-pm-diameter carbon fiber microelectrode into a premixed solution of gluteraldehyde, BSA and GOD [60]. However, while the use of electrochemical pretreatments resulted in a substantial improvement in sensitivity to hydrogen peroxide at a carbon surface (0.1 pM hydrogen peroxide detection limit), oxygen and hydrogen gases evolved during the course of this severe pretreatment lead to the deactivation of the enzyme over the course of time. In another approach, Ma and Cleemann treated a 2-5-pm-diameter ammonia gas sensing microelectrode with a solution of gluteraldehyde after a mixture of urease and BSA was first evaporated onto the electrode surface under vacuum [66]. Since the response time of the enzyme-modified ammonia sensor (15-20s for 95% response) was only slightly slower than the 10-15s response observed at the bare ammonia sensor, this immobilization technique appears to have resulted in a thin layer of surface-localized enzyme. Another novel method of enzyme immobilization was described by Suaud-Chagny and Gonon [75], where a 500-pm long, 12-pm-diameter pyrolytic carbon fiber microelectrode was electrochemically coated with an inert sheath of BSA before lactate dehydrogenase was added to the activated carbon surface. Differential normal pulse voltammetry of enzyme-generated NADH was used to demonstrate a response that was sensitive (10 pM detection limit for NADH), fast (ca. 500 ms), and stable (ca. 150 h). A key feature of this technique was that the thickness of the inert protein sheath was found to be amenable to experimental control (e.g., BSA concentration) as evaluated by the magnitude and stability of the sensors’ response. More recent work on the use of adsorption and various cross-linking agents to immobilize peroxidases in combination with oxidases has also been performed on carbon fibers [76]. Various procedures were used to immobilize horseradish peroxidase (HRP) and heat treatment of the carbon fiber surface was found to dramatically improve the current response and linear dynamic range of the sensor.

3.4. Platinization

Ikariyama and co-workers have performed a tremendous amount of research with respect to the adsorption of enzymes onto various forms of platinized microelectrodes [77-791. The electrochemical deposition of platinum onto the platinum microelectrode surface results in a highly dispersed layer known as “platinum black”. A platinum black surface is characterized by the presence of a deposited surface layer of porous microplatinum particulates which possess a very large surface area, a high catalytic activity for electrolytic reactions, and a strong tendency to facilitate the adsorption of protein.

Amperometric glucose oxidase microelectrodes with tip dia- meters ranging from 10-200 pm were constructed by adsorbing glucose oxidase onto a platinized platinum surface where the added stabilization of the enzyme layer was realized following the application of a thin layer of gluteraldehyde/BSA [77, 781. With this approach, the platinum particles play the dual role of transducing material and matrix for enzyme attachment; typical response times were between 0.5-3.0s (100% signal) and a stable response could be readily observed over a period of 4 weeks. In association with pulse voltammetric detection, these highly sensitive sensors were utilized to determine the real time concentration of glucose in a stationary droplet of blood (ca. lOpL) [80, 811 and in a flow injection analysis system [79].

Since platinized platinum surfaces are very reactive, care must be taken in the methods used to prepare these surfaces for the optimized performance of both enzyme immobilization and electron transfer duties. This has become especially critical since the electrochemical deposition of the platinized particles can also be performed simultaneously with the electrochemical adsorption of enzyme molecules [61, 821. Improvements in the selectivity of these sensors towards glucose in the presence of other saccharides can be controlled by an anodic pretreatment of the platinized platinum microelectrode before the enzyme immobilization step [83]. Similarly, the effect of performing an anodic pretreatment on a 100-pm-diameter platinum wire microelectrode before the platinization reaction was evaluated by Beh et al. [84]. Not only was the stability of the platinized GOD microelectrode increased, but the detection limit was lowered by an order of magnitude relative to control without any significant change to the response time (ca. 25 s).

