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Electrophoresis 2015, 00, 1–6
ˇ Pavel Podesva ˇ Frantisek Foret Institute of Analytical Chemistry AS CR, v.v.i., Brno, Czech Republic
Received April 13, 2015 Revised May 13, 2015 Accepted May 13, 2015
Research Article
Metal nano-film resistivity chemical sensor In this work, we present a study on reusable thin metal film resistivity-based sensor for direct measurement of binding of thiol containing molecules in liquid samples. While in bulk conductors the DC current is not influenced by the surface events to a measureable degree in a thin metal layer the electrons close to the surface conduct a significant part of electricity and are influenced by the surface interactions. In this study, the thickness of the gold layer was kept below 100 nm resulting in easily measureable resistivity changes of the metal element upon a surface SH-groups binding. No further surface modifications were necessary. Thin film gold layers deposited on a glass substrate by vacuum sputtering were photolithographically structured into four sensing elements arranged in a Wheatstone bridge to compensate for resistance fluctuations due to the temperature changes. Concentrations as low 100 pM provided measureable signals. The surface after the measurement could be electrolytically regenerated for next measurements. Keywords: Adsorption / Chemiresistor / Nano-film / Thiol binding / Thiol sensing DOI 10.1002/elps.201500190
1 Introduction The main use of thin metal layers in bioanalysis relates to the measurements based on the surface plasmon resonance (SPR). Typically a well conducting thin metal layer (gold, silver) is deposited on a glass surface and illuminated by a light source from the glass side. The reflected light intensity is monitored. The surface plasmon, generated at certain angle, responds to the surface refractive index changes allowing detection of the binding events. The evanescence field penetrating a fraction of the wavelength above the metal surface allows monitoring of not only the primary binding on the metal surface, but provides information about binding events within the evanescence field distance. Thus it is possible to monitor not only the primary binding but also the affinity interactions on the chemically modified surface. Surface plasmon resonance is one of the best developed optical techniques for affinity-binding studies, especially for biomolecular research [1,2] and commercial instrumentation is available from a number of manufacturers. On the other hand, the SPR instrument is relatively complex and provides only an indirect optical information about the binding events with little information about the first chemical layer binding. Monitoring of basic electric properties (current, voltage) represents one of the most precise measurement techniques. Besides the simplicity and precision, electrical measurements also provide a large dynamic range. The related equipment is
ˇ Correspondence: Dr. Pavel Podesva, Institute of Analytical Chemistry AS CR, v.v.i., Veveri 97, 902 00 Brno, Czech Republic E-mail: [email protected]
Abbreviations: GSH, glutathione; HDMS, hexamethyldisilasane; SPR, surface plasmon resonance C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
readily available and inexpensive, resulting in the widespread use not only in electronics measurements but also in a variety of sensors used, e.g. in physics, chemistry, environmental sciences, or medicine. While many physico–chemical interactions can be readily converted into electric signals using a suitable sensor (e.g. in optical sensing), other interactions result in a direct electric change upon the interaction with the sensor material. For example, chemiresistors, a family of electronic sensing components, change their electric resistance depending on the physic-chemical processes occurring at its active surface, including charge induced field effect [3], formation or dissolution of self-assembling monolayers [4], swelling of a conductive layer [5], molecular or ionic adsorption-desorption, redox changes [6, 7], etc. Typical chemiresistor is an inexpensive, sensitive, small, robust, and low power device commonly used for detection in gas or liquid phases [8, 9]. The metal layer itself is generally not considered as a suitable sensing element for chemiresistors since its ability to conduct DC electric current is given by all electrons available in the bulk metal and the contribution of any surface events is negligible. However, as the thickness of the metal layer decreases, the contribution of the electrons close to the surface to the total conductivity of the material increases and their mobility and scattering depends on the surface interactions. While this effect has been theoretically described over 60 years ago [10], only few papers have investigated the effect experimentally [11] and to our knowledge there is no work describing a practical application of a simple thin metal sensing probe except to our previous report on the free floating nanoelectrodes for electrochemiluminiscence generation [12]. The strength of the Au–S bond is about 50% of the C–C bond (170–210 kJ/mol) [13, 14] and, once formed, the release
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of the attached compounds in water or acetate buffer environment (pH = 5–7) is slow, on the order of days [15]. The full coverage of the 40 nm gold film sensor by the monolayer of the thiol compound leads to the increase of resistance on the order of 4% largely independent on the type of organic residue connected to the SH group [16, 17]. This resistivity change of the layer is; however, influenced by the number of sulphur atoms connected to the surface. In this work, we have investigated the DC electric resistance of thin (45–60 nm) Au films upon binding of organic thiol containing compounds in water solutions. The Au sensing elements were formed either as a part of a flow-through microfluidic device or as a droplet sensing area. By real-time resistance monitoring of the Au element is was possible to observe the rate of coverage during the experiment. By applying the electric potential between the working and counter electrode it was also possible to release the attached thiol containing species and reuse the Au element for further measurements.
