Sensing Conductive Hydrogels for Rapid Detection of Cytokines in … · 2016. 3. 18. · ©2016...

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Sensing Conductive Hydrogels for Rapid Detection of Cytokines in Blood

Dong-Sik Shin, Zimple Matharu, Jungmok You, Christian Siltanen, Tam Vu, Vijay Krishna Raghunathan, Gulnaz Stybayeva, Ashley E. Hill, and Alexander Revzin*

Prof. D.-S. Shin, Dr. Z. Matharu, Prof. J. You, C. Siltanen, T. Vu, Dr. G. Stybayeva, Prof. A. RevzinDepartment of Biomedical EngineeringUniversity of CaliforniaDavis, CA 95616, USAE-mail: [email protected]. D.-S. ShinDepartment of Medical & Pharmaceutical SciencesSookmyung Women’s UniversitySeoul 140-742, Republic of KoreaDr. Z. MatharuDepartment of Electrical and Computer EngineeringUniversity of CaliforniaDavis, CA 95616, USAProf. J. YouDepartment of Plant & Environmental New ResourcesKyung Hee UniversityYongin 446-701, Republic of KoreaDr. V. K. RaghunathanDepartment of Surgical & Radiological SciencesSchool of Veterinary MedicineUniversity of CaliforniaDavis, CA 95616, USAProf. A. E. HillCalifornia Animal Health and Food Safety LaboratoryUniversity of CaliforniaDavis, CA 95616, USA

DOI: 10.1002/adhm.201500571

To prove this concept we chose to detect bovine-interferon-γ (B-IFN-γ)—an inflammatory cytokine produced by T-cells that can be used for diagnosing tuberculosis (TB) in cows. Caused by Mycobacterium bovis, bovine TB is a chronic bacterial dis-ease that is communicable to humans through aerosol or con-sumption of unpasteurized milk.[8] Therefore, development of new means for detecting bovine TB has relevance to human health. At present, detection of B-IFN-γ is carried out by enzyme linked immunosorbent assay (ELISA); however, ELISA calls for a multistep protocol involving centrifugation, washing station, and plate readers. Therefore, this test is performed in a cen-tralized laboratory. We wanted to utilize conductive hydrogels for the detection of IFN-γ without the need for benchtop instru-mentation such as centrifuges as well as washing and labeling steps. As shown in Scheme 1, this biosensor consisted of con-ductive hydrogel layer impregnated with cytokine-specific anti-bodies (Abs) residing on top of an electrode. Capture of antigen molecules (IFN-γ in our case) caused a quantifiable change in electrochemical redox properties, providing the basis for spe-cific and sensitive detection. While there have been previous reports of conductive polymers functionalized with Abs for detection of antigens,[9] our study takes this concept to the next level by creating a conductive sensing material that possesses mechanical and nonfouling properties of a hydrogel making it durable and functional in a complex solution such as whole blood. Discussed below are the various aspects of the project leading up to the functional biosensor, including: (1) conductive hydrogel synthesis, (2) characterization of chemical composi-tion and mechanical properties, (3) demonstration of sensitivity and specificity in response to B-IFN-γ, and (4) detection of IFN-γ in bovine blood.

A miniature sensing electrode array is shown in Figure 1A. PEG hydrogel is photopolymerized on top of miniature gold or indium tin oxide (ITO) electrodes (diameter = 1.5 mm and PEG diameter = 1.7 mm). The glass substrates containing electrode arrays were silanized using (3-acryloxypropyl) trichlo-rosilane prior to gel patterning to promote PEG attachment.[10] PEG gel coated electrodes were placed into custom-made elec-trochemical cell and used as working electrodes (Scheme S1, Supporting Information). Ag/AgCl reference electrode and Pt wire counter electrode were placed into electrochemical cells, after which 1.1 V potential was applied to drive electropoly-merization reaction. Formation of PEDOT is associated with a characteristic color change (from transparent white to bluish)[11] and could be easily visualized, as shown in Figure 1B,C. The quality of the conductive film is a function of several factors including monomer concentrations, the types of dopants used, and duration of electropolymerization. We have converged on an aqueous solution containing EDOT and EDOT-COOH

