Design variations of a polymerenzyme composite biosensor for glucose: Enhanced analyte sensitivity without increased oxygen dependence

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    Available online 3 May 2005

    redox center of the enzyme and the electrode surface

    [7,8]. Using the avoenzyme glucose oxidase (GOx), the

    GOx/FADH2 +O2!GOx/FAD+H2O2 2

    H2O2!O2 +2H +2e 3where FAD is the oxidized form of the prosthetic group,avin adenine dinucleotide. Reaction (3) represents the

    .

    * Corresponding author. Tel.: +353 1 7162314; fax: +353 1 7162127.

    E-mail address: Robert.ONeill@UCD.ie (R.D. ONeill).

    Journal of Electroanalytical Chemist

    Journal ofElectroanalytical0022-0728/$ - see front matter 2005 Elsevier B.V. All rights reserved1. Introduction

    The incorporation of biological molecules into sensing

    devices is an important strategy in the development of

    fast, ecient and inexpensive assays for a wide range of

    analytes in clinical, industrial and environmental applica-

    tions [16]. In particular, amperometric biosensors fabri-cated using immobilized enzymes can be categorized

    broadly as rst, second or third generation devices

    depending on the mode of electron transfer between the

    majority of these sensors have focused on the detection

    of glucose, both as a model system and because of the

    importance of glucose as a target analyte [1,911].

    The overall enzymatic process for GOx, that is simi-

    lar for many oxidases, may be written as Reactions (1)

    and (2)

    b-d-glucose +GOx/FAD

    !d-glucono-d-lactone +GOx/FADH21Abstract

    The glucose sensitivity and oxygen dependence of a variety of implantable biosensors based on glucose oxidase (GOx), incorpo-

    rating an electrosynthesized poly-o-phenylenediamine (PPD) permselective barrier on 125-lm diameter Pt disks (PtD) and cylinders(PtC, 1-mm length), were measured and compared. Full glucose calibrations and experimental monitoring of solution oxygen con-

    centration allowed us to determine apparent MichaelisMenten parameters for glucose and oxygen. In the linear region of glucose

    response, the most sensitive biosensor design studied was PtD/PPD/GOx (enzyme deposited over polymer) that was 20 times moresensitive than the more widely used PtC/GOx/PPD (enzyme immobilized before polymer deposition) conguration. The oxygen

    dependence, quantied as KM(O2), of both active and less active designs was surprisingly similar, a nding that could be rationalized

    in terms of an increase in KM(G) with increased enzyme loading. The PtD/PPD/GOx design will now enable us to explore glucose

    concentration dynamics in smaller and layered brain regions with good sensitivity and minimal interference from uctuations in

    tissue pO2.

    2005 Elsevier B.V. All rights reserved.

    Keywords: Poly(o-phenylenediamine); Enzyme modied electrode; Amperometry; Glucose oxidase; Electropolymerization; Hydrogen peroxide;

    Brain monitoringDesign variations of a polymfor glucose: Enhanced analy

    oxygen d

    Colm P. McMahon, Sarah J

    Chemistry Department, University C

    Received 20 January 2005; received in revisedoi:10.1016/j.jelechem.2005.03.026enzyme composite biosensorsensitivity without increasedendence

    illoran, Robert D. ONeill *

    e Dublin, Beleld, Dublin 4, Ireland

    m 22 March 2005; accepted 24 March 2005

    www.elsevier.com/locate/jelechem

    ry 580 (2005) 193202

    Chemistry

  • 194 C.P. McMahon et al. / Journal of Electroanalytical Chemistry 580 (2005) 193202electrochemical step and in rst generation devices is

    generally carried out amperometrically, either directly

    on the substrate surface at relatively high applied poten-

    tials [12,13], or catalytically at lower potentials [14,15].

    Thus, a two-substrate model is necessary to describe

    the kinetics of GOx under conditions of varying concen-tration of both glucose and O2 [16]. When the concen-

    tration of the co-substrate is constant, however, the

    two-substrate equation simplies to the one-substrate

    MichaelisMenten form (Eq. (4)), where the current

    density for the biosensor glucose response, Jgluc, is a

    measure of the overall rate of the enzyme reaction,

    and Jmax is the Jgluc value when the surface enzyme is

    saturated with glucose (G). Dierent values of Jmax,determined under the same conditions, reect dierences

    in the amount of active enzyme on the surface provided

    the sensitivity of the electrode to H2O2 (Reaction (3))

    does not vary, as is the case for the PPD-modied Pt cyl-

    inders and disks used here [17,18]

    J gluc Jmax1 KMGG

    : 4

    The true Michaelis constant for the one-substrate

    case, KM, is dened (Eq. (5)) in terms of the rate con-

    stants for the generalized reactions describing the overall

    conversion of substrate to product, catalyzed by the en-zyme: substrate binding to enzyme (k1) and its reversal

    (k1), and the conversion step itself (k2 or kcat) [19,20]

    KM k1 k2k1 : 5

    However, when Eq. (4) is used to approximate the

    two-substrate case (e.g., glucose and oxygen), the

    KM(G) is more complex, containing co-substrate terms.

    KM(G) is then the apparent Michaelis constant and phe-

    nomenologically denes the concentration of substrate

    that gives half the Jmax response. If the concentrationof co-substrate is high and constant (as was the case in

    the rst part of this study), the main factors aecting

    changes in KM(G), for glucose biosensors of similar de-

    sign, are those given in Eq. (5). Thus, the value of

    KM(G) is sensitive to the binding constant, k1, and has

    often been interpreted in terms of barriers to substrate/

    enzyme binding [21,22].

    Alternatively, if the concentration of glucose is xedand O2 levels are changed, then Eq. (6) can be used to

    analyze the oxygen dependence of the glucose signal

    [23], where J 0max is the maximum (plateau) response fora particular concentration of glucose, and KM(O2) is

    the apparent Michaelis constant for oxygen. The option

    of using a single two-substrate equation, such as Eq. (32)

    in [16] or Eq. (1) in [24] that expresses explicitly the true

    Michaelis constants for both substrate and co-substrate,was not used in this analysis because the apparentMichaelis constants dened separately are analyticallymore straightforward, are useful for dening the linear

    range for glucose (1/2KM(G), Eq. (4)) and oxygen(1/2KM(O2), Eq. (6)) responses, and inuence the glu-cose slope in the linear region, i.e., Jmax/KM(G). Inaddition, the concentration of surface enzyme determin-

    able in studies involving drop evaporation onto macro-disk electrodes [24] is not available for enzyme deposited