A number of electrode materials other than platinum have been platinized in order to enhance the performance of GOD- modified microelectrodes. A novel method which combined covalent attachment of the enzyme to the carbonaceous region of a partially platinized carbon surface was utilized by Yacynych and co-workers to fabricate a glucose oxidase-based sensor [85]. The partial platinization of reticulated vitreous carbon electrode surfaces effectively delegated the task of enzyme immobilization and electron transfer to separate areas of the same surface (reducing the overpotential required for the oxidation of hydrogen peroxide); improvements were observed in the areas of precision, sensitivity, signal-to-noise ratio (S/N) and sampling rates. This work was extended to carbon fiber microelectrodes, where a glucose sensor was fabricated using an 8-pm-diameter carbon fiber microelectrode [86]. The platinized carbon fiber was very sensitive to glucose and had a response time of 15-45 s (100% response). Wang and Angnes simultaneously deposited rhodium microparticles and GOD onto 7- pm-diameter carbon fiber microelectrodes to produce a fast- responding sensor (3 s) that exhibited a catalytic superiority over platinized surfaces with respect to the oxidation of hydrogen peroxide [87]. In contrast to the 600-800 mV (vs. Ag/AgCl) potential used at most platinized surfaces, a considerable response for the oxidation of hydrogen peroxide was observed at 250mV (vs. Ag/AgCl) at the rhodium microparticles. The obvious advantage of this approach is that the lower applied potential minimizes the response of other electroactive species that are normally present in many samples.

Some of the fastest responding microbiosensors that have been constructed and used in hostile biological environments have been fabricated with platinization methods for enzyme immobilization. Kim and Lee silanized the glass insulation of a 3-p-m-diameter platinum microelectrode to covalently bind gluteraldehyde across the probe tip [67]. Following the evaporation (under vacuum) of a dilute solution of glucose

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oxidase onto this cross-linked layer, the amperometric probe was finally coated with a 10% solution of polystyrene acrylonitrile. These microprobes were operated at 700 mV (vs. Ag/AgCl) to detect enzyme-generated hydrogen peroxide and were observed to respond to a step change in glucose concentration within 200 ms (63% response time). While the accuracy of the measurement might have been affected by local concentration of pH and dissolved oxygen, these microelectrodes were capable of measuring local glucose concentrations in a 150-pm-thick microorganism (Corynebacterium glutarni- ciurn) that was grown on an agar plate. Extremely small glucose- oxidase-modified microelectrodes (2 pm diameter platinum disks) with very fast response times were constructed by Karube and co-workers [88].

Abe and co-workers modified a carbon ring microelectrode by codepositing glucose oxidase and platinum particles onto the surface to produce an intracellular amperometric glucose probe with response times as low as 270ms (10-90% of the full response) [72, 891. These sensors were used to demonstrate the feasibility of measuring glucose inside a single cell at atmos- pheric levels of oxygen tension. 2pL of glucose was iontophor- etically injected into the cytoplasm of the pond snail, Planorbis corneus and the current response from the GOD microsensor was used to measure the time course of glucose metabolism and uptake into cellular compartments. Glucose levels in the submillimolar range could be measured without complications until the oxygen concentration dipped below a level of 200 pM.

3.5. Conducting Organic Salts

In 1965, Melby synthesized planar organo-metallic complexes whose extended pie-orbital systems could be arranged in a stacked configuration to display metallic like conductivity [90]. Since then, the exceptional physical, chemical, and electrochemical properties of these conducting salts (i.e., their high electrocatalytic properties, long term stability, and minimal susceptibility to fouling) have been the subject of numerous investigations [9 1-95]. To date, the most frequently employed complexes are synthesized using 7,7,8,8-tetracyanoquinodi- methane (TCNQ) as the electron donor, and tetrathiafulvalene (TTF) or N-methylphenazinium (NMP) as the electron acceptor. In biosensor applications, the first reason for the widespread usage of these complexes is their ability to play host to an enzyme through a large number of different immobilization strategies (e.g., enzymes and conducting salts can be comixed into a carbon paste or polymeric film). In fact, these conducting salts can be compressed and molded to become the actual electrode material/surface and an enzyme can be immobilized through simple adsorption. The second reason for their widespread usage is that they have been successfully employed with all three of the main classes of redox enzymes (i.e., the FAD-dependent oxidases, the PQQ-dependent dehydrogenases and NAD-dependent dehydrogenases).