Electrically conducting polymers are promising coating mate-rials for biomolecular electronics applications such as immuno-biosensors, DNA biosensors, neuron signaling, and temporal growth factor release.[1] However, when deposited as thin films these materials are often brittle[2] and are easily fouled through nonspecific adsorption of proteins or cells. The goals of this study were to (1) design a novel hydrogel that is conductive, nonfouling, and nonbrittle, and (2) deploy this hydrogel coating in the development of a biosensor for detection of cytokines in blood. To achieve these objectives we focused on combining poly(ethylene glycol) (PEG) hydrogels with poly(3,4-ethylenedi-oxythiphene) (PEDOT). Of a number of conducting polymers described in the literature,[3] PEDOT is often preferred due to its biocompatibility,[4] electrochemical stability,[5] and ease of modulating electrical properties with dopants.[6] PEG on the other hand is a biocompatible polymer known for its excellent nonfouling properties.[7] We posited that by combining PEDOT and PEG one may be able to create an electrode coating that will be conductive and nonfouling, permitting for electrochemi cal measurements to be carried out in complex media such as blood.

Adv. Healthcare Mater. 2016, DOI: 10.1002/adhm.201500571

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http://doi.wiley.com/10.1002/adhm.201500571

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monomers along with a dopant LiClO4 as optimal for elec-tropolymerization (optimization data not shown). We found that keeping the time of electropolymerization constant yielded gels of varied redox properties, likely due to slight variations in thickness of the PEG hydrogel. Instead we chose to keep deposited charge constant and found 0.8 mC to result in gel films with optimal redox properties (prominent reduction peaks observed in Figure S1, Supporting Information).

It must be noted that COOH functionalized EDOT provides functional groups for antibody attachment via 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide-N-hydroxy-succinimide (EDC-NHS) crosslinking while EDOT monomer

assists PEDOT-COOH electropolymerization. A conduc-tive polymer with 1 to 4 molar ratio of EDOT (2.5 × 10−3 m) to EDOT-COOH (10 × 10−3 m) provided the right balance of maximizing the number of Ab binding sites while enabling electropolymerization.

Once electropolymerization of PEDOT in PEG gel was opti-mized, the number of COOH groups was quantified. These moieties are very important as they provide handles for cova-lent attachment of Abs (described later). The quantification of COOH groups was performed using Toluidine Blue O (TBO) staining (Figure S2, Supporting Information)[12] and was deter-mined to be 3.6 × 1016 molecules cm−2 and 1.7 × 1016 molecules



Scheme 1. Fabrication of sensing PEDOT/PEG hydrogels. Step 1. PEG gel film formed on electrode surface. PEDOT is electropolymerized in the PEG gel, making it conductive. Step 2. Biorecognition molecules (Ab) are incorporated into the gel via PEDOT-COOH groups. Step 3. Capture of B-IFN-γ molecules causes a change in electrical signal of the gel. Step 4. Change of reduction peak current is recorded as analyte binds.

Figure 1. Conductive hydrogels on gold and ITO electrodes. A) An array of gold electrodes (with leads and contact pads) used for depositing and characterizing conductive hydrogels in our study. PEG gel was photopolymerized on top of the electrodes. For PEDOT deposition, electrode arrays were placed into an electrochemical cell such that only electrodes were exposed. Scale bar: 3 mm. B–C) Electropolymerization of PEDOT inside PEG gel. Optically transparent ITO electrodes were used for ease of visualization. B) Note that larger of the two concentric circles is the PEG gel disk, smaller is ITO electrode. C) As PEDOT polymerize the gel becomes dark blue. Scale bar: 300 μm. D,E) SEM images reveal differences in morphology of pure D) PEDOT and E) PEDOT/PEG conductive hydrogel.