    by dip evaporation, as used here for microdisks and cyl-

    inders; this also limits the scope of the kinetic analysis of

    the individual reaction steps

    J gluc J0max

    1 KMO2O2: 6

    The three main problems associated with reliable glu-

    cose detection using rst generation amperometric bio-

    sensors are: analyte sensitivity (Reactions (1)(3)); the

    inuence of changes in pO2 on the biosensor signal(Reaction (2)); and interference by electroactive compo-

    nents present in the target medium (Reaction (3)). A

    variety of strategies have been developed to optimize

    the sensitivity, selectivity and stability of glucose biosen-

    sors [1]. Recent novel approaches to the ecient detec-

    tion of GOx-generated H2O2 include the use of

    platinized platinum cylinders [25], platinized boron-

    doped diamond microbers [26], iron-loaded carbonnanotubes [27], and horseradish peroxidase coupled to

    redox polymers [28]. Targeted synthesis of enzyme

    immobilization agents [29], incorporation of oxygen

    storage media into the biosensor design [30], and the

    use of inventive enzyme entrapment procedures [31]

    have also led to advances in achieving these goals. The

    choice of design for a given application, however, should

    depend on the conditions, such as levels of glucose pres-ent, properties of specic electroactive interference, and

    the relevant range of pO2.

    A number of laboratories have described the develop-

    ment and characterization of a rst generation glucose

    sensor based on aGOx immobilized in poly(o-phenylene-

    diamine) on Pt electrodes (PtGOxPPD) that possesses

    a variety of properties indicating suitability for neuro-

    chemical applications [22,3237]. These attributes in-clude fast response time, linearity over the relevant

    concentration range, eective elimination of interference

    by endogenous reducing agents such as ascorbic acid,

    freedom from protein and lipid fouling, stability in vivo,

    and ease of miniaturization [38]. More recently, we have

    demonstrated that a particular conguration of the Pt

    GOxPPD design also displays minimal O2 interference

    in applications both in vitro and in vivo over limited glu-cose and O2 concentration ranges [23]. In this paper we

    systematically explore variations in the PtGOxPPD

    design, in attempts to increase glucose sensitivity with-

    out compromising the range of pO2 over which these

    biosensors can operate reliability. Designs involvingthe incorporation of enzyme in the monomer solution,

  • environments to perform all amperometric experiments

    troanaand to collect, plot, and do a preliminary analysis ofthe data.

    2.3. Working electrode preparation

    The basic protocol for fabrication of the working

    electrode biosensors has been described in detail recently

    [47,48], and a number of variations are reported here.

    Cylinders (1 mm long) of freshly cut Teon-insulatedPt wire (125 lm diameter, Advent Research Materials,a strategy that has been used frequently for PPD-based

    biosensors [22,25,33,34,3946], were avoided here be-

    cause dip-evaporation deposition of the enzyme prior

    to electropolymerization has been shown to give good

    eciency of enzyme loading compared to the co-

    deposition of enzyme and polymer [13], and to enableus to develop protocols that could be used formore expen-

    sive enzymes, such as glutamate oxidase, where co-immo-

    bilization would not be a cost-eective option [13].

    2. Materials and methods

    2.1. Chemicals and solutions

    The enzymes glucose oxidase (GOx) from Aspergillus

    niger (EC 1.1.3.4, type VII-S), was obtained from Sigma

    Chemical Co, as was o-phenylenediamine (1,2-diamino-

    benzene), a-D-(+)-glucose and bovine serum albumin(BSA, fraction V). The L-ascorbic acid (AA, BDH Bio-

    chemical grade) and all other reagents were used as sup-

    plied. A stock 1 M solution of glucose was prepared inwater, left for 24 h at room temperature to allow equil-

    ibration of the anomers, and then stored at 4 C. Allexperiments were carried out in a phosphate-buered

    saline (PBS) solution, pH 7.4; NaCl (BDH, AnalaR

    grade, 150 mM), NaH2PO4 (BDH, AnalaR grade, 40

    mM) and NaOH (Sigma, 40 mM). A 200 U/mL solution

    of GOx was prepared by dissolving 15 mg in 10 mL of

    PBS. The 300 mM o-phenylenediamine (o-PD) mono-mer solution was prepared using 0.81 g of o-PD and

    125 mg of BSA in 25 mL of PBS and sonicating at 25

    C until dissolved.

    2.2. Instrumentation and software

    Experiments were computer controlled with data col-

    lection accomplished using either a Biodata Microlinkinterface or a National Instruments (NI, Austin, TX)

    AT-MIO-16 data acquisition board linked to a low-

    noise, low-damping potentiostat (Biostat II, Electro-

    chemical and Medical Systems, Newbury, UK).

    In-house software was written in QuickBASIC (version

    4.0) and NI LabWindows (version 2.1) QuickBASIC

    C.P. McMahon et al. / Journal of ElecSuolk, UK), or disks made from the same wire, weredipped into a 200 U/mL solution of GOx to allow

    adsorption of the enzyme (5 min), and then removed

    to dry. This procedure was repeated four more times, ex-

    cept that the electrodes were removed immediately fol-

    lowing immersion. This dip-evaporation procedure

    optimizes the loading of GOx on the Pt wire [13]. Someof these Pt/GOx electrodes were exposed to glutaralde-

    hyde (GA) vapor for 20 min to cross-link the enzyme;

    these enzyme-modied cylinders and disks are termed

    PtC/GOxGA (geometric area, 4.05 103 cm2) andPtD/GOxGA (1.23 104 cm2) biosensors, respec-tively. Other samples of the enzyme-coated wire were

    immersed in PBS containing the o-PD monomer and

    BSA, and electropolymerization carried out at +700mV vs. SCE for 15 min to produce PtC/GOx/PPD

    BSA and PtD/GOx/PPDBSA sensors. Finally, a set of

    biosensors were fabricated by dip-evaporation of GOx

    onto polymer-modied electrodes followed by cross-

    linking with GA, i.e., PtC/PPDBSA/GOxGA and

    PtD/PPDBSA/GOxGA. The BSA was included in

    the polymerization medium because this, and other pro-

    teins [17], have been shown to increase the interferencerejecting properties of the PPD polymer [49,50].

    A self-calibrating commercial membrane-covered

    amperometric oxygen sensor (1 cm diameter) was usedto quantify solution oxygen concentration. The model

    used was a CellOx 325 connected to an Oxi 340A meter

    (Wissenschaftlich-Technische Werkstatten GmbH from

    Carl Stuart Ltd., Dublin, Ireland), incorporating a tem-

    perature probe for automatic compensation. Reliablequantication of O2 using this device required constant

    stirring of the solution at a rate of 5 Hz using a glasspropeller mounted parallel to the shaft of the CellOx

    325. The sensor range was 0.0199.9% O2 (100% corre-

    sponding to air saturation) with a resolution of 0.1%.