While there are many advantages stemming from the combination of electrochemical mediators and enzyme electrodes, there is one disadvantage which can be a severe detriment when a mediated biosensor is employed in an in vivo measurement. As these mediators shuttle electrons between the enzyme and the electrode, it is inevitable that some of the mediator molecules will desorb from the surface [21]. In addition to the effect this will have on the performance of the biosensor (and the ability for a meaningful determination of the postcalibration response), as was noted earlier, the vast majority of mediators are toxic. The additional advantages of a biosensor

which utilizes TTF-TCNQ in an in vivo setting are two-fold. First, when this conducting salt is compressed under high pressure to serve as the actual electrode material, the amount of salt that desorbs over an extended period of time is essentially nil; and second, as reported in a recent study conducted by Kulys et al., TTF and TCNQ were found to be relatively nontoxic [96].

As a result of these features, one very successful and two very promising TTF-TCNQ microbiosensors have been designed for in vivo applications. Hale and Wightman reported the construction of an enzyme-modified TTF-TCNQ microelectrode for the determination of the neurotransmitter, acetylcholine, and its synthetic precursor, choline. In this work, acetylcholine esterase and choline oxidase were adsorbed separately (or jointly) onto the conducting salt surface without the need for any other type of supporting membrane [97]. In this configuration, an applied potential of 150 mV (vs. SSCE) was sufficient for the direct electrochemical determination of acetylcholine (or choline) without the need of any externally supplied electron transfer mediator. The feasibility of employing this type of probe in an in vivo environment was addressed by a kinetic analysis of the sensor's response. Since this analysis indicated that the overall response of these sensors was limited by the rate(s) of the catalytic enzyme(s), future work was proposed to employ a Nafion polymeric coating which would increase the range of linearity of these sensors by ensuring that their responses would be mass transport limited.

Kawagoe and co-workers constructed two miniaturized enzyme-modified organic conducting salt microelectrodes by electrodepositing TTF-TCNQ crystals in the recessed tip of a 7- pm-diameter carbon fiber microelectrode [98]. Acetylcholine esterase and choline oxidase were utilized to form a micro- biosensor for acetylcholine; this sensor was operated at an applied potential of 150mV (vs. SSCE) and exhibited a fast response time ( 3 s for 90% of the maximum response). In addition, glucose oxidase was used for the fabrication of a glucose microsensor. In order to minimize the interferences stemming from oxygen and ascorbate, a layer of ascorbate oxidase was applied over the GOD layer. In this configuration, the ascorbic acid oxidase enzymology was separated from the GOD/TTF-TCNQ response, and both interferences were minimized when the analyte solution was saturated with oxygen.

A very stable GOD-modified microelectrode was constructed by Boutelle et al. based on the TTF-TCNQ electrodes of Albery et al. [95]. No dramatic degradation of the sensor's response was observed following 6 h periods of use in brain tissue [99]. In brief, the electrodes were made by packing the TTF+-TCNQ- crystals into a 250-pm-diameter Teflon-insulated cavity where electrode contact was made with a silver microelectrode. Since the adsorption of GOD was sufficiently strong such that no additional immobilization membranes were necessary, the fast response of these sensors permitted their use for a physiologically relevant measurement of glucose dynamics in vivo (the 90% response to the stepwise addition of glucose was on the order of 20 s). The electrode potential was held at 0.0 V (vs. Ag/ AgCl) and the signal from ascorbate decreased following the adsorption of GOD onto the electrode surface as well as following the implantation of the probe into brain tissue. Additionally, a differential amperometric measurement was employed to negate any possible interferences from fluctuating levels of ascorbic acid in the measurement of glucose in the brain of a freely moving rat. In this configuration, these sensors were shown to be capable of reliably monitoring the changes in the levels of brain glucose following a local application of insulin.

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Enzyme-Modified Microelectrodes for In Vivo Neurochemical Measurements 41 1

3.6. Carbon Paste

The attractive feature of modified electrodes based on carbon paste may be ascribed to the ease of their preparation, short response time and to the reservoir-like property of the electrode matrix. Carbon paste is commercially available but most laboratories apply their own recipe for paste formulation, preparation, modification and electrode resurfacing. Indeed, inherent to their heterogeneous and pasty nature, modified carbon paste electrodes suffer from poor mechanical stability especially under dynamic flow conditions (i.e., there is risk of leaching of the mediator and/or the biocomponent out of the electrode matrix into the solution). Various strategies have already been developed to solve these problems (e.g., electrostatic trapping of the mediator into the paste, use of less soluble or hydrophobic mediators of high molecular weight, chemical immobilization of the entrapped bio- catalyst, covering the carbon paste surface with appropriate membranes, and the use of a robust graphite powder/polymer matrix).