2 Materials and methods 2.1 Reagents and equipment All reagents were of analytical grade quality. (–)-Glutathione, oxidized 98% (HPLC) (Sigma) (GSSG); (–)-Glutathione, reduced 98.0% (Sigma-Aldrich) (GSH); Gly-Pro-Glu 98% (HPLC) (CAS 32302-76-4) (GPG); Hexamethyldisilasane reagent grade, 99% (HDMS) were purchased from SigmaAldrich inc. Sylgard 184 silicone elastomer was purchased from Dow Corning GmbH (made in USA). Acetic acid 98% p.a. and Toluene p.a. from Lach-Ner, s.r.o., Czech Republic. Acetonitrile p.a. was purchased from Penta chemicals, Czech Republic. MicropositTM LOLTM 2000 Lift Off Layer (LOL 2000) and MicropositTM remover 1165 were purchased from Rohm and Haas Electronic Materials LLC. The ma-P 1225 photoresist and ma-P 1225 developer were purchased from Micro resist technology GmbH, Germany; Nanofilm glass – Borofloat 1.5 mm, 1000A˚ Cr layer, 5300A˚ AZ1518 Photoresist layer from Nanofilm Inc. USA; 3 inch 0.65 mm borofloat glass wafers from S. I. Howard Glass, Cambridge, MA, USA; Demineralized water (18 M⍀/cm) was used for all experiments. The Au layers were prepared by vacuum magnetron sputtering using Bal-Tec SCD 500 equipped with a QCM thickness monitor (Bal-Tec, Schalksm¨uhle, Germany). Clarity Lite (DataApex, Prague, Czech Republic) two channel A/D converter module and a laboratory constructed nanomicroampere adjustable constant current source was used for resistance monitoring. For improved hydrophilic properties the glass surfaces were modified using a radiofrequency plasma pen developed at the Masaryk University, Brno, Czech Republic [18].
2.2 Device fabrication The 3 inch borosilicate substrate was consecutively cleaned with detergent, water, piranha solution (H2 SO4 :H2 O2 /3:1), C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
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washed thoroughly with water, dried with nitrogen stream, baked at hotplate for 1 h at 170°C, then treated in 1% solution of HDMS (submersion) for 2 min and left to dry. The first spin-coated layer was the LOL 2000 spun at 3000 rpm for 30 s followed by baking for 5 min at 170°C on a hotplate. The second layer was Ma−P 1225 spin-coated for 30 s at 3000 rpm and baked at 100°C for 90 s. The wafer was exposed with PG 101 lithograph, developed in ma-P 1225 developer, washed with water and dried in a clean air stream. Next, a 3 nm indium-tin oxide adhesion layer and 45 nm of gold was sputtered on the wafer using the Bal-Tec SCD 500. Resists were then stripped with Microposit remover 1165. The resulting wafer with the Au motive was finally annealed in furnace for 240 min at 450°C in the air atmosphere to decrease porosity and reduce resistance by 50% [19]. Resistance of each resistor in the bridge was 790 ⍀ after annealing. Two electro-mechanical arrangements have been prepared. In both cases the gold surface sensor was arranged in a serpentine shape to provide desired resistance. The serpentine layout was complemented by a comb-like electrode interdigitated into the measuring electrode meanders. This electrode was tested as a cleaning electrode for electrochemical cleaning of the sensor after the measurement. The first design was arranged as a flow through device for on-line operation while the second device was prepared as an open structure allowing measurement in liquid drops pipetted on the electrode surface. The microfluidic part of the first design was casted from PDMS resin using a glass master microfabricated by photolithography on the Nanofilm glass. The resulting fluidic channels were 50 m deep. The form for casting of the PDMS part was milled in a PMMA sheet and the glass master was treated by HDMS for easy separation from the cured PDMS resin. Fused silica capillaries (100 m id × 360 m od, Polymicro Technologies, USA) were connected to the PDMS block with the LabSmith (LabSmith, Inc.; Livermore, CA, USA) ferrules type C360100 half-casted into PDMS block to ensure tight input/output without large void volumes. Completion of the chip was performed using plasma pencil [18] producing a stream of cool O3 containing Ar plasma for activation of the surface of the glass and PDMS to the Si-OH state. The time needed for sufficient treatment of both the PDMS and glass surfaces was about 