    This percentage was converted to an estimated concen-

    tration of O2, taking 200 lM to correspond to 100%[51,52].

    2.4. Amperometric experiments

    All calibrations were performed in a standard three-

    electrode glass electrochemical cell containing 20 mL

    PBS at room temperature (22 C). A saturated calomelelectrode (SCE) and a large stainless steel needle served

    as the reference and auxiliary electrodes, respectively.The applied potential for amperometric studies was

    +700 mV vs. SCE. Glucose calibrations were performed

    in quiescent air-saturated PBS by adding aliquots of glu-

    cose stock solution to the electrochemical cell, and stir-

    ring for 5 s. In oxygen dependence experiments, the

    electrochemical cell was contained within an Atmos-

    bage (Sigma) to avoid contamination by air [23]. Theaddition of glucose was not carried out until the elec-trodes were well settled giving a steady background in

    lytical Chemistry 580 (2005) 193202 195air-saturated PBS. After the addition of a single aliquot

  • of dierences observed between the glucose and oxy-

    been used in the past to probe barriers to substrate/en-

    zyme binding [21,22]. Finally, the value of KM(O2),

    which is the concentration of oxygen needed to give half

    the maximum biosensor response for a given xed con-

    centration of glucose (Eq. (6)), provides an index of the

    oxygen dependence of the immobilized enzyme as afunction of glucose concentration for each design of

    0 20 40 60 80 1000

    3

    KM(G)

    J glu

    c

    0 20 40 60 80 1000

    10

    20

    30

    40

    50

    60

    70

    2. PtC/GOx/PPD-BSA

    1. PtC/GOx-GA

    3. PtC/GOx/PPD-BSA/GOx-GA

    [glucose] / mM

    [glucose] / mM

    J glu

    c /

    A c

    m-2

    0 1 2 3 4 50.0

    0.2

    Fig. 1. The eects of electrode geometry and enzyme distribution on

    the glucose response of GOx-based biosensors. Points are means

    SEM of current densities; curves were generated using non-linear

    regression and Eq. (4) (R2 > 0.95). (Top) The geometric eect is

    illustrated by glucose calibrations performed using cylinder- and disk-

    based biosensors of the type where the GOx was adsorbed on the Pt

    before electropolymerization of the PPDBSA composite. Although

    the Jmax values were greater for the smaller electrodes (see Table 1), the

    KM for glucose, KM(G), was not signicantly dierent for these two

    designs: 15 2 mM (, n = 21); 15 1 mM (, n = 4); p > 0.94. (Inset)Part of the linear region glucose response for the PtC/GOx/PPDBSA

    conguration, showing the regression line with R2 = 0.995 and using

    the same axes as the main graph. (Bottom) The eect of enzyme

    distribution was demonstrated on cylinder electrodes by sequentially

    depositing GOx, and calibrating at each stage: rst on the bare Pt (1.

    PtC/GOxGA); the PPDBSA matrix was then generated (2. PtC/GOx/

    PPDBSA); nally, GOx was again deposited over the polymer layer

    (3. PtC/GOx/PPDBSA/GOxGA). MichaelisMenten parameters for

    larger populations of biosensors are given in Table 1.

    roanagen responses for the various designs was calculated

    using Students two-tailed unpaired t tests on theabsolute current densities, slopes or MichaelisMenten

    parameters.

    3. Results and discussion

    A number of sophisticated mathematical models of

    the behavior of enzymes immobilized in surface layers

    have been described [20,5356]. These complex analy-

    ses are often needed to understand and optimize the

    behavior of thick and/or conducting lms [57], andhave been used, for example, to determine the site

    of H2O2 oxidation in polypyrrole-based biosensors

    [39]. However, a recent study has shown that substrate

    diusion is not limiting for PPD layers incorporating

    GOx, due to their small thickness [25]. Therefore,

    the basic MichaelisMenten parameters used here pro-

    vide more readily accessible insights into factors aect-

    ing the responsiveness of biosensors fabricated fromultra-thin (1030 nm [34,49,58,59]) insulating poly-

    mers, such as PPD.

    Thus, dierences in Jmax values (Eq. (4)), determined

    under the same conditions, reect dierences in the

    amount of active enzyme on the surface, provided the

    sensitivity of the electrode to H2O2 (Reaction (3)) does

    not vary. Changes in the apparent Michaelis constant

    for glucose (KM(G), Eq. (4)), in the presence of a highand constant concentration of oxygen, are sensitive toof glucose, the bag was sealed and high-grade N2(

  • 3.1. Glucose response of a cylinder- and disk-based

    polymerenzyme conguration

    The use of PPD as a permselective polymer in biosen-

    sor applications is well documented (see Section 1). Two

    signicant advantages of PPD are its high permeabilityto the oxidase transduction molecule, H2O2 [17,49,60],

    and its ultra-thin nature when formed in neutral media

    that allows it to immobilize enzymes eciently

    [22,33,34,61]. Moreover, for neurochemical monitoring

    in vivo, the PPDenzyme matrix is stable over weeks

    Fig. 1 (top). The sensitivity of PtC and PtD electrodes

    to H2O2 was therefore measured as the slope of the cal-

    ibration plot (0100 lM) that was linear in both cases.

    en parameters, Jmax and KM(G), determined using non-linear regression and

    calibration slopes in the linear region, for biosensors of dierent designs

    cm2) KM(G) (mM) Slope (lA cm2 mM1)

    .8 14.7 2.0 0.42 0.09

    .3 6.4 0.3 3.0 0.3

    .0 5.0 1.4 6.6 1.9

    4 9.0 1.4 20.5 1.6

    .1 15.0 1.0 0.82 0.18

    Fig. 2. Schematics representing the biosensor design themes explored

    in this study. Immobilization of enzyme before (top left) or after

    (bottom left) deposition of the polymer (PPD) by electrosynthesis on

    either 125-lm diameter Pt cylinder (PtC, 1-mm length, top right) ordisk (PtD, bottom right) electrodes. Enzyme was immobilized on top of

    the polymer using GA vapor. Polymer thickness and enzyme diameter

    are roughly to scale. For clarity, the polymer modier (BSA) has been

    omitted from the scheme. Glucose oxidase (GOx) biosensors based on

    all four permutations of these two themes were characterized: PtC/

    GOx/PPDBSA; PtC/PPDBSA/GOx; PtD/GOx/PPDBSA; and PtD/

    PPDBSA/GOx.