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cm−2 for PEDOT/PEG and PEDOT, respectively. This result may be interpreted as either pointing to the presence of more EDOT-COOH repeat units in PEG/PEDOT coating or to better accessibility of these functional groups in the gel. In either case, higher density of COOH groups is beneficial for improved loading of Ab molecules into the sensing layer.

The electrode surfaces covered with conductive coating were characterized and compared using scanning electron micro-scopy (SEM). PEDOT films deposited on gold electrodes had rough and granular morphology (Figure 1D), whereas PEDOT/PEG hydrogel coatings on gold were smooth (Figure 1E). In fact, PEDOT/PEG layer had similar morphology to PEG gel (data not shown). Also, morphology of conductive hydrogels formed on ITO and Au electrodes was similar (data not shown).

It is thought that conductive polymers can grow vertically from the electrode surface, spanning cross-section of the gel.[13] Other studies investigating hydrogel properties suggested that incorporation of conductive polymer had only minimal effect on the water uptake of the gel.[14] Therefore, it is likely that physical properties of the PEG gel in our study (mesh size and water content) were minimally affected by incorporation of con-ductive polymer. As discussed below no statistically significant differences were observed between mechanical properties of conductive hydrogels and pure PEG gels.

Mechanical properties of conductive coatings were analyzed by atomic force microscopy. These experiments (Figure 2) revealed elastic moduli to be 4.95 ± 0.25 kPa (mean ± standard error), 1.28 × 103 ± 149.50 kPa, and 7.58 ± 0.84 kPa for PEG, PEDOT, and PEDOT/PEG, respectively. While the difference

in modulus of PEDOT/PEG and PEG was not statistically sig-nificant, PEDOT/PEG gel modulus was 170 times lower than the modulus of pure PEDOT. These results are consistent with previous reports of PEDOT possessing high elastic modulus and being brittle.[2] These results also highlight the fact that mechanical properties of our conductive hydrogels are similar to pure PEG hydrogels. As discussed in the next paragraph, conductivity of the composite gel was similar to that of pure PEDOT, suggesting that our hydrogel strikes the right balance of materials properties and holds significant promise for bio-logical applications.

The goal of the next set of experiments was to investigate electrochemical properties of conductive hydrogels. Analysis of PEG gel-coated electrodes by cyclic voltammetry (CV) revealed appearance of characteristic redox peaks upon electropolym-erization of PEDOT (Figure 3A).[15] As seen from these data, electrochemical properties of conductive hydrogel coating were practically indistinguishable from those of a PEDOT film while PEG gel was not electroactive. To summarize our find-ings up to this point, PEG/PEDOT hydrogels possessed supe-rior mechanical properties, had more sites for Ab conjugation, and showed similar electrochemical properties to pure PEDOT films. The next sets of experiments were aimed at the construc-tion and characterization of biosensors using these conductive hydrogels.

The development of a biosensor began with covalent conju-gation of IFN-γ Abs using EDC-NHS chemistry.[16] Incubation with Ab molecules resulted in 10% to 15% decrease in reduc-tion peak current indicating Ab molecules became incorpo-rated into the gel. This change in redox properties was used as an indicator of Ab immobilization. The redox properties, or more specifically prominent reduction peak appearing at −0.3V versus Ag/AgCl, were used as basis for sensing. Spiking IFN-γ into PBS caused the reduction peak to decrease proportionally with IFN-γ concentration (Figure 3B,C). While precise sensing mechanism is yet to be fully elucidated, it is likely that capture of IFN-γ molecules hindered electron transfer through the PEDOT chains. An indirect validation of this hypothesis was an observation that only conductive gels formed from EDOT-COOH monomers and functionalized with Abs were respon-sive to IFN-γ. Hydrogels with COOH groups present on PEG as opposed to PEDOT did not change redox properties in response to the analyte. This suggested that antigen-Ab binding needs to happen proximally to PEDOT chains for redox properties to be affected.