10 s at which both surfaces were strongly hydrophilic with almost zero contact angle (checked visually with drop of water). In the next step, the PDMS part was placed on the glass sensor and cured for 5 h on a 100°C hotplate, leading to a permanent bond between the glass and PDMS. For easier manipulation with the activated PDMS block before pressing it onto the activated glass few drops of the methanol were applied on its surface [20]. This allowed precise adjustment of the PDMS block position before establishing the permanent contact. In the next step, ferrules with capillaries were inserted into port holes. For best mechanical endurance edges of the PDMS as well as the top sides of the ferrules were sealed by PDMS to form one compact piece. www.electrophoresis-journal.com
Microfluidics and Miniaturization
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The open surface device was fabricated on the same glass substrate as the previous one. Also the sensor circuit is the same with the only difference being the round shape of the resistive elements. The surface of the final device except the sensor pads was coated by hydrophobic layer of Teflon (PTFE Disp. 30, DuPont, Wilmington, DE, USA). During the coating the sensor pads were protected by a photoresist pattern and after hardening of the Teflon layer the photoresist was removed exposing the base Au sensor surface. This treatment allowed spatial confinement of the pipetted water sample solutions up to 200 L. Details of electrode structures and actual layouts of both devices including the electric connections for measurements are in Fig. 1. Each design contained four identical sensor pads on the chip surface arranged in the Wheatston bridge (see Fig. 1). Each of the sensors could be used for measurement with the remaining three serving for compensation of the temperature drift. For additional temperature stability the chip was placed on a computer processor heat sink using thermally conducting paste purchased in a local computer store. A stabilized laboratory DC power supply (Hewlett-Packard E3610A; Engelwood, CO, USA) was used to supply the Wheatstone bridge. Electric connection was provided by the 20-pin test clip from 3M (923690-16, St. Paul, MN, USA) and the resistance changes were monitored using a Clarity Lite A/D converter (C40, DataApex, Prague, Czech Republic) connected to a Windows PC. Data were monitored at 6 Hz. The measured signal was in mV and, if needed, the relative resistance change, ⌬ R, can be, for small changes, estimated according to the Eq. (1): ⌬ R = 2⌬ UR/U
(1)
Where R is the resistance of the sensor element, ⌬ U is the measured voltage change, and U the applied voltage.
3 Results and discussion The first experiments were directed toward characterizing the sensor response for different water-based solutions. Since all four sensors had practically identical resistance (790 ⍀) the measuring bridge was very well balanced and, without the sample, practically zero signal was recorded. Also the temperature drift was practically eliminated in the bridge arrangement. In the early stage of the work, we have also tested a device with a single sensing element; however, the temperature drift, even on a temperature controlled surface, made reliable measurements difficult. Since the gold sensor surface was in direct contact with the sample solution, we have tested the influence of different buffers and buffer concentrations. In this experiment, water was driven across the chip until stable baseline was reached, followed by 20 mM acetate buffer, pH = 5 for 25 min and water again. The change from low conductivity water (1 S/cm) to the highly conductive electrolyte (2 mS/cm) resulted in the signal of –2.02 mV. Practically no difference was observed when flushing the C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
Figure 1. (A) Layout of resistive elements with interdigitated comb counter electrode. Wide electrodes 250 m wide; narrow electrodes 50 m wide. Actual dimensions of the sensing element – 10 × 15 mm. (B) Detail of the low-through PDMS top including a serpentine microchannel and connectors for sample delivery. (C) Similar device for measurements in 50 L pipetted drops and electric scheme of the measurement.