    C.P. McMahon et al. / Journal of Electroanalytical Chemistry 580 (2005) 193202 197of continuous recording [48], and its excellent ascor-

    bate-rejecting properties are a feature of PPD that has

    allowed unambiguous detection of glucose under a vari-ety of physiological and pharmacological conditions in

    relatively large brain structures, such as the corpus stri-

    atum dorsalis [23,48,62]. For studies of smaller brain

    regions, and in other applications where miniaturization

    is desirable, the cylinder geometry is not suitable [18].

    We have therefore investigated the properties of a range

    of PPD-based glucose biosensors, specically the eects

    of decreasing electrode size and increasing GOx loadingon analyte sensitivity and oxygen dependence.

    The behavior of each biosensor design was character-

    ized by a full calibration in quiescent air-saturated buf-

    fer, and analyzed in terms the MichaelisMenten

    constants (Jmax and KM(G), Eq. (4)) and the slope of

    Jgluc in the linear region (LR) of glucose response. The

    rst factor examined was geometry. Fig. 1 (top) shows

    averaged calibrations for one cylinder and one diskdesign, both with the same enzyme conguration (poly-

    merization following enzyme immobilization): PtC/GOx/

    PPDBSA and PtD/GOx/PPDBSA. Clearly, the Jmaxfor the disk-based electrode was signicantly greater

    than that for the cylinder with the same polymer

    enzyme arrangement (see Table 1 for calibration

    parameters).

    Before using Jmax as an index of active enzymeloading when dierent geometries and scales were in-

    volved, it was necessary to determine whether subtle dif-

    ferences in mass transport to the PtC and PtD electrodes

    (see Fig. 2), or a disparity between the condition of the

    cylinder and freshly cut disk surfaces, might also con-

    tribute to the variations in Jmax values observed in

    Table 1

    Means SEM (n = number of biosensors) of apparent MichaelisMent

    Eq. (4) for glucose calibrations (see Fig. 1), together with the glucose

    calibrated at +700 mV vs. SCE in air-saturated buer (pH 7.4)

    Design n-value Jmax (lA

    PtC/GOx/PPDBSA 21 6.0 0

    PtC/PPDBSA/GOxGA 10 29.0 3

    PtC/GOxGA 6 35.8 7

    PtD/GOxGA 28 249 3

    PtD/GOx/PPDBSA 4 15.8 3PtD/PPDBSA/GOxGA 4 128 33 7.1 1.1 10.5 1.7

  • roanaThere was no dierence between the two slopes (PtC,

    225 15 nA cm2 lM1, n = 23; PtD, 220 10 nAcm2 lM1, n = 51), and both were signicantly higherthan that observed for H2O2 at Pt macrodisk electrodes

    [37], indicating that radial diusion is equally ecient

    for PtC and PtD electrodes of these dimensions overthe long time scales pertinent to these measurements.

    Thus, given the similarity of H2O2 sensitivity for the

    two geometries, the Jmax values in Fig. 1 (top) and Table

    1 indicate that there was three times more active GOx

    per unit area for PtD/GOx/PPDBSA compared with

    PtC/GOx/PPDBSA biosensors. This is consistent with

    retention of a dome of enzyme solution around the disk

    tip (see Fig. 2) as it is removed vertically from the liquid,as expected from surface tension considerations. Upon

    evaporation, the density of GOx on the disk surface

    was therefore higher than that achieved when the corre-

    sponding cylinder geometry (see Fig. 2) was prepared.

    (The eect of the drop that formed on the bottom of

    the PtC electrodes had a negligible contribution to the

    overall response of the PtC-based biosensors because

    the disk is 30 times smaller than the cylinder.)

    3.2. Glucose response of cylinder-based biosensors with

    dierent enzyme congurations

    Biosensors fabricated from PtC wires (Fig. 2) were

    used for an initial determination of the eects of enzyme

    distribution with respect to the polymer matrix. Starting

    by calibrating biosensors where the GOx was immobi-lized with GA on bare Pt (i.e., containing no polymer,

    PtC/GOxGA), and then re-calibrating following elec-

    tropolymerization, Fig. 1 (bottom) shows that the

    PPD matrix caused a signicant lowering the glucose

    sensitivity, and that re-loading GOx over the polymer

    restored the response.

    This nding was explored further using a larger pop-

    ulation of similar biosensors (Table 1). There was no sig-nicant dierence between the Jmax values for biosensors

    where the GOx was on the surface, either as PtC/GOx

    GA (36 7 lA cm2, n = 6) or PtC/PPDBSA/GOxGA (29 4 lA cm2, n = 10, p > 0.35). The KM(G)values were also indistinguishable (5.0 1.4 mM, n = 6

    vs. 6.4 0.3, n = 10, p > 0.23) and close to the value of

    5 mM cited for unhindered GOx [22]. In contrast, when

    polymer was formed after enzyme immobilization, as inPtC/GOx/PPDBSA, the Jmax was lower (6 1

    lA cm2, n = 21, p < 0.001 compared with no PPD)and the KM(G) was higher (15 2 mM, n = 21,

    p < 0.02). Thus, approximately 80% of the GOx was

    inactivated by the electropolymerization process (Fig. 1

    and Table 1), either by being denatured, blocked from

    access to solution glucose, or lost from the surface com-

    pletely. The increase in average KM(G) caused by thePPD suggests that the enzyme remaining active in the

    198 C.P. McMahon et al. / Journal of Electpolymer layer has a diminished anity (k1) for glucose(Eq. (5)) that can be represented schematically (see

    Fig. 2, top left). Thus, with a thickness that has been

    determined to be in the range 10 and 30 nm for the insu-

    lating form of the polymer generated at neutral pH

    [34,49,58,59], it is likely that PPD hinders a population

    of surface GOx molecules (9 nm diameter [22]) that be-come distributed within the polymer matrix. More de-

    tailed trends in KM(G) values are examined below.

    3.3. Glucose response of biosensors with dierent

    geometries and enzyme congurations

    The Jmax values for six combinations of geometry and

    enzyme distribution are compared in Table 1. A numberof trends are evident. The largest mean value of Jmax was

    achieved for disk biosensors containing no permselective

    polymer, PtD/GOxGA. When compared with the cor-

    responding cylinder design, PtC/GOxGA, a seven times

    lower Jmax-value was obtained, demonstrating again the

    eectiveness of the disk shape in dip-evaporation en-

    zyme deposition. Electropolymerization after GOx

    immobilization on the bare disk, as observed for the cyl-inders, led to a substantial reduction in glucose sensitiv-

    ity, that was counteracted when enzyme was deposited

    over the PPD-coated metal, i.e., for PtD/PPDBSA/

    GOx.