Wang and co-workers have shown that the simultaneous incorporation of ferrocene derivatives and glucose oxidase into a carbon paste matrix results in an effective microelectrode for sensing glucose [loo]. The close proximity of the enzymatic, redox mediating and sensing sites offers extremely short response times (&o = 18 s) as compared to early ferrocene- based glucose sensors. In this work, the influence of paste composition, operating potential, glucose concentration and other variables was described. The incorporation of stearic acid into the enzyme-containing paste greatly reduced the interference due to ascorbic acid.

Glutamate dehydrogenase has been immobilized on glassy carbon and carbon paste electrodes by Amine and Kauffmann for the preparation of a glutamate biosensor [69]. The biosensor was characterized using hexacyanoferrate(II1) as the electrochemical mediator or phenazine methylsulfate as a redox mediator. Linear calibration plots were obtained between 50 pM and 1.3 mM glutamate. The electrode remained stable for about 11 days. The time response of these electrodes was approximately 12 s. The selectivity of the sensor, as well as most sensors which use a redox mediator, is somewhat compromised by the reactivity of the redox mediator for other easily oxidized species.

3.7. Polymers

The formation of polymer films on electrode surfaces and the use of polymer-film-coated electrodes for bioanalytical purposes has been known for some time [ l l , 17, 45, 101-1251. Electrochemically active polymers, which are intrinsically conductive towards both ions and electrons, can be divided into three basic groups. Redox polymers (e.g., polyvinyl ferrocene) contain redox sites or localized electronic states that are electronically conductive as a result of a process described as a “self-exchange” of electrons (i.e., electrons hop between oxidized and reduced states). Electronically conductive polymers [e.g., polypyrrole doped with Fe(CN):-] are organic metals that are more conductive than their redox polymer counterparts since the charge transfer process takes place through delocalized metallike band structures. Finally, ion exchange polymers (e.g., Nafion) are made electroactive by exchanging counterions for redox-active ionic species. Undoubtedly, the ease and variety of methods with which a variety of polymer films can be cast has certainly played a role in

the intense interest and rapid progress in this field. For example, preformed polymers can be dip-coated, spin-coated, screen- painted, evaporated, bound in a covalent manner, or simply physisorbed or chemisorbed onto a suitable surface. Addition- ally, a solution of monomer can be initiated to polymerize via methods of electrostatic precipitation, via potentiostatic or galvanostatic control, via a radio frequency plasma discharge, or via several sources of electromagnetic radiation [ 126-1281.

3.8. Electropolymerized Films

Since the original work of Diaz and co-workers [129], numerous refinements continue to appear in the literature concerning the methods by which conducting polymers can be synthesized. As the trend in the field of biosensors continues to be towards the development and characterization of miniaturized microsensors, few technologies have contributed more to the recent advances in this field than electronically conducting polymers. Many of the important aspects of this technology have been reviewed recently by Bartlett and Cooper [130]. The main advantage of these electronically conducting polymer films is that they can be manipulated on a molecular level in order to impart specific properties and functions to a modified electrode interface. The ability to control the immobilization conditions of an enzyme and/or electron transfer mediator has facilitated the successful localization of enzymes onto miniaturized electroactive surfaces. In conjunction with being able to choose a polymer on the basis of the state of its charge and/or permeability, the ability to reproducibly control the amount of polymer that is deposited (i.e., its thickness) allows some control over the temporal response and the concentration of immobilized enzyme.

For example, both the structure and the conductivity of poly- N-methylpyrrole (in addition to the activity of the enzyme entrapped in it) can be influenced by numerous experimental parameters such as the choice of solvent, the monomer concentration, the nature of the supporting electrolyte and its concentration, the pH, the temperature, and the amount of charge passed during the electropolymerization process [ 1 191. The only disadvantages of this immobilization method come from the variability in the adherence of some polymers to some electrode materials, and the incompatibility of some enzymes to certain polymeric matrices and/or applied potential waveforms [113].