system with 10 mM or 100 mM sodium acetate buffer or with 100 mM PBS solution. Small signal could be recorded when an air bubble entered the channel. An example of the sensor selectivity for two compounds differing just by the thiol group (beta-alanine does not contain –SH group) is in www.electrophoresis-journal.com
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Figure 2. Comparison of the signal response for thiolated and nonthiolated compounds measured in the flow through device with 1 mM Cysteine or beta alanine dissolved in 20 mM acetate buffer. At the start of the experiment the system was filled with 20 mM acetate buffer. The rise of the signal corresponds to the time when the sample solution reached the sensor element. Sample flow rate was 14 L/min delivered from a pressurized vial. In case of Beta-alanine there is actually resistance shift due to ionic adsorption on the order of 10−4 in comparison to cysteine, but after washing the sensor with water or acetate buffer, this deviation came back to the initial value. This effect was not observed at cysteine sample.
Figure 3. Graph of signal responses to various concentrations cysteine (left) and glutathione (right) in 20 mM acetate buffer. Different times of the signal rise correspond to different start of the sample flow. The negative onset in 1 M glutathione curve is caused by thermal shock from not well thermally equilibrated sample.
Fig. 2 where signals corresponding to 1 mM cysteine and beta alanine are plotted. To minimize side effects of electrochemical reactions or Joule heating the voltage source supplying the Wheatstone bridge was set to 0.2 V, resulting in 0.1 V voltage drop on each sensor element. Thus the dissipated heat was only 13 micro Watt per sensor. Figure 3 illustrates the change of the measured signal for series of solutions prepared in 20 mM acetate buffer at pH = 5. The signal was about two times higher in case of cysteine. This might correspond to the size of the molecule where the larger molecules of glutathione provide lower density of the SH-Au bonds due to the steric hindrance. The signal increase was fast in the first minute, slowing down later before reaching the saturation. This corresponds well with the previous observations using the quartz crystal microbalance measurements [21, 22]. Shortly, in the initial stage the rough, sparse monolayer is built very quickly, followed by C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
much slower reorganisation of the layer, forming higher packing density and uncovering new binding sites. This effect is less noticeable with decreasing concentration of GSH, probably because of the feed rate of new GSH molecules is going to be equal or lower than speed of the initial assembly. Since the molecules attached to the gold surface create a monomolecular layer the process should comply with the Langmuir isotherm. The plot in Fig. 4 shows the signal response at different concentrations of the analytes in the solution recorded five minutes after the exposure of the gold surface. The curves were fitted to the data using modified Langmuir isotherm equation y = A × (K × c)/(1 + K × c) , where c is the analyte concentration, A is a constant of proportionality between the degree of surface coverage and the resistivity change (response factor) and K is the equilibrium constant. Similar curves were obtained also at different times of the surface-analyte contact. While the fit is very good, it should be noted that the recorded signal is only a measured www.electrophoresis-journal.com
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Figure 4. Signals corresponding to the binding of glutathione (black squares) and cysteine (white squares) at different concentrations 5 min after the gold surface exposure to the respective solutions. The solid lines correspond to curve fittings according to the Langmuir isotherm.