    The overall pattern, therefore, in the data presented

    in Fig. 1 and Table 1 is that more active enzyme can

    be deposited on disks, compared with cylinders, under

    the same conditions of immobilization. The GOx wasmost active when present on a surface (either bare metal

    or polymer) compared with biosensors in which the PPD

    was generated after the enzyme. These ndings are in

    line with expectation and consistent with the schematic

    models given in Fig. 2, but only reect the amount of ac-

    tive GOx in each design. Equally important, are the cor-

    responding KM(G) values, because this is a key

    parameter in determining both the biosensors sensitivityto glucose in the linear response region (Jmax/KM(G))and the extent of the linear range (1/2KM(G)). Wetherefore carried out a detailed correlation analysis of

    the dependence of KM(G) on enzyme loading, both with-

    in and between designs.

    3.4. Correlation analyses for KM(G) vs. Jmax

    Fig. 3 (top) shows the correlation between KM(G)

    and Jmax (enzyme loading) for the two PtC-based biosen-

    sor designs incorporating PPD. Unexpectedly, KM(G)

    values for PtC/GOx/PPDBSA showed a statistically sig-

    nicant negative correlation (R2 = 0.37, p < 0.01) with

    Jmax, ranging from 15 mM for electrodes with the low-est GOx loading to 5 mM for those with the highestloading. Thus, the change in KM(G) here cannot be as-cribed to changes in oxygen demand as this factor would

    lytical Chemistry 580 (2005) 193202have the opposite eect [52]. Instead, an explanation is

  • troanaPtC/PPD~GOx

    0 10 20 30 40 500

    5

    10

    15

    20

    slope = 0.026 0.023R2 = 0.14 (n=10)

    p > 0.28

    PPD/GOx:

    slope = -1.15 0.36R2 = 0.37 (n=19)

    p < 0.01

    GOx/PPD:

    Jmax / A cm-2

    KM

    (G) /

    mM

    Pt/PPD~GOx25 PtC/GOx/PPD

    PtC/PPD/GOx

    C.P. McMahon et al. / Journal of Elecprovided by the model of the GOx/PPD matrix shown in

    Fig. 2. For biosensors of this design with very low active

    enzyme levels, much of the GOx is buried in the PPD

    that hinders binding by glucose, decreasing k1 and there-fore increasing KM(G); see Eq. (5). As the Jmax increases,

    a greater fraction of the GOx lies towards the surface of

    the polymer so that the average enzymesubstrate bind-

    ing barrier diminishes, decreasing KM(G). Although this

    interpretation is speculative, it is strongly supported by

    the behavior of KM(G) vs. Jmax for the other cylinder

    PPD-based design: PtC/PPDBSA/GOx; see Fig. 3

    (top). Here, the GOx was placed over the polymer andthe observed values of KM(G) are all similar, i.e.,

    independent of Jmax over a four times larger range of

    Jmax than for the PtC/GOx/PPDBSA design. Moreover,

    the values of KM(G) for the GOx-on-surface congura-

    tions clustered around the minimum value observed

    for the PtC/GOx/PPDBSA design: 5.0 1.4 mM

    0 100 200 300 400 500 6000

    5

    10

    15

    20

    slope = 0.035 0.004R2 = 0.766 (n=30)

    p < 0.001

    PtD/PPD/GOx

    Jmax / A cm-2

    K M(G

    ) / m

    M

    Fig. 3. KM(G) as a function of enzyme loading and distribution.

    Correlation plots of the apparent MichaelisMenten parameters (Eq.

    (4)), KM(G) vs. Jmax for Pt/PPD electrodes incorporating the enzyme

    GOx in a variety of congurations: as 125-lm diameter cylinders (PtC)and disks (PtD); and with the GOx immobilized either before (GOx/

    PPD) or after (PPD/GOx) deposition of the PPD polymer. (Top)

    Cylinder biosensors showing the dierent correlation behaviors of the

    two enzyme congurations. (Bottom) Systematic increase in KM(G)

    with enzyme loading (Jmax) for PtD-based biosensors with surface GOx

    compared with the cylinder geometry behavior, as detailed above

    (top).(PtC/GOxGA, n = 6, R2 vs. Jmax = 0.16, p > 0.42); and

    6.4 0.3 mM (PtC/PPDBSA/GOxGA, n = 10, R2 vs.

    Jmax = 0.14, p > 0.28), values also consistent with litera-

    ture KM(G) for unhindered GOx molecules [22].

    Table 1 shows that there were some similarities be-

    tween the average KM(G) values for cylinder and disktype biosensors. Thus, there were no dierences be-

    tween: the high values of KM(G) observed for enzyme-

    hindered PtC/GOx/PPDBSA and PtD/GOx/PPDBSA

    designs (15 1 mM, p > 0.94; see also Fig. 1) that have

    been reported previously for GOx trapped in polymers

    [63,64]; the low values for the exposed enzyme in

    PtC/PPDBSA/GOxGA (6.4 0.3 mM) and PtD/

    PPDBSA/GOxGA (7.1 1.1 mM, p > 0.40); or thosefor PtC/GOxGA (5.0 1.4 mM) and PtD/GOxGA

    (9.0 1.4 mM, p > 0.20).

    Because the values for PtD/PPDBSA/GOxGA and

    PtD/GOxGA electrodes were on the high side of the

    averages, it is possible that the very high GOx loading

    for these biosensor designs, discussed above, might be

    aecting KM(G) as observed previously [52]. Therefore,

    correlation analyses (KM(G) vs. Jmax) were performedfor disk-based designs showing high average Jmax values

    (GOx exposed on the surface) to compare with cylinder

    results (Fig. 3). There was no dierence between the

    linear regression correlation slopes calculated for

    PtD/GOxGA (31 4 lM lA1 cm2, R2 = 0.963,

    p < 0.02) and PtD/PPDBSA/GOxGA (36 4 lMlA1 cm2, R2 = 0.772, p < 0.001) designs (p > 0.63).The correlation points for these two sensor types weretherefore pooled to compare with cylinder behavior.

    The positive correlation between KM(G) and Jmax for

    the pooled disk-based biosensor designs (35 4

    lM lA1 cm2, n = 30, R2 = 0.766, p < 0.001) was verydierent to that for the cylinder designs (negative or zero

    correlation); see Fig. 3 (bottom). It appears, therefore,

    that the extensive build-up of enzyme on the disk sur-

    face, possibly through a stacking arrangement previ-ously observed [65], led to the signicant positive

    correlation between KM(G) and Jmax for these PtD-based

    biosensors, associated with a reduction in GOx e-

    ciency. This loss of eciency could be due to either

    the development of a barrier to glucose binding by

    GOx itself as the enzyme loading increased, or to a lim-

    itation in oxygen supply with excess loading. This issue

    is addressed in Section 3.6.

    3.5. Linear region glucose sensitivity

    Having established patterns in enzyme loading (Jmax)and KM(G) for the six biosensor designs listed in Table 1,

    it is now possible to compare rationally the analytically

    more relevant parameter of LR glucose slope which, in

    the limit as glucose concentration approaches zero, is

    Jmax/KM(G). Based on the analyses described above, it

    lytical Chemistry 580 (2005) 193202 199was not surprising that the lowest LR slope, determined

  • PtD/PPD-BSA/GOx in 1 mM glucose

    0 20 40 60 800

    1

    2

    3

    4

    5

    6

    7

    R2 = 0.996

    KM(O2)

    J'max

    [O2] / M

    J glu

    c /

    A cm

    -2

    Fig. 4. Example of raw data for the response (current density) of a

    PtD/PPDBSA/GOx biosensor to 1 mM glucose as a function of

    continuously changing concentration of dissolved oxygen. The curve

    was generated using non-linear regression and Eq. (6) (R2 = 0.996).

    The KM(O2) values obtained for dierent concentrations of glucose are

    plotted in Fig. 5.

    Oxygen dependence of PtD/PPD-BSA/GOx

    0.0 0.5 1.0 1.5 2.00

    4

    8

    12

    16

    20

    24

    28

    [glucose] / mM

    K M(O

    2) /

    M O

    2

    Fig. 5. Values of KM(O2) for the PtD/PPDBSA/GOx design

    (means SD, n = 4; see Fig. 4) plotted as a function of glucose

    concentration. The improved experimental design allowed more

    reproducible results at low concentrations of glucose and a clear

    demonstration that KM(O2) is a linear function in glucose

    (R2 = 0.9995) with an intercept at the origin. The slopes of similar

    plots for dierent biosensor designs were used to compare the O2

    roanalytical Chemistry 580 (2005) 193202experimentally by calibration in the LR, was for the cyl-

    inder design with hindered enzyme (PtC/GOx/PPDBSA,

    0.4 0.1 lA cm2 mM1; Table 1), whereas the highestvalue was that for a disk-based design with surface en-

    zyme (PtD/GOxGA, 21 2 lA cm2 mM1), a value

    50 times more sensitive than the cylinder design usedfor neurochemical monitoring in vivo [23,38,48]. How-

    ever, the absence of the permselective PPD barrier in

    the PtD/GOxGA sensors makes them unsuitable for

    applications in real samples containing interference

    species, such as ascorbic and uric acids. The most sensi-

    tive biosensor involving the PPD layer was PtD/PPD

    BSA/GOx (11 2 lA cm2 mM1) that, with a 25times greater sensitivity (LR current density slope), anda 30 times smaller area, than the standard PtC/GOx/

    PPDBSA design, could be very useful for in vivo anal-

    ysis provided its oxygen dependence was suitable for

    these applications.

    The limit of detection (LOD) for the dierent biosen-

    sor congurations was determined as 3r of the corre-sponding background currents. The values ranged

    from 19 6 lM for the lowest LR sensitivity design(PtC/GOx/PPDBSA) to 3 1 lM for the highest sensi-tivity implantable biosensor (PtD/PPDBSA/GOx).

    Thus, all these congurations were suitable for applica-

    tions involving low glucose levels, for example, in brain

    ECF (as low as 350 lM [66]). Other GOx designs, suchas those reported recently involving platinization of Pt

    cylinders, can show excellent sensitivity, but much

    higher LOD values [25].With a correspondingly higher glucose turnover rate

    for the disk-based designs compared to PtC/GOx/

    PPDBSA sensors that have previously been shown

    not to be limited by oxygen over relevant physiological

    ranges for both glucose and pO2 [23], it was important

    to evaluate and compare the oxygen dependence of the

    four variations of PPD-containing biosensors. Such a

    comparison should also reveal whether oxygen supplylimitations are responsible for the positive correlation

    of KM(G) with GOx loading observed in Fig. 3 for

    disk-based devices.

    3.6. Oxygen dependence

    The experimental protocol for determination of oxy-

    gen dependence was similar to that reported recently[23]. Fig. 4 shows the variation of Jgluc, recorded with

    a PtD/PPDBSA/GOx biosensor in 1 mM glucose, as a

    function of solution oxygen concentration that was

    monitored simultaneously using the self-calibrating

    CellOx probe. The data tted the adapted Michaelis

    Menten equation well (Eq. (6), R2 = 0.996) and provided

    KM(O2) values that can be used as an index of oxygen

    dependence, with lower values of KM(O2) beingpreferred. In the previous study that characterized

    200 C.P. McMahon et al. / Journal of ElectPtC/GOx/PPDBSA electrodes [23], the lowest glucoseconcentration for KM(O2) studies was 0.5 mM, and a

    linear relationship for KM(O2) vs. glucose concentration

    suggested based on an R2 value of 0.96. Because baseline

    glucose levels in certain brain regions are below 0.5 mM

    [62,66,67], this linearity was re-investigated down to a

    concentration of 0.2 mM. Fig. 5 conrms this importantlinearity through the origin (R2 = 0.9995), and allows

    one to interpolate a value of KM(O2) for any

    concentration of glucose for the corresponding biosen-

    sor design.

    Using this protocol and analysis (Fig. 5) for a range

    of discrete concentrations of glucose, the slope of

    KM(O2) vs. glucose concentration was determined forsensitivities of the diverse devices.

  • each of the four PPD-containing biosensor types. This

    slope (lM O2 per mM glucose) was remarkably similarfor all four designs: PtC/GOx/PPDBSA (12.2 1.3,

    n = 12); PtC/PPDBSA/GOx (15.7 1.0, n = 12);

    PtD/GOx/PPDBSA (13.3 0.6, n = 4); and PtD/PPD

    BSA/GOx (12.1 0.6, n = 4). These results are surpris-ing given the very dierent LR glucose slopes for the

    four designs (Table 1) and the corresponding greater

    oxygen demand of the disk-based devices, even in the

    LR of glucose response.

    The trends in these average KM(O2) values can, how-

    ever, be understood better in terms of the correlation be-

    tween KM(O2) and KM(G) for individual sensors. Fig. 6

    (top) shows that, not only did the KM(O2) not increaseas the LR glucose slope increased, there was a statisti-

    cally signicant, albeit modest, decrease in KM(O2) for

    sensors with higher glucose sensitivity. A possible cause

    for this behavior is revealed in Fig. 6 (bottom). There

    was a parallel weak negative correlation between

    KM(O2) and KM(G). Thus, the sensors that had high

    GOx loading had high LR slope (Table 1), but also a

    higher KM(G). There was more enzyme in the biosensinglayer, but each GOx molecule had a lower anity for

    glucose and a lower turnover rate that lowered oxygen

    can therefore be understood, not as a limitation in oxy-gen supply, but as a decrease of the anity of glucose

    oxygen interference [23], it remains to be seen whether

    the more active designs with surface enzyme can with-

    16

    18

    20 slope = -0.240.09R2 = 0.213 (n=30)p < 0.02

    M-1

    gluc

    ose

    C.P. McMahon et al. / Journal of Electroana0 3 6 9 12 1510

    12

    14

    glucose LR slope / A cm-2 mM -1

    4 6 8 10 12 14 16 1810

    12

    14

    16

    18

    20 slope = -0.270.14R2 = 0.20 (n=18)p = 0.06

    KM(G)/ mM glucose

    KM

    (O2)

    /M

    O2

    mM

    -1

    gluc

    ose

    KM

    (O2)

    /M

    O2

    m

    Fig. 6. Correlation analyses for oxygen dependence (KM(O2) slope; see

    Fig. 5) vs. glucose sensitivity quantied as both the linear region

    glucose slope (top) and the KM(G), the apparent Michaelis constant

    for glucose (bottom). There was no evidence for increased oxygen

    dependence as the glucose sensitivity increased; moreover, there was a

    slight statistically signicant trend for the oxygen dependence todecrease with increased glucose sensitivity.stand the conditions of direct tissue contact in in vivo

    monitoring, or whether a protective barrier such as a

    microdialysis membrane will be needed for protection.

    In either case, the PtD/PPDBSA/GOx biosensor will

    now enable us to explore glucose concentration

    dynamics of smaller and layered brain regions withgood sensitivity and minimal interference from uctu-

    ations in tissue pO2.

    Acknowledgements

    This work was funded in part by Science Foundation

    Ireland (04/BR/C0198). We thank Enterprise Ireland fora postgraduate scholarship (CMcM), and UCD forfor GOx (Eq. (5)) associated with steric hindrance by

    neighboring enzyme molecules. A similar, but more

    pronounced, trend has been reported recently for gluta-

    mate oxidase in PPD-based biosensors, where glutamate

    binding was even more sensitive to enzyme loading, an

    eect due, at least in part, to electrostatic interactions

    between the anionic substrate and the polyanionicenzyme [68].

    4. Conclusions

    The glucose sensitivity and oxygen dependence of a

    variety of implantable GOx-based biosensors incorpo-

    rating a PPDBSA polymer permselective barrier weremeasured and compared. Dip-evaporation of the en-

    zyme was more ecient on disks than on cylinders. Be-

    cause polymerization after GOx deposition deactivated

    a large fraction of the immobilized enzyme, the most

    sensitive biosensor design studied was PtD/PPDBSA/

    GOx in the linear region of glucose response, that was

    20 times more sensitive than the more widely usedPtC/GOx/PPDBSA conguration. The oxygen depen-dence, KM(O2), of both active and less active designs

    was surprisingly similar, a nding that could be rational-

    ized in terms of an increase in KM(G) with increased en-

    zyme loading.

    Although a basic PPD-GOx polymerenzyme com-

    posite biosensor design has been used for continuous

    recording over weeks in the harsh electrochemical

    environment of brain tissue [48] without signicantdemand on a molecular level. For sensors with high glu-

    cose sensitivity, the enzyme layer as a whole had a high

    oxygen demand, but this was spread out, possibly over a

    3D stacking arrangement of the GOx [65]. The positive

    correlation for KM(G) vs. Jmax shown in Fig. 3 (bottom)

    lytical Chemistry 580 (2005) 193202 201nancial support.

  • References

    [1] J. Wang, Electroanalysis 13 (2001) 983.

    [2] I. Karube, Y. Nomura, J. Mol. Catal. B 10 (2000) 177.

    [3] C.M. Braguglia, Chem. Biochem. Eng. Quart. 12 (1998) 183.

    [4] S. Cosnier, Anal. Bioanal. Chem. 377 (2003) 507.

    [5] S.J. Dong, B.Q. Wang, Electroanalysis 14 (2002) 7.

    [6] L.D. Mello, L.T. Kubota, Food Chem. 77 (2002) 237.

    [7] J. Wang, Anal. Chem. 65 (1993) R450.

    [8] F. Scheller, D. Kirstein, F. Schubert, D. Pfeier, C. McNeil,

    Russian J. Electrochem. 29 (1993) 1334.

    [9] E. Magner, Analyst 123 (1998) 1967.

    [10] G.S. Wilson, Y.B. Hu, Chem. Rev. 100 (2000) 2693.

    [36] T. Yao, T. Yano, H. Nishino, Anal. Chim. Acta 510 (2004) 53.

    [37] R.D. ONeill, S.C. Chang, J.P. Lowry, C.J. McNeil, BiosensorsBioelectron. 19 (2004) 1521.

    [38] R.D. ONeill, J.P. Lowry, in: R. Meyers (Ed.), Encyclopedia ofAnalytical Chemistry, Wiley, Chichester, 2000.

    [39] P.N. Bartlett, R.G. Whitaker, J. Electroanal. Chem. 224 (1987)

    37.

    [40] N.F. Almeida, J. Wingard, M.K. Malmros, Ann. NY Acad. Sci.

    613 (1990) 448.

    [41] J.P. Lowry, R.D. ONeill, Anal. Chem. 64 (1992) 453.[42] E. Dempsey, J. Wang, M.R. Smyth, Talanta 40 (1993) 445.

    [43] E. Ekinci, A.A. Karagozler, A.E. Karagozler, Electroanalysis 8

    202 C.P. McMahon et al. / Journal of Electroanalytical Chemistry 580 (2005) 193202[11] S. Cosnier, Biosensors Bioelectron. 14 (1999) 443.

    [12] S. Cosnier, C. Innocent, L. Allien, S. Poitry, M. Tsacopoulos,

    Anal. Chem. 69 (1997) 968.

    [13] M.R. Ryan, J.P. Lowry, R.D. ONeill, Analyst 122 (1997) 1419.[14] N.V. Kulagina, L. Shankar, A.C. Michael, Anal. Chem. 71 (1999)

    5093.

    [15] E. Mikeladze, A. Schulte, M. Mosbach, A. Blochl, E. Csoregi, R.

    Solomonia, W. Schumann, Electroanalysis 14 (2002) 393.

    [16] J.K. Leypoldt, D.A. Gough, Anal. Chem. 56 (1984) 2896.

    [17] J.P. Lowry, R.D. ONeill, Electroanalysis 6 (1994) 369.[18] C.P. McMahon, S.J. Killoran, S.M. Kirwan, R.D. ONeill, Chem.

    Commun. (2004) 2128.

    [19] J.J. Kulys, Enzyme Microb. Technol. 3 (1981) 344.

    [20] W.J. Albery, P.N. Bartlett, J. Electroanal. Chem. 194 (1985) 211.

    [21] D. Compagnone, G. Federici, J.V. Bannister, Electroanalysis 7

    (1996) 1151.

    [22] S.V. Sasso, R.J. Pierce, R. Walla, A.M. Yacynych, Anal. Chem.

    62 (1990) 1111.

    [23] B.M. Dixon, J.P. Lowry, R.D. ONeill, J. Neurosci. Methods 119(2002) 135.

    [24] T. Ikeda, I. Katasho, M. Kamei, M. Senda, Agric. Biol. Chem. 48

    (1984) 1969.

    [25] J.I.R. De Corcuera, R.P. Cavalieri, J.R. Powers, J. Electroanal.

    Chem. 575 (2005) 229.

    [26] H. Olivia, B.V. Sarada, K. Honda, A. Fujishima, Electrochim.

    Acta 49 (2004) 2069.

    [27] M. Gao, L.M. Dai, G.G. Wallace, Electroanalysis 15 (2003) 1089.

    [28] N.V. Kulagina, A.C. Michael, Anal. Chem. 75 (2003) 4875.

    [29] C. Mousty, B. Galland, S. Cosnier, Electroanalysis 13 (2001) 186.

    [30] J. Wang, L. Chen, M.P. Chatrathi, Anal. Chim. Acta 411 (2000)

    187.

    [31] A.A. Karyakin, E.A. Kotelnikova, L.V. Lukachova, E.E. Kar-yakina, J. Wang, Anal. Chem. 74 (2002) 1597.

    [32] P.N. Bartlett, P.R. Birkin, Anal. Chem. 66 (1994) 1552.

    [33] J.P. Lowry, K. McAteer, S.S. El Atrash, A. Du, R.D. ONeill,Anal. Chem. 66 (1994) 1754.

    [34] C. Malitesta, F. Palmisano, L. Torsi, P.G. Zambonin, Anal.

    Chem. 62 (1990) 2735.

    [35] E.R. Reynolds, A.M. Yacynych, Electroanalysis 5 (1993) 405.(1996) 571.

    [44] L.I. Netchiporouk, N.F. Shram, N. Jarezic-Renault, C. Martelet,

    R. Cespuglio, Anal. Chem. 68 (1996) 4358.

    [45] P.N. Bartlett, J.H. Wang, W. James, Analyst 123 (1998) 387.

    [46] J.M. Cooper, P.L. Foreman, A. Glidle, T.W. Ling, D.J. Pritchard,

    J. Electroanal. Chem. 388 (1995) 143.

    [47] J.P. Lowry, M. Miele, R.D. ONeill, M.G. Boutelle, M. Fillenz, J.Neurosci. Methods 79 (1998) 65.

    [48] R.D. ONeill, J.P. Lowry, M. Mas, Crit. Rev. Neurobiol. 12(1998) 69.

    [49] J.D. Craig, R.D. ONeill, Analyst 128 (2003) 905.[50] K. McAteer, R.D. ONeill, Analyst 121 (1996) 773.[51] C. Bourdillon, V. Thomas, D. Thomas, Enzyme Microb. Technol.

    4 (1982) 175.

    [52] Y.N. Zhang, G.S. Wilson, Anal. Chim. Acta 281 (1993) 513.

    [53] T. Ikeda, K. Miki, M. Senda, Anal. Sci. 4 (1988) 133.

    [54] P.N. Bartlett, K.F.E. Pratt, Biosensors Bioelectron. 8 (1993) 451.

    [55] C. Phanthong, M. Somasundrum, J. Electroanal. Chem. 558

    (2003) 1.

    [56] R. Baronas, F. Ivanauskas, F. Ivanauskas, J. Kulys, J. Math.

    Chem. 35 (2004) 199.

    [57] R. Baronas, F. Ivanauskas, J. Kulys, Sensors 3 (2003) 248.

    [58] T.W. Sohn, P.W. Stoecker, W. Carp, A.M. Yacynych, Electro-

    analysis 3 (1991) 763.

    [59] S. Myler, S. Eaton, S.P.J. Higson, Anal. Chim. Acta 357 (1997)

    55.

    [60] L.J. Murphy, Anal. Chem. 70 (1998) 2928.

    [61] J.M. Cooper, D.J. Pritchard, J. Mater. Sci. 5 (1994) 111.

    [62] J.P. Lowry, M. Fillenz, Bioelectrochemistry 54 (2001) 39.

    [63] G. Fortier, M. Vaillancourt, D. Belanger, Electroanalysis 4 (1992)

    275.

    [64] T. Kaku, H.I. Karan, Y. Okamoto, Anal. Chem. 66 (1994)

    1231.

    [65] G.E. De Benedetto, C. Malitesta, C.G. Zambonin, J. Chem. Soc.,

    Faraday Trans. 90 (1994) 1495.

    [66] J.P. Lowry, R.D. ONeill, M.G. Boutelle, M. Fillenz, J. Neuro-chem. 70 (1998) 391.

    [67] A.E. Fray, M. Boutelle, M. Fillenz, J. Physiol. (London) 504

    (1997) 721.

    [68] C.P. McMahon, R.D. ONeill, Anal. Chem. 77 (2005) 1196.

    Design variations of a polymer -- enzyme composite biosensor for glucose: Enhanced analyte sensitivity without increased oxygen dependenceIntroductionMaterials and methodsChemicals and solutionsInstrumentation and softwareWorking electrode preparationAmperometric experimentsData analysis

    Results and discussionGlucose response of a cylinder- and disk-based polymer ndash enzyme configurationGlucose response of cylinder-based biosensors with different enzyme configurationsGlucose response of biosensors with different geometries and enzyme configurationsCorrelation analyses for KM(G) vs. JmaxLinear region glucose sensitivityOxygen dependence

    ConclusionsAcknowledgementsReferences

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