Clearly the largest group of direct electron transfer biosensors is based on coimmobilization of an enzyme in a conducting polymer of polypyrrole (PPy) [102, 1041. These conductive polymer films can be regarded as porous organometallic complexes with a high surface-to-volume ratio and an average pore size of SO-lOOi%. However, while such features can be used to selectively repel interfering species, they can also influence the loss of enzyme where the rate of desorption depends upon the enzyme’s net charge [l 11. One of the reasons why considerable attention has been given to PPy/GOD- electrodes is because they facilitate the attachment of additional molecular redox species onto their surfaces (e.g., ferrocene modified PPy/GOD-electrodes) [ 1261. An alternative to polypyrrole is poly(pheno1) [131], which has been used to entrain GOD onto 125- and 25-pm-diameter platinum microelectrodes. The poly(pheno1) was electrodeposited after GOD was adsorbed to the electrode surface, resulting in a very stable electrode (> 40 days) with a reasonably fast response time (< 10 s).

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3.9. Selective Polymer Layers

Another reason for the widespread usage of polymer- modified biosensors stems from their ability to simultaneously perform multiple tasks. For example, they can serve as a matrix for the preconcentration of an analyte(s), as a matrix that serves as a transport barrier for detrimental interferences or as a support for the immobilization of one or more enzymes (with or without an electron transfer mediator) [lll]. In an in vivo measurement, one of the most severe limitations of electrochemical detection schemes is the interference originating from the oxidation of endogenous compounds that are easily oxidizable. This is especially true for ascorbic acid since its concentration is quite large and it is well-known to undergo dramatic fluctuations in response to stimulated neuronal activity [132, 1331.

A miniaturized glucose oxidase-based needle-type glucose microsensor was developed for subcutaneous glucose monitoring [134]. The sensor is equivalent in shape and size to a 26- gauge needle (0.45 mm i.d.), where GOD and polymer films were deposited so that sensors with characteristics suitable for in vivo use could be prepared in high yield (greater than 60%). The upper limit of the linear range was greater than 15mM, the response time was less than 5 min, and sensitivity yielded a 5 : 1 signal-to-background ratio at normal basal glucose levels. The sensor response was largely independent of oxygen tension in the normal physiological range. Acetaminophen has been one of the most serious electrochemical interferences to oxidase-based amperometric biosensors that measure H202. A composite membrane of cellulose acetate and Nafion was found to eliminate acetaminophen and other electrochemical interferences effectively while at the same time maintaining reasonable diffusivity for hydrogen peroxide [ 1351. The excellent in vivo performance of the sensor was attributed not only to significantly reduced steady-state sensitivity to acetaminophen but also to a very slow acetaminophen response. These features, combined with rapid acetaminophen clearance pharmaco- kinetics, led to the decreased response as demonstrated in the rat.

The electrochemical polymerization of a selective polymer layer has found a great deal of success for its ability to reject a number of interferences without dramatically changing the time response of the biosensor. For example, Malitesta et al. used o-phenylenediamine to electrochemically polymerize GOD in a strongly adherent, highly reproducible membrane, where the thickness of the layer was on the order of lOnm [136]. This one- step procedure produced a glucose oxidase-modified platinum electrode with a response time less than 1 s and a wide linear range. Moreover, as a result of the permselectivity of the membrane, the access of ascorbate to the electrode surface was effectively minimized.

3.10. Photochemical Initiation of Polymerization

Photochemical techniques for the entrapment of bioactive materials are of particular interest owing to their ease of employment, and the mild conditions that are required to generate a polymeric network [ l l S , 137-1411. For example, consider the photo-activated cross-linkable method utilizing polyvinyl alcohol-styryl pyridinuim (PVA-SBQ) that was developed by Ichimura [142]. Since the PVA-SBQ method is rapid and does not change the pH or temperature of reactants/ reagents, most enzyme structures are hardly affected by this method of immobilization. The most significant advances and