potential difference and does not directly correspond to the absolute surface coverage of the gold layer. While this work was focused mainly on the fundamental properties of the Au based thin layer sensor, we have also tested its reusability after electrochemical redox cleaning of the gold surface. In this experiment, the used chip surface was regenerated for 13 min in a flowing acetate buffer a by series of 10 voltage pulses of ±1.8 V applied between one end of the resistor and the interdigitated comb counter electrode. The polarity of the applied pulses was periodically switched and ended at –1.8 V on the sensor to prevent forming of an oxide layer on the active element. Pulse duration was 1 min with 20 s intervals in between. After this procedure the sensor resistance has returned to its original value and was ready for next measurement. All the curves shown in Fig. 3 were obtained on the same sensor element with the regeneration steps in between measurements. It is important to stress that the use of the same sensor was necessary for absolute signal comparisons shown in Fig. 3. This relates to variations in the thickness of the Au layer resulting in variations of the sensor response as disclosed by comparing the maximum signals recorded in Figs. 2 and 3 for the same concentration of cysteine. Of course much better batch to batch repeatability and uniformity of the Au layer thickness on the substrate could be obtained by using a semiconductor grade sputtering machine instead of the inexpensive one for electron microscope specimen preparation available for this work. All the above experiments were performed using the flow through device shown in the upper part of the Fig. 1. The second tested device was designed for measurements in stationary droplets. While we could monitor the signals as in the previous device, the water evaporation during the extended measuring periods resulted in a significant signal drift. We have attempted to minimize the solvent evaporation effects by enclosing the device in a plastic box filled with C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
water saturated air. Unfortunately, the manipulation of small sample volumes (50–150 L) proved to be difficult and the evaporation still caused both temperature drift and concentration changes. On the other hand, the open structure provides a good access to the Au surface. Thus we are currently testing this device for selective sample preconcentration of thiol containing compounds from complex samples.
4 Concluding remarks The experiments demonstrate that a thiol-specific sensor, capable to monitor concentration range of six orders of magnitude, can be based on the simple thin Au layer resistance measurements. Response of the sensor is fast in the first stage of coverage followed by slower resistivity rise in the saturation phase. After each measurement the sensor can be easily regenerated by series of voltage pulses. Such a device allows performing, robust, and simple analyses and preconcentration of thiol containing species from diluted samples and releasing them for further analyses. During the experiments the sensor exhibited resistance to interference from higher concentrations of nonthiolated species in the sample. While desorption of the molecules from the sensor surface could not be measured using the DC powered Wheatstone bridge, it could be achieved by monitoring the resistance in the alternating current mode. This will be included in the next generation design. Given the cumulative nature of the sensor’s response we did not evaluate the LOD; however, for a 14 L/min flow rate and 30 min measurement time we can estimate the LOD to be at 10−10 M for cysteine. The authors wish to thank Dr. Miloˇs Kl´ıma from the Masaryk University in Brno, for the plasma pen surface treatments. This work was supported by GACR P20612G014 and institutional funding RVO: 68081715. Part of the work was
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realized in CEITEC - Central European Institute of Technology with research infrastructure supported by the project Czech Republic.1.05/1.1.00/02.0068 financed from European Regional Development Fund. The authors have declared no conflict of interest.
5 References [1] Homola, J., Vaisocherova, H., Dostalek, J., Piliarik, M., Methods 2005, 37, 26–36. [2] Homola, J., Yee, S. S., Myszka, D., Optical Biosensors: Today and Tomorrow, 2nd edition, Elsevier Science, 2008, pp. 185–242. [3] Wang, F., Gu, H. W., Swager, T. M., J. Am. Chem. Soc. 2008, 130, 5392–5393. [4] Fried, G. A., Zhang, Y. M., Bohn, P. W., Thin Solid Films 2001, 401, 171–178. [5] Steinecker, W. H., Rowe, M. P., Zellers, E. T., Anal. Chem. 2007, 79, 4977–4986.
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[10] Turkevich, J., Stevenson, P. C., Hillier, J., Disc.Faraday Soc. 1951, 11, 55–75. [11] Juskova, P., Foret, F., J. Sep. Sci. 2011, 34, 2779– 2789. [12] Juskova, P., Neuzil, P., Manz, A., Foret, F. Lab Chip 2013, 13, 781–784. [13] Nuzzo, R. G., Fusco, F. A., Allara, D. L., J. Am. Chem. Soc. 1987, 109, 2358–2368 ¨ [14] Gronbeck, H., Curioni, A., Andreoni, W., J. Am. Chem. Soc. 2000, 122, 3839–3842. [15] Flynn, N. T., Tran, T. N. T., Cima, M. J., Langer, R., Langmuir 2003, 19, 10909–10915. [16] Zhang, Y., Terrill, R. H., Bohn, P. W., J. Am. Chem. Soc. 1998, 120, 9969–9970. [17] Ossendorfova, N., Pradac, J., Koryta, J., J. Electroanal.Chem. Interfacial Electrochem. 1970, 28, 311–316. ˇ J., Kl´ıma, M., Slav´ıcek, ˇ ˇ ´ L., Surf. Coat. [18] Janca, P., Zaj´ıckov a, Technol. 1999, 116, 547–551. [19] Main, C., Lenihan, J. M. A., J. Phys. E 1978, 11, 1123.
[6] Mansurov, G. e. a., Dokl. Akad. Nauk. SSSR 1977, 236, 153.
[20] Herold, K. E., Rasooly, A., Fabrication and Microfluidics, Volume 1, Caister Academic Press, Norfolk, UK 2009.
[7] Muller, L., Mansurov, G. N., Petrii, O. A., J. Electroanal. Chem. 1979, 96, 159–164.
[21] Hepel, M., Tewksbury, E., J. Electroanal. Chem. 2003, 552, 291–305.
[8] Cai, Q. Y., Zellers, E. T., Anal. Chem. 2002, 74, 3533–3539.
[22] Bieri, M., Burgi, T., Phys. Chem. Chem. Phys. 2006, 8, 513–520.
[9] Podesva, P., Foret, F., Curr. Anal. Chem. 2013, 9, 642–652.
C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
www.electrophoresis-journal.com
Electrophoresis 2015, 00, 1–6
ˇ Pavel Podesva ˇ Frantisek Foret Institute of Analytical Chemistry AS CR, v.v.i., Brno, Czech Republic
Received April 13, 2015 Revised May 13, 2015 Accepted May 13, 2015
Research Article
Metal nano-film resistivity chemical sensor In this work, we present a study on reusable thin metal film resistivity-based sensor for direct measurement of binding of thiol containing molecules in liquid samples. While in bulk conductors the DC current is not influenced by the surface events to a measureable degree in a thin metal layer the electrons close to the surface conduct a significant part of electricity and are influenced by the surface interactions. In this study, the thickness of the gold layer was kept below 100 nm resulting in easily measureable resistivity changes of the metal element upon a surface SH-groups binding. No further surface modifications were necessary. Thin film gold layers deposited on a glass substrate by vacuum sputtering were photolithographically structured into four sensing elements arranged in a Wheatstone bridge to compensate for resistance fluctuations due to the temperature changes. Concentrations as low 100 pM provided measureable signals. The surface after the measurement could be electrolytically regenerated for next measurements. Keywords: Adsorption / Chemiresistor / Nano-film / Thiol binding / Thiol sensing DOI 10.1002/elps.201500190
1 Introduction The main use of thin metal layers in bioanalysis relates to the measurements based on the surface plasmon resonance (SPR). Typically a well conducting thin metal layer (gold, silver) is deposited on a glass surface and illuminated by a light source from the glass side. The reflected light intensity is monitored. The surface plasmon, generated at certain angle, responds to the surface refractive index changes allowing detection of the binding events. The evanescence field penetrating a fraction of the wavelength above the metal surface allows monitoring of not only the primary binding on the metal surface, but provides information about binding events within the evanescence field distance. Thus it is possible to monitor not only the primary binding but also the affinity interactions on the chemically modified surface. Surface plasmon resonance is one of the best developed optical techniques for affinity-binding studies, especially for biomolecular research [1,2] and commercial instrumentation is available from a number of manufacturers. On the other hand, the SPR instrument is relatively complex and provides only an indirect optical information about the binding events with little information about the first chemical layer binding. Monitoring of basic electric properties (current, voltage) represents one of the most precise measurement techniques. Besides the simplicity and precision, electrical measurements also provide a large dynamic range. The related equipment is
ˇ Correspondence: Dr. Pavel Podesva, Institute of Analytical Chemistry AS CR, v.v.i., Veveri 97, 902 00 Brno, Czech Republic E-mail: [email protected]
Abbreviations: GSH, glutathione; HDMS, hexamethyldisilasane; SPR, surface plasmon resonance C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
readily available and inexpensive, resulting in the widespread use not only in electronics measurements but also in a variety of sensors used, e.g. in physics, chemistry, environmental sciences, or medicine. While many physico–chemical interactions can be readily converted into electric signals using a suitable sensor (e.g. in optical sensing), other interactions result in a direct electric change upon the interaction with the sensor material. For example, chemiresistors, