Role of reaction kinetics and mass transport in glucose sensing with nanopillar array electrodes
© Anandan et al; licensee BioMed Central Ltd. 2007
Received: 12 September 2007
Accepted: 10 October 2007
Published: 10 October 2007
The use of nanopillar array electrodes (NAEs) for biosensor applications was explored using a combined experimental and simulation approach to characterize the role of reaction kinetics and mass transport in glucose detection with NAEs. Thin gold electrodes with arrays of vertically standing gold nanopillars were fabricated and their amperometric current responses were measured under bare and functionalized conditions. Results show that the sensing performances of both the bare and functionalized NAEs were affected not only by the presence and variation of the nanoscale structures on the electrodes but also by the reaction kinetics and mass transport of the analyte species involved. These results will shed new light for enhancing the performance of nanostructure based biosensors.
Biosensors are important devices for monitoring biological species in various processes of environmental, fermentation, food and medical concerns. The main challenges biosensors face include low sensitivity, poor specificity and proneness to fouling. The advent of nanotechnology presents some promising solutions for alleviating these problems. While many efforts have been devoted to improving the performances of biosensors by taking advantage of nanostructures or macrostructures with nanoscale features, the effort to elucidate the underlying mechanism governing the performances of biosensors enhanced with nanostructures is still scant.
This study intends to investigate the role of reaction kinetics and mass transport in biosensing when electrodes with nanoscale features are used. For this purpose, we used glucose biosensor as a model system. In a typical glucose biosensor, an enzyme such as glucose oxidase is immobilized onto the electrode surface [1, 2]. The performance of such functionalized electrodes can be improved by either adjusting the spatial distribution of the enzyme or by modifying the morphology of the electrode surface. To achieve a high efficiency in immobilizing an enzyme onto the electrode surface, various techniques have been developed, such as the use of self-assembled monolayer [1–4], conducting polymers [5, 6] and sol-gels . Among these methods, the self-assembled monolayer (SAM) approach offers a better control for enzyme distribution at the molecular level, a high degree of reproducibility in enzyme immobilization and a short distance between the immobilized enzyme and the electrode surface [1, 4]. The SAM approach, however, is limited by the amount of the enzyme that can be immobilized onto the electrode surface, which in turn will affect the sensing performance of the biosensor . To increase the amount of immobilized enzyme various nanostructures such as nanotubes, nanoparticles and nanorods have been explored in order to increase the active surface area of the electrodes. For example, nanostructures like gold nanotubes , carbon nanotubes [5, 9] and gold nanoparticles  have been incorporated into electrode surfaces and they exhibited better performance than conventional flat electrodes.
Recently Wang et al.  used nanostructured platinum electrodes functionalized with glucose oxidase for glucose detection. These electrodes showed a significant (two orders of magnitude) increase in glucose detection sensitivity as compared with a flat electrode, but the response of these electrodes to K4Fe(CN)6 was just 2.3 times that of the flat electrode. They attributed such sensitivity enhancements for glucose detection to the increased enzyme loading and improved retention of hydrogen peroxide near the electrode surface without examining systematically the role of reaction kinetics and mass transport. We believe that the electrical current response of these nanostructured electrodes is controlled by reaction kinetics, mass transport and the geometric topography of the nanostructures. Thus, to be able to understand the mechanism governing such an electrochemical process for the purpose of improving the performance of nanostructure based electrodes, it is necessary to investigate the role of reaction kinetics and mass transport in biosensing when nanostructured electrodes are used.
Nanopillar array electrodes (NAEs) with three different pillar heights were fabricated using a template method . In fabricating these electrodes, a layer of gold film about 150 nm thick was first sputter-coated onto one side of a porous anodic alumina (PAA) circular disc (d = 25 mm; Whatman Inc, Maidstone, England) having an average pore diameter of 150 nm using a SPI sputter coater (Structure probe Inc, West Chester, PA). Then, a thicker gold layer was electrodeposited on top of the sputtered gold film to form a strong supporting base in an Orotemp24 gold plating solution (Technic Inc, Cranston, Rhode Island) with a current density of 5 mA/cm2 for two minutes. This supporting base was masked with Miccrostop solution (Pyramid plastics Inc., Hope, Arkansas) for insulation. After that, gold nanopillars were electrodeposited through the open pores of the PAA disc from the uncoated side under an electrical current density of 5 mA/cm2 at 65°C. The deposition time was varied for achieving nanopillars of different heights. For this study, specimens with three different nanopillar heights were prepared with the electrodeposition time controlled at 1, 7 and 15 minutes. After nanopillar deposition, the PAA disc was dissolved in 2.0 M NaOH resulting in a thin gold sheet with arrays of vertically standing gold nanopillars. The fabricated specimens were cut into small square pieces (about 3.2 × 3.2 mm2) and they were grouped into specimens A, B and C by their nanopillar height. For connecting the electrodes, a copper tape were attached to the backside of an electrode with the exposed part of the copper tape insulated using Miccrostop. Of these small specimens, some were used for scanning electron microscopy (SEM) imaging analyses, and some for electrochemical experiments (bare and functionalized conditions). For controls, two flat gold electrodes (one for bare and one for functionalized condition) with the same geometric size were prepared by depositing a thin film (300 nm) of gold on a titanium coated glass plate using a thermal evaporator (built in house). Prior to the electrochemical experiments, all electrodes (NAEs and flat) were cleaned by running cyclic voltammetry (CV) in 0.3 M H2SO4 between -500 mV and 1500 mV until a stable CV curve was obtained for each specimen, and then washed with deionized water.
These electrodes were characterized in either bare or functionalized conditions. In the bare condition the cleaned electrodes were used directly, and in the functionalized condition the cleaned electrodes were further functionalized prior to use. To functionalize the electrodes, their surfaces were first modified with a SAM layer by placing them in a 75% ethanol solution containing 10 mM 3-mercaptopropionic acid. Then the SAM modified electrodes were rinsed in 75% ethanol and immersed in a 0.1 M 2-(N-morpholino) ethanesulfonic buffer solution (pH of 3.5) containing 2 mM 1-ethyl-3-(3-dimethylaminopropyl) carbodiimide hydrochloride and 5 mM N-hydroxysuccinimide for activation for two hours. After washing in phosphate buffer solution (PBS), the activated NAEs were placed in PBS solution at pH 7.4 containing 1 mg/ml of glucose oxidase for two hours under constant stirring. The reason for setting the immobilization time to two hours is that according to literature , enzyme loading reaches its maximum in about 2 hours and it saturates afterwards. From the electrochemical experiments, the amperometric current responses of both bare and functionalized NAEs along with flat controls were measured using a conventional three-electrode cell with an Ag/AgCl reference electrode and a platinum counter electrode with the Multistat 1480 (Solartron Analytical, Houston TX, USA) electrochemical system.
For the bare-electrodes, their amperometric current responses at different concentrations of K4Fe(CN)6 to each incremental addition of 80 μl of 1 M K4Fe(CN)6 to a 20 ml solution containing 0.5 M Na2SO4 (equivalent to a 4 mM increase in K4Fe(CN)6concentration) were measured at a constant electrode potential of 350 mV (vs. Ag/AgCl), and the change in the current response upon the change in K4Fe(CN)6 concentration for both the NEAs and flat electrode was determined. For the functionalized NAEs, the amperometric current responses to each incremental addition of 50 μl of 1 M glucose to a 20 ml PBS solution (equivalent to a 2.5 mM increase in glucose concentration) containing 3 mM p-benzoquinone as a mediator were measured at a constant potential of 350 mV (vs. Ag/AgCl). In all experiments, the background current of all electrodes was allowed to stabilize before drops of target species were added. Prior to these experiments the electrolyte solution was deaerated with nitrogen and during experiments the solution was blanketed with nitrogen and stirred constantly at 600 rpm.
For glucose detection, the electrode reactions in the present study can be described by the following cascading events . With the catalysis of glucose oxidase (GOX-FAD), glucose was first oxidized into gluconolactone with GOX-FADH2 as a by-product. The GOX-FADH2 was then converted back to its oxidized form (GOX-FAD) by the p-benzoquinone mediator in the solution. The mediator itself was converted back to its original form by oxidation at the electrode surface, through which free electrons were generated and picked up by the electrode to produce a current response. These cascading events can be expressed by the following reactions:
Glucose + GOX-FAD → Gluconolac tone + GOX-FADH2 (1)
GOX - FADH2 + 2 Mediatorox → GOX - FAD + 2 Mediatorred + 2H+ (2)
2Mediatorred → 2Mediatorox + 2e- (3)
Where I s is the steady-state current measured at each glucose concentration, Imax the maximum current attainable, K m the apparent Michaelis-Menten constant, and S the concentration of the target species (i.e., glucose in this case). The parameter K m describes the enzymatic activity of glucose: the smaller the K m value is, the more efficient the enzymatic reaction is. When K m of an enzymatic biosensor is larger than the K m value of the freely dissolved enzyme, it usually implies that the enzyme immobilized in the biosensor is less efficient in oxidizing glucose than the dissolved enzyme . In this study, the K m values for the functionalized electrodes were determined by performing nonlinear curve fit using Eq.4 to the measured current-concentration data.
For the electrode reaction at the functionalized NAEs, we assumed that glucose was consumed at a flux of J g at the electrode surface to produce the mediator in its reduced-form at a flux of J M . Here J g and J M can be described by the following equations:
J glucose = kc G (5)
J M = kc G - k0c M exp(-αF(E - E std )/RT) (6)
where k represents the rate constant for Eq. 1, c G the concentration of glucose, c M the concentration of mediator, k0 the standard rate constant, α the charge transfer coefficient, F the Faraday constant, E the electrode potential, and E std the standard potential of the mediator. To simulate the actual event, the electrode was held at a constant overpotential of 350 mV. Under this condition, the reduced-form mediator was oxidized at the electrode surface to generate a current flux of J c :
J C = -2k0c M exp(-αF(E-E std )/RT) (7)
With these considerations, the amperometric current response of the electrodes in response to a drop of glucose was determined while the electrolyte solution was constantly stirred by a swirling vortex force applied at the center of the cell.
For the electrode reaction at the bare-electrode, we considered the redox of K4Fe(CN)6 with the reduction flux of K4Fe(CN)6 governed by:
J F = -k0FcF 1exp(-αF(E - E std ')/RT) + k0FcF 2exp(-αF(E - E std ')/RT) (8)
where k0Fis the electron transfer rate for both ferrocyanide and ferricyanide (assumed to be the same), cF 1the concentration of ferrocyanide, cF 2the concentration of ferricyanide, E the electrode potential, and E std ' the standard potential of ferro- and ferri-cyanide.
Besides the reaction kinetics discussed above, the mass transport in these electrochemical processes was mainly governed by diffusion and convection for the mobile species such as glucose and K4Fe(CN)6. The electromigration was ignored because of the presence of the supporting electrolyte in a high concentration.
Material constants and kinetic parameters used in the simulation 
Standard rate constant
1.5 × 10-3 (m/s)
Charge transfer coefficient
200 × 10-9 (m)
5 × 10-6 (m)
Surface reaction rate constant
5 × 10-4, 5 × 10-5, 5 × 10-7 (m/s)
Molecular weight of water
8.31 (J/K. mol)
9.648 × 104 (C/mol)
Diffusivity of ferro- and ferri-cyanide
where A represents the solute (e.g., glucose or the mediator) and B the solvent (e.g., water), ε B the association factor of the solvent, M B the molecular weight of the solvent, μ the viscosity of solution, V A the molar volume of solute glucose, and T the absolute temperature.
Results and Discussion
The roughness ratio, detection sensitivity, Imax and K m obtained from experiments
Sensitivity of bare electrodes to K4Fe(CN)6 (μA·mM-1·cm-2)
Sensitivity of functionalize electrodes to glucose (μA·mM-1·cm-2)
Imax glucose (μA)
K m glucose (mM)
One surprising observation, however, was that the sensitivity of these bare NAEs did not increase with the increase of the roughness ratio. This implies that the benefit of the increased surface area due to nanopillars has not been fully realized. It seems that only the top part of the nanopillars has contributed to the increase of active electrode surface for electron transfer, which may explain why there is only a two-fold increase in the current responses of all the NAEs as compared with the flat electrode. We speculate that the electroactive species K4Fe(CN)6 may encounter certain difficulties in its transport to the small spaces between the bare nanopillars as the result of either a low diffusivity or a fast electron transfer rate constant. With a low diffusivity, it would be difficult for K4Fe(CN)6 to diffuse deep into the small spaces between the nanopillars, while with a fast electron transfer rate constant, most of the species K4Fe(CN)6 would get oxidized near the top ends of the nanopillars before it diffuses down the gaps. Under such a circumstance, it is conceivable that only the top regions of the nanopillars are serving their active duty in transferring electrons.
Figure 4B shows the variations of the steady-state amperometric current with glucose concentration (from 2.5 mM to 15 mM) along with the corresponding linear regression lines. By taking the slope of the regression lines and normalizing it with respect to the geometric area of the electrode in each case, we obtained the sensitivity measurement for the functionalized electrodes (NAEs and flat). From the obtained sensitivity values listed in Table 2, we observed that unlike in the bare electrode cases, the sensitivity of NAEs increases as the roughness ratio increases. The highest sensitivity value (Nano C) is about 3.13 μA·mM-1·cm-2 (about 12 times higher than that for a flat electrode) which is significantly higher than the value reported for a gold nanotube electrode (0.4 μA·mM-1·cm-2) . So for the functionalized NAEs, increasing the surface roughness of the NAEs does contribute to an increase in detection sensitivity.
In comparing the bare with the functionalized electrodes, we found that the highest nanostructure-induced sensitivity increase for the functionalized electrodes (12 times) is higher than that for the bare electrodes (2 times). This could be due to the difference in electrochemical species involved (i.e., glucose versus K4Fe(CN)6). These two electroactive species, however, have a similar diffusivity value (8 × 10-10m2/s for K4Fe(CN)6 and 7.6 × 10-10m2/s for glucose) . This fact suggests that the difference in the reaction rate constant at the bare and functionalized electrodes may play a more dominate role in affecting the current response. It is also possible that such an increase in the sensitivity of functionalized NAEs is the result of heightened retention of the mediator during glucose detection .
Steady-state amperometric current density obtained at different rate constants from computer simulation
Reaction rate constant (m/s)
Current Density (mA·cm-2)
5 × 10-4 for K4Fe(CN)6
5 × 10-5 for glucose
5 × 10-7 for glucose
Here and are the thiele moduli calculated based on the transverse and longitudinal diffusion times, and they are defined by = 2kr p /D G , = 2kL2/r p D G , where r p is the pore radius, L is the pore length, D G is the glucose diffusivity and k is the surface rate constant. For a surface rate constant of 5 × 10-4 m/s, we calculated = 0.131 and = 328.8. Under this condition, η is found to be η ≈ 1/ϕ = 0.055. This low effectiveness factor will hinder the transport of the target species to the spaces between the nanopillars. For a lower surface reaction rate constant of 5 × 10-7 m/s, we calculated = 131.5 × 10-6 and = 0.328, and under this condition η will be close to unity (η ≈ 1). This high effectiveness factor will surely enable more efficient transport of glucose to the functionalized surfaces in between the nanopillars.
The above results clearly indicate that the enhanced current response in glucose sensing with functionalized NAEs can be attributed to the effective mass transport facilitated by the relatively lower reaction rate constant of glucose. This fact suggests that to reap the true benefit of using nanostructured electrodes for enhancing the performance of biosensors, it is necessary to optimize the geometry of the nanopillars (their diameter, spacing and height) in order to accommodate the specific analyte species in terms of its reaction kinetics and mass transport.
The role of reaction kinetics and mass transport in biosensing using electrodes integrated with nanopillars of different heights was investigated. In an electrochemical based detection, the increased active surface area due to the addition of nanopillars may lead to enhanced sensing performances only when the reaction rate constant of the target species is low. At a higher reaction rate constant, only the top part of the nanopillar modified electrodes will serve the purpose for transferring electrons. To reap the benefit of using nanostructured electrodes for improving the sensing performances, it is necessary to optimize the geometry of the nanopillars to accommodate the specific analyte species in terms of its reaction kinetics and mass transport.
This work was partially supported by the National Science Foundation (ECS-0304340), the Faculty of Engineering and the College of Agricultural and Environmental Science at the University of Georgia.
- Berchmans S, Sathyajith R, Yegnaraman V: Layer-by-layer assembly of 1,4-diaminoanthraquinone and glucose oxidase. Materials Chemistry and Physics 2003, 77: 390-396. 10.1016/S0254-0584(02)00016-0View ArticleGoogle Scholar
- Gooding JJ, Erokhin P, Hibbert DB: Parameters important in tuning the response of monolayer enzyme electrodes fabricated using self-assembled monolayers of alkanethiols. Biosensors & Bioelectronics 2000, 15: 229-239. 10.1016/S0956-5663(00)00080-4View ArticleGoogle Scholar
- Gooding JJ, Praig VG, Hall EAH: Platinum-catalyzed enzyme electrodes immobilized on gold using self-assembled layers. Analytical Chemistry 1998, 70: 2396-2402. 10.1021/ac971035tView ArticleGoogle Scholar
- Losic D, Gooding JJ, Shapter JG, Hibbert DB, Short K: The influence of the underlying gold substrate on glucose oxidase electrodes fabricated using self-assembled monolayers. Electroanalysis 2001, 13: 1385-1393. 10.1002/1521-4109(200111)13:17<1385::AID-ELAN1385>3.0.CO;2-LView ArticleGoogle Scholar
- Gao M, Gordon LD: Biosensors Based on Aligned Carbon Nanotubes Coated with Inherently Conducting Polymers. Electroanalysis 2001, 15: 1089-1094. 10.1002/elan.200390131View ArticleGoogle Scholar
- Uang YM, Chow TC: Criteria for Designing a Polypyrrole Glucose Biosensor by Galvanostatic Electropolymerization. Electroanalysis 2002, 14: 1564-1570. 10.1002/1521-4109(200211)14:22<1564::AID-ELAN1564>3.0.CO;2-HView ArticleGoogle Scholar
- Qiaocui S, Tuzhi P, Yunu Z, Yang CF: An Electrochemical Biosensor with Cholesterol Oxidase/Sol-Gel Film on a Nanoplatinum/Carbon Nanotube Electrode. Electroanalysis 2005, 17: 857-861. 10.1002/elan.200403162View ArticleGoogle Scholar
- Delvaux M, Demoustier-Champagne S: Immobilisation of glucose oxidase within metallic nanotubes arrays for application to enzyme biosensors. Biosensors & Bioelectronics 2003, 18: 943-951. 10.1016/S0956-5663(02)00209-9View ArticleGoogle Scholar
- Wang J, Musameh M: Carbon Nanotube/Teflon Composite Electrochemical Sensors and Biosensors. Anal Chem 2003, 75: 2075-2079. 10.1021/ac030007+View ArticleGoogle Scholar
- Bharathi S, Nogami M: A glucose biosensor based on electrodeposited biocomposites of gold nanoparticles and glucose oxidase enzyme. Analyst 2001, 126: 1919-1922. 10.1039/b105318nView ArticleGoogle Scholar
- Wang J, Nosang M, Minhee Y, Harold M: Glucose oxidase entrapped in polypyrrole on high-surface-area Pt electrodes: a model platform for sensitive electroenzymatic biosensors. Journal of Electroanalytical Chemistry 2005, 575: 139-146. 10.1016/j.jelechem.2004.08.023View ArticleGoogle Scholar
- Anandan V, Rao YL, Zhang G: Nanopillar array structures for high performance electrochemical sensing. International journal of nanomedicine 2006, 1: 73-79. 10.2147/nano.2006.1.1.73View ArticleGoogle Scholar
- Gooding JJ, Erokhin P, Losic D, Yang W, Policarpio V, Liu J, Ho F, Situmorang M, Hibbert DB, Shapter JG: Parameters important in fabricating enzyme electrodes using self-assembled monolayers of alkanethiols. Anal Sci 2001,17(1):3-9. 10.2116/analsci.17.3View ArticleGoogle Scholar
- Harper AC: Modified electrodes for amperometric determination of glucose and glutamate using mediated electron transport. In PhD Thesis. Virginia Polytechnic Institute and State University, Chemistry Department; 2005.Google Scholar
- Winkler K: The kinetics of electron transfer in Fe(CN) 6 4-/3- redox system on platinum standard-size and ultramicroelectrodes. Journal of Electroanalytical Chemistry 1995, 388: 151-159. 10.1016/0022-0728(94)03847-VView ArticleGoogle Scholar
- Heinsohn RJ, Cimbala JM: Indoor Air Quality Engineering: Environmental Health and Control of Indoor Pollutants. New York: Marcel Dekker Inc; 2003.View ArticleGoogle Scholar
- Calvo EJ, Wolosiuk A: Supramolecular Architectures of Electrostatic Self-Assembled Glucose Oxidase Enzyme Electrodes. ChemPhysChem 2004, 5: 235-239. 10.1002/cphc.200300930View ArticleGoogle Scholar
- Balakotaiah V, Gupta N: Controlling regimes for surface reactions in catalyst pores. Chemical Engineering Science 2000, 55: 3505-3514. 10.1016/S0009-2509(00)00011-7View ArticleGoogle Scholar
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