developments with this technology for application in the field of micro-biosensors have come from the laboratory of Karube and co-workers. In a series of papers, they have reported on the use of PVA-SBQ as an immobilization agent for enzymes on various microelectrode surfaces selective for the neurochemically interesting molecules, acetylcholine and glutamate [ 143-1461. Navera et al. have fabricated a ZOO-pm-diameter acetylcholine sensor where a carbon fiber microelectrode was used to detect the enzymatic production of H202 at 1.2 V (vs. Ag/AgCl) from the sequential enzyme reactions of acetylcholinesterase and choline oxidase [ 1421. In this initial work, the co-immobilization of enzymes in PVA-SBQ was demonstrated where the response time of the sensor was less than one minute. Dramatic improvements in the performance characteristics of this sensor were reported by Tamiya et al. who utilized a 500-pm- long, 7-pm-diameter carbon fiber cylinder microelectrode as the electrochemical transducer/enzyme support for this acetylcholine sensor [144, 1451. In this work, a complex potential waveform was used to increase the sensitivity of the hydrogen peroxide determination and to improve the selectivity with respect to ascorbic acid. In addition, this waveform was shown to be effective in minimizing the effects of protein fouling; presumably, the production of oxygen and hydrogen gases by this triangular potential waveform etched the carbon surface and therefore stripped adsorbed protein from the electrode surface. Remarkably, the deactivation of the immobilized catalytic proteins immobilized within the PVA-SBQ matrix was not as dramatic as might be expected; the half-life of these sensors was on the order of one week. Finally, the thickness of the enzyme layer was decreased and a corresponding reduction was observed in the measurement time for acetylcholine (this was decreased to 5 s).

A variation of this immobilization scheme was utilized by Tamiya et al. to construct a microamperometric sensor for the determination of glutamate [145, 1461. In this work, the carbon fiber disk electrode (7 pm diameter) was platinized to increase the sensitivity of a pulsed amperometric determination of hydrogen peroxide at 600mV (vs. Ag/AgCl). In this work, glutamate oxidase and bovine serum albumin were premixed before being immobilized to the platinized surface by PVA- SBQ, and the photo-initiated polymerization was followed by treatment of the layer with gluteraldehyde. The current response of these sensors to a stepwise addition of glutamate was very stable and the response times were on the order of 12 s. While this sensor was used to show the feasibility of monitoring the potassium stimulated change in the concentration of a species in cerebellar cortex, the authors did not address how the signal from glutamate could be differentiated from the contributions to the signal stemming from ascorbate and the capacitive changes that will accompany the injection of a large concentration of potassium ion [1451. A recent improvement to the design of this sensor involves the elimination of the traditional Ag/AgCl reference electrode through the use of a platinized-coated glass capillary tube such that the totally integrated sensor has a tip diameter on the order of 20pm [146].

3.11. Redox Polymers

An idea that has found widespread usage is that of entrapping redox enzymes in redox polymers, which can function as both the mediator and enzyme immobilization matrix. The “redox conducting epoxy cement” of Gregg and Heller is among the more novel redox films of recent years [147]. These films

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Enzyme-Modified Microelectrodes for In Vivo Neurochemical Measurements 41 3

are formed from a two-component mixture containing an osmium-bipyridine complex and a commercially available poly(ethy1ene glycol) diglycidyl ether. The resulting films are strongly attached to electrode surfaces and highly permeable to solution species. The glucose response time of the resulting electrodes is less than 10 s. Unfortunately, the responses were decreased by 30-50% in air-saturated solutions because of competition between oxygen and the 0s”’ complex for electrons from the reduced enzyme.

Glucose microelectrodes have been fabricated by Pishko and co-workers who immobilized GOD in a poly((viny1pyridine) Os(bipyridine),Cl) derivative-based redox hydrogel onto the surface of carbon fiber disk microelectrodes (7 pm diameter) [148]. The current density and sensitivity are 10 times higher than with macroelectrodes made with the same hydrogel. Furthermore, the current is less affected by a change in the partial pressure of oxygen. The higher current density and lower oxygen sensitivity point to the efficient collection of electrons through their diffusion in the redox hydrogel to the electrode surface. These results contrast with those observed for enzyme electrodes based on diffusing mediators, where loss of the enzyme-reduced mediator by radial diffusion to the solution decreases the current densities of microelectrodes relative to similar macroelectrodes. The response times of these sensors were on the order of 5 s (10-90% of the maximum signal).

