J Electroanal Chem.2016 Aug;775:135-145
Layer-by-layer Constructed First Generation Uric Acid Biosensors Using Xerogels
Grace E. Conway, Raef H. Lambertson, Margaret A. Schwarzmann, Michael J. Pannell, Helene W. Kerins, Kristen J. Rubenstein, Jonathan D. Dattelbaum, Michael C. Leopold.
Department of Chemistry, Gottwald Center for the Sciences, University of Richmond , Richmond, Virginia 23173, United States.
Objectives: Developing a strategy of modifying an electrode with a specific combination of polymeric and xerogel materials to create an effective layer-by-layer (LbL) constructed 1st generation biosensor for uric acid (UA). UA, the predominant end product of purine nucleotide catabolism in the human body, is a clinically relevant molecule implemented in pregnancy induced hypertension (PIH) diagnosis.
Methods: Electrochemically cleaned platinum working electrodes were modified using four functional layers (Figure 1): (1) an initial enzyme-doped xerogel layer, (2) a diffusion-limiting xerogel layer, (3) an inner selective electropolymer layer, and (4) an outer polyurethane (PU) selective membrane. Each layer was systematically investigated and its properties specifically tailored for UA permeability and interferent discrimination. The influence of the PU hydrophobic character and its UA permeability was established while the inner enzyme-doped and outer diffusional xerogel layers were evaluated for uricase (UOx) species/loading and silane precursor dependence, respectively. Complete sensors were held at +0.65 V during standard UA or interferent injections while monitoring current response over time, a current-time (I-t) curve.
Results: LbL systematic evaluation revealed the specific combination of hydroxylmethyl tri-ethoxy silane (HMTES) xerogel bi-layer, a polyluminol-aniline (PL-A) electropolymer, and 100% hydrophilic polyurethane yielded impressive UA sensing performance: effective sensitivity (0.8 nA/μM), linear response across physiologically relevant UA concentrations (100–700 μM), fast response times (~10 s), low limits of detection (< 10 μM), and effective selectivity against most common interferents. Toward the specific application of PIH risk assessment, the optimized sensor exhibited 10-day stability as well as effective shelf-life exceeding 35 days.
Conclusions: The presented system rivals or exceeds UA biosensor performance found in the literature and offers the possibility of miniaturization for in situ or in vivo remote diagnostic sensing. The successful adaptation suggests that the strategy and materials may be applicable to detecting/monitoring other medically significant molecules via sensor development.
Over the past couple of decades, fundamental sensor research has disproportionately focused on the development of glucose biosensing schemes (1, 2) due to the increasing need for improved diabetes management as well as the idea that glucose represents a well-understood, relatively inexpensive model system for fundamental sensor studies (3). In most cases, however, the systematic transition of adapting an effective strategy, materials, and constructs for glucose biosensing to a dramatically different target molecule, such as uric acid has not been adequately demonstrated and is largely unexplored.
Uric acid (UA), the predominant end product of purine nucleotide catabolism in the human body, present in human blood serum, plasma, urine, and saliva. Due to its poor solubility in biological fluids, it exists predominantly as urate anion at physiological pH (4, 5). UA imbalances in the body have been linked to many diseases. Elevated levels of UA in blood serum are a significant independent predictor for atherosclerotic cardiovascular disease in high-risk individuals (6), peripheral artery disease (7), and chronic kidney disease (8). Higher concentrations of UA in blood serum have also been shown to be a risk factor for non-alcoholic fatty liver disease (9) and a risk factor for silent brain infarction (10). Long-standing hyperuricemia also contributes to gout, which is caused by the nucleation and growth of mono-sodium urate crystals in tissues in and around the joints following the saturation threshold of UA in serum (11). Hyperuricemia is one of the earliest and most consistent observations noted in pregnancy-induced hypertension (PIH), a condition that can lead to a dangerous disorder called pre-eclampsia (12). Elevated concentrations of UA have been shown to effectively identify a subset of women with PIH who are at greater risk for maternal or fetal morbidities from pre-eclampsia. An effective UA biosensor with the potential for in-vitro remote sensing or in-vivo operation as an intravenous implant with selectivity and sensitivity to UA between 0.1 and 0.6 mM would accelerate risk assessment of pregnancies by identifying higher risk PIH patients, expediting drug therapy and/or surgical intervention – all of which lower perinatal mortality rates and reduce risk to mothers as well as lower treatment cost.
Figure 1. Scheme of a layer-by-layer UA biosensor; and typical i-t curve during successive injections of UA.
Figure 2. (A) Calibration curves during successive 1 mL injections of UA standard (1.86 mM) at a Pt electrode modified OTMS xerogel bilayers capped with PU (100% HPU) where the first OTMS xerogel layer is doped with different species of uricase enzyme* (UOx) including: Bacillus fastidiosus (BF), BF (Sigma) and Candida sp. (Sigma); (B) Calibration curves during successive 1 mL injections of uric acid standard (1.86 mM) at a platinum electrode modified OTMS xerogel bilayers with (closed symbols) and without (open symbols) a PU (100% HPU) capping layer where the first OTMS layer has different concentrations of uricase enzyme (UOx) created from different loadings (3, 6, or 9 mg of UOx) of the deposition mixture.
UA biosensing schemes based on amperometric detection have been developed using a range of materials including uricase (UOx) enzyme, polymer and self-assembled monolayer modified electrodes as well as electrodes incorporating carbon paste or nanomaterials(13, 14). Our research focuses on layer-by-layer (LbL) construction of a highly effective 1st generation amperometric biosensor for UA (Figure 1). While this study stems from prior, proof-of-concept work with a model glucose biosensing system, it represents a significant advancement in sensor research and development, providing a strategy and materials optimization that can be adapted to a clinically relevant target species of interest. With the construction of LbL system for UA detection, we hypothesized that each of the components of the LbL system either individually or in concert has a significant effect on the selectivity and/or sensitivity of UA detection. This study presents systematic characterization and evaluation of each functional layer within the sensing scheme which, for the first time, establishes a thorough understanding of each material comprising these layers. The work is supportive that the overall strategy itself can be optimized toward a specific target species other than the glucose model. That is, systematic variations of the materials involved may allow for strategic design of schemes flexible enough to be adapted toward a variety of target molecules, thereby spawning the development of additional diagnostic sensors. This study culminates with the development and demonstrated performance of a highly functional UA biosensor that features high sensitivity, effective selectivity, and excellent stability attributes – collectively an unmatched combination of performance parameters and a system that maintains the potential to be further developed for in-vitro or in-vivo operation in PIH risk assessment (12).
The general design of the LbL biosensing scheme is depicted in Figure 1. To identify the suitable uricase enzyme type for the system performance, careful examination was made using various UOx enzymes doped within octyl-trimethoxysilane (OTMS) xerogel, a system previous employed for glucose sensing films. Results indicate that the in-house optimized B. fastidiosus UOx enzyme showed significant sensitivity over a wide concentration range (Figure 2A). Similarly, the effect of UOx loading (i.e., enzyme concentration) in the first xerogel layer on the sensitivity of the constructed biosensor was investigated. Results showed that 3-4 mg of UOx is sufficient for effective UA sensitivity (Figure 2B).
To analyze the effect of hydrophobicity of the outer semi-permeable polyurethane (PU), we used a blend of hydrophilic polyurethane (HPU) and hydrophobic polyurethane (TPU) based on our previous work and that of other labs investigating glucose sensing strategies (3, 15). Cyclic voltammetry of ferricyanide (FeCN) in solution at these interfaces, along with hydrogen peroxide (H2O2) amperometric injections were used to assess the porosity and properties of the different PU blends. This study revealed that the PU layer most conducive to FeCN redox probes and H2O2 oxidation is the 100% HPU coating (Figure 3A), which also shows a substantial difference in the morphology from the 100% TPU according to SEM and AFM microscopy (Figure 3B, C). Ionic strength (μ) of the sample matrix is particularly important to evaluate given the goals of eventually using these findings toward in-vitro, remote bedside sensors or in-vivo implantable devices for continuous UA monitoring. Electrochemical sensing is dependent on solution activity rather than simple concentration, making evaluation of the sensor at biological levels of μ an important facet of the sensor development. To evaluate m dependence of the sensor response to UA, OTMS bi-layer systems were used. The UA sensitivity of the composite film increases with μ to a point (65.55 mM; μ=150 mM), which is similar to blood plasma, before plateauing at even higher values (Figure 3D). Finally, the effect of varying xerogel bilayer system was investigated. To this effect, based on prior evaluation and reporting of silane precursors for xerogel formation (3), a subset of silane precursors, including HMTES, OTMS, IBTMS, PTMS, and APTMS, were identified for UA biosensing development (Figure 4A). This study revealed that HMTES xerogel full system to be the most effective optimized system tested in terms of both sensitivity and interferent discrimination (Figure 4B).
Through the development of a viable UA sensor poised for subsequent development toward commercialization, this study has presented an unprecedented level of detail evaluating the LbL strategy and materials involved in xerogel-based 1st generation biosensor design. Based on this work, biomedical devices, perhaps incorporated at the tip of an intravenous needle/wire within a catheter (Figure 4C), capable of real-time bedside monitoring of UA levels can be envisioned. The development of such devices could aid in the detection, monitoring, and treatment of PIH thereby decreasing both cost and risk for expectant mothers and their newborns.
Importance of the study: Our data indicates a successful adaptation of the LbL xerogel-based sensing scheme developed from the glucose model system to a new and separate target molecule of clinical relevance, UA. The study provides a functional sensor prototype that can be further developed and miniaturized for commercial application as well as a strategic template for designed additional sensors for other medically relevant target molecules.
Figure 3. (A) Summary comparison of the oxidative peak current for FeCN voltammetry (■, ip,a) verses the average amperometric response during H2O2 injections (▲, ia) at various PU coatings; Scanning electron microscopy images and atomic force microscopy images (insets) of (B) 100% TPU (hydrophobic) and (C) 100% HPU (hydrophilic) polyurethane coated substrates; and (D) uric Acid sensitivity (μA·μM-1) dependence on ionic strength.
Figure 4. (A) Calibration curves during successive 1 mL injections of UA standard at a Pt electrode modified with various UOx-doped xerogels, undoped xerogels, as well as selective membranes of electropolymerized PL-A film and PU (100% HPU); (B) Specific i-t curve example for successive injections of common interferent species and uric acid at the full HMTES xerogel system; (C) example of a needle electrode.
This research was generously supported by funding from the National Science Foundation (CHE-1401593), Virginia’s Commonwealth Health Research Board, and the Beckman Foundation (MJP) as well as student fellowships from the University of Richmond’s Undergraduate Research Committee (GEC, RHL). We would like to thank Dr. Mulugeta Wayu for his important contributions to this work as well.
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