To confirm that IFN-γ was indeed captured in the conductive hydrogel we developed fluorescence-based sandwich immuno-assay utilizing sensing gel coated on optically transparent ITO electrodes. As shown in Figure S3 (Supporting Information), fluorescence signal from the electrode increased as a function of analyte concentration further validating that sensing gels were indeed capturing B-IFN-γ molecules.

To test specificity, sensing hydrogel electrodes were chal-lenged with various nonspecific cytokines and proteins com-monly present in serum, including TGF-β, IL-6, TNF-α, and IgG. As seen from Figure 4A, in most cases the sensing elec-trodes showed <10% response to nonspecific proteins present at the same concentration (100 ng mL−1) as bovine IFN-γ. In fact, sensing electrodes designed to detect bovine IFN-γ were



Figure 2. Box plot of elastic moduli PEDOT/PEG measured by AFM: elastic modulus profiles obtained from PEG, PEDOT only, and PEDOT/PEG hydrogels obtained from a minimum of n = 75 force curves for each sample. The red line indicates mean, while the solid dots rep-resent outliers. Y-axis is adjusted to normal logarithmic scale (base 10). ***p < 0.001 compared with PEG sample, ANOVA followed by Dun-nett’s multiple comparison test. No differences in elastic moduli were observed between PEG and PEDOT/PEG samples, while the modulus of PEDOT only samples is at least three orders of magnitude higher. This implies that the hydrogels, while able to possess the conductive proper-ties of PEDOT, exhibit the intrinsic mechanical properties of native PEG hydrogels.


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only minimally responsive when challenged with human IFN-γ (see Figure 4A). It should be noted that we attempted to fabri-cate Ab-conjugated PEDOT films for head-to-head comparison with PEG/PEDOT-coated electrodes. However, our attempts were unsuccessful due to delamination of PEDOT film from gold and ITO electrodes (see Figure S4, Supporting Informa-tion, for example). The fragility of PEDOT-covered electrodes was in stark contrast to robustness of conductive hydrogels and underscored the utility of the latter material for biosensing applications.

The key challenge for electrochemical biosensors oper-ating in complex samples such as blood is fouling of electrode through deposition of proteins or cells. The process of fouling may change electrical properties of the bioelectrode and may make it impossible to decouple specific response of the sensor from changes in its electrical properties due to nonspecific adsorption. The challenge of electrode fouling is commonly resolved by creating a separate filtering or nonfouling layer that resides on top of the biorecognition layer and protects bioelectrode during measurements. In contrast, we describe

the development of a single composite hydrogel layer that combines biorecognition and nonfouling functions. To under-score this point, we challenged hydrogel-coated electrodes with bovine blood diluted 1:1 in PBS. As seen from Figure 4B, while reduction peak shifted toward a more negative potential (from −0.35 to −0.6 V versus Ag/AgCl), its amplitude did not change and was comparable to the amplitude of reduction peak in pristine buffer. For comparison, in our previous work with gold electrodes functionalized with methylene blue-labeled aptamers, we observed approximately sixfold decrease in the amplitude of redox current.[17] To rescue redox activity of an electrode a filter needed to be placed atop the sensing electrode (unpublished results). Herein we demonstrate a biosensor that retains redox activity upon exposure to blood without the need for additional filtering. While shift of reduction peak to a more negative potential is not ideal, we should note that the sensor still operates in the safe zone where electrochemical interfer-ence is not significant. Typically electrochemical interference in blood comes from oxidation of uric and ascorbic acid at higher positive potentials (>0.6 V versus Ag/AgCl).



Figure 3. Electrochemical properties and IFN-γ responsiveness of conductive hydrogels. A) Cyclic voltammetry (CV) comparison of electrodes cov-ered with PEG gel, PEDOT, and PEDOT/PEG gel. Note that the redox activity of pure PEDOT is similar to that of PEDOT impregnated into PEG gel. B) Challenging Ab-carrying conductive hydrogels with different concentrations of B-IFN-γ results in diminution of reduction peak current. C) A calibration curve may be constructed by plotting changes in reduction peak current (Ip) versus concentration of B-IFN-γ.


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This suggests that redox properties and conductivity of the gel coating were only minimally affected by exposure to blood. Spiking different IFN-γ concentrations into bovine blood caused the reduc-tion peak to decrease in a concentration-dependent manner.

As seen from the calibration curve in Figure 4C, the limit of detection for IFN-γ in whole blood was 1 ng mL−1 compared to 0.5 ng mL−1 in pristine buffer (Figure 3B). One should note that sensor saturation in blood occurred at ≈50 ng mL−1 IFN-γ, whereas this point was not reached with 150 ng mL−1 in pris-tine buffer.

While the reasons for decreased dynamic range in blood need to be investigated further, it is possible that fewer binding sites are available in complex media due to coating of the gel with serum proteins. Overall, despite changes in electrochem-ical properties and dynamic range, we conclude that electrodes coated with conductive hydrogel layer remain functional and sensitive in bovine blood.

As a final demonstration of the advantages of the biosensor developed here, we monitored temporal dynamics of IFN-γ

release from leukocytes. Temporal aspects of cellular secretions may be important for immunology applications for example in determining how primed or activated immune cells are and how fast the immune cells are able to process and respond to specific antigens. Our lab and others have been developing aptamer-based electrodes for dynamic detection of cell secretions.[18] While aptamers are promising as biorecognition elements, only a limited repertoire of aptamers have been developed to date. The sensing strategy described here relies on monoclonal antibodies as biorecognition elements and therefore has much wider immediate impact for label-free, dynamic detection of cell function. Figure 4D demonstrates temporal monitoring of IFN-γ release from bovine leukocytes (peripheral blood mononuclear cells or PBMCs) stimulated with mitogens. As seen from these data, biosensing conductive hydrogels allowed for monitoring the production of IFN-γ from leukocytes over time by recording changes in reduction current associated with PEDOT peak.

In conclusion, this study demonstrates the development of a novel composite hydrogel designed for sensing in complex

Figure 4. Assessing specificity of conductive sensing hydrogels and their performance in blood. A) Specificity is determined by comparing responses of gel bioelectrode to nonspecific proteins and B-IFN-γ spiked at the same concentration (100 ng mL−1) into cell culture medium (RPMI 1640 w/o serum). B) Comparing redox properties (reduction current) of conductive hydrogel in RPMI1640 and whole bovine whole blood. As seen from these results, exposure to bovine whole blood did not significantly alter redox properties. C) Calibration curve of biosensor constructed by spiking known concentrations of B-IFN-γ into bovine whole blood. D) Monitoring of cytokine release from mitogen activated peripheral blood mononuclear cells (PBMCs). These data demonstrate that Ab-impregnated conductive hydrogels may be used for dynamic detection of cell secretions.


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samples such as blood. This material is similar to PEDOT in terms of conductivity but possesses mechanical properties of a hydrogel providing for a robust electrode coating layer. When functionalized with Abs, our conductive hydrogel may be used for sensing antigens/analytes in a label-free manner where redox signal intrinsic to PEDOT in the gel changes directly upon addition or appearance of the analyte. Importantly, we demonstrate that conductive hydrogel combines specificity and sensitivity for analyte with nonfouling properties to allow for operation in blood or other complex samples. Most importantly, our conductive hydrogels utilize monoclonal antibodies to convert antigen binding events into an electrochemical signal. Therefore, we envision the conductive hydrogel described here as a sensing platform technology where specificity to the pro-tein of interest may be tuned by selection and incorporation of appropriate monoclonal antibodies.

Supporting InformationSupporting Information is available from the Wiley Online Library or from the author.

AcknowledgementsD.-S.S. and Z.M. contributed equally to this work. Financial support for this work was provided by grants from California Dairy Research Foundation (Grant No. P14-003-UCD-AR-SS). Additional funding came from “Research Investments in Science Engineering” from UC Davis.

Received: July 22, 2015Revised: November 16, 2015

Published online:

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