Michael and co-workers used the same principles to produce amperometric sensors for peroxide, choline, and acetylcholine [ 1491. Horseradish peroxidase (HRP), choline oxidase, and acetylcholinesterase were immobilized in a cross-linked redox polymer deposited on glassy carbon electrodes. Peroxide sensors prepared by immobilization of HRP alone, gave detection limits of lOnM and a linear response up to ca. 1 mM. Coimmobilization of HRP and glucose oxidase was used to establish the feasibility of highly efficient bienzyme sensors at low substrate levels. Replacing glucose oxidase with choline oxidase produced sensors with submicromolar detection limits and a linear response up to 0.8mM choline. These electrodes were used to monitor choline transients in vivo in the rat brain following local injection of choline near the electrode [ 1501. Addition of acetylcholinesterase to the sensors generated a relatively small response to acetylcholine that demonstrates the feasibility of trienzyme sensors. At low substrate concentrations, no loss in sensitivity during a I-day experiment was observed. The response times of these sensors are all less than 30s with 2 s response times achieved in some cases.

3.12. Electrical Wiring of Redox Enzymes

Degani and Heller carried out the first direct electrical communication between glucose oxidase and an electrode surface. These researchers modified the enzyme structure by opening the active site using urea, thus making it more accessible to arrays of electrode-transfer relays comprising of a ferrocyl acetamide-modified electrode [ 151, 1521. This chemical modification of the enzyme has led to the production of “electrically wired” enzymes [ 1531. Under normal circum- stances, the transfer of electrons between the active site of reduced GOD and an electrode takes place slowly or not at all. This is because the distance between the active site and the electrode is too large for the electrons to cross. Attachment of an electron mediator to the native enzyme can open a direct path of electrical communication between the metal electrode and the enzyme redox center.

In order to improve the applicability of redox polymer- modified enzyme electrodes in biological systems, Maidan and Heller constructed a multilayer electrode that consisted of an immobilized horseradish peroxidase (HRP) film coated onto a film of electrically “wired” GOD [154]. The “wired” enzyme was connected by a redox epoxy network to a vitreous carbon electrode. In this configuration, the common electrooxidizable interferants (ascorbic acid, uric acid and acetaminophen) that normally interfere with amperometric glucose assays are completely and rapidly oxidized by horseradish peroxidase in a multilayer electrode. The current from the electrooxidizable interferants was decreased by a factor of 2500 by the peroxidase- catalyzed preoxidation, and the glucose-interference current ratio was increased 1000-fold.

3.13. Covalent Derivatization

While covalent methods of enzyme immobilization are more labor intensive than those involving the mere adsorption of enzyme (i.e., steps must be taken to choose and prepare the surface, reagents and reaction conditions), covalent methods are also more widely applicable as a result of increased stability of the enzyme and/or mediator once bound. Covalent immobilization can produce an extremely thin (ca. monolayer), perma- nently bound, enzyme layer in a reproducible fashion. Enzymes are usually anchored to an insoluble support by covalent attachment via an intermediate linkage. The variety of linkages available make immobilization attractive, provided the support surface can be functionalized. Carbon and graphite electrodes can be conveniently functionalized to serve as amperometric enzyme electrodes. Yacynych and co-workers were the first to demonstrate the covalent immobilization of enzymes onto macroscopic carbon electrode surfaces [ 155, 1561. While various methods of covalent immobilization of enzymes have been investigated at macroscopic electrodes, there has been little success of the transfer of this technology to microelectrode surfaces.

Considerable work has been done on the construction of an enzyme-modified carbon fiber microelectrode using a covalent attachment of the enzyme to the carbon surface through a hydrophilic tether which employs biotin-avidin technology [157-160]. The interaction between one molecule of the egg white protein avidin and four molecules of biotin (vitamin H) is one of the strongest known noncovalent biological recogni- tion’s between a protein and ligand (Kd = M) [161]. In association with the covalently linked tether, the biotin-avidin “molecular sandwich” attaches the enzyme close (within 100 .$) to the carbon surface. This derivatization scheme permits a great deal of control over the extent of coverage and localization of the enzyme across the electrode surface and can do so without creating any barriers to diffusion which can limit temporal responses.

The biotin/avidin-modified surface not only provides a stable environment for the enzyme, but can do so with minimum interference to the electron transfer properties of oxidase-based redox mediators [ 1581. For example, the electrochemical properties of a horseradish peroxidase and a glucose oxidase- modified microelectrode were examined via flow injection analysis experiments, and each showed response times on the order of 200ms, with a very high current density (64mA/cm2, 20 mM /3-glucose).

The temporal resolution of dehydrogenase-modified electrodes is comparable to that observed at the oxidase-modified microelectrodes, also showing subsecond response times, and

Electroanalysis 1995, 7 , No. 5


several dehydrogenases (including glutamate, alcohol and glyceraldehyde 6-phosphate dehydrogenase) were biotinylated and successfully attached to the electrode surface with a biotin/avidin tether [I 59, 1601. Unfortunately, the alteration of the carbon fiber surface by the derivatization procedure produces a surface less optimal for NADH oxidation, which promoted an investigation into the effects the derivatization had on the chemical architecture of the carbon fiber surface [ 162-1641, Thus, the optimization of this derivatization requires careful characterization of the chemical properties of the surface after each derivatization step. This is because the carbon fiber surface is the site of both enzyme attachment and electron transfer; in other words, the electrode surface must be modified in such a way to provide a stable environment for the enzyme without interfering in electron transfer to the redox cofactor. A fluorescent probe (FITC- labeled avidin) was attached to the derivatized carboxyl groups across the carbon fiber surface via the biotinylated Jeffamine tether to characterize the effects of surface derivatization [ 1621. The fluorescence observed from this modified surface after excitation at 488 nm was imaged with a fluorescence microscope equipped with a cooled CCD camera, yielding a spatial map of the distribution of derivatized carboxyl groups on the surface of the carbon fiber with 50.5pm resolution. Additionally, the electrochemically generated chemiluminescence (ECL) of luminol was imaged across the carbon fiber surface to determine the spatial distribution of the derivatized electrode's electron transfer sites [ 1631. Luminol ECL imaged at the electrode surface was compared to the spatial distribution of FITC-ExtrAvidin- modified carboxylates indicated by the fluorescence image. The combination of these techniques at the same electrode allowed the distribution of derivatized carboxyl groups to be correlated with the distribution of electroactive sites for the oxidation of luminol.

These glutamate-selective microelectrodes have been used for the measurement of the dynamics of glutamate release in the mammalian brain on the millisecond time scale [165]. Fast scan cyclic voltammetry, where the voltammetric information is obtained in a few milliseconds, circumvents many of the drawbacks of traditional superfusion detection methods and has been proven ideal for the study of monoamine release in brain slices from experimental animals. 'The outstanding temporal resolution is matched by the comparatively high spatial resolution afforded by the microelectrode. Brain slices (350 pm thick) were prepared for perfusion in a conventional manner containing saggital sections of the hippocampus. One micropositioner is used to place two tungsten stimulating electrodes 150 pm apart; another micropositioner is used for the placement of the 32pm carbon fiber microelectrode at a position approximately 150 pm from either stimulating electrode. Signal was observed as a function of time following a 100Hz, 4s local stimulation of the CAI region of the hippocampus. The selectivity of the enzyme-generated response could be demonstrated easily. When the glutamate response was desired, the brain slice was continuously perfused with ACSF that contained the enzyme cofactor (3mM NAD'). The presence of an excess of the enzyme cofactor (NAD') turned on the enzyme (glutamate dehydrogenase) and allowed the formation of NADH as a function of glutamate concentration. The nonspecific response was obtained by omitting the NAD+ from the perfusion buffer for approximately 1 h prior to recording. Application of an identical stimulation did not produce a signal, indicating that this response was enzymatic in nature.

4. Conclusions

Many significant advances have allowed the construction of enzyme-modified microelectrodes that have the requisite characteristics necessary to make physiologically relevant measurements in vivo. The dominant trend in biosensor development involves the progression towards improving the sensitivity and long-term stability of miniaturized chemical sensors since there is a great need for these sensors in many neurochemical and clinical investigations [51]. Advances in sensitivity, stability and most importantly, the response times of these sensors will allow the measurement of many dynamic processes in the chemical environment necessary to characterize biochemical processes which govern physiology and behavior.

5. Acknowledgements

This work was supported by the National Institutes of Health (Grant GM44112-01). W.G.K. is the recipient of a Presidential Young Investigator Award (NSF CHE-8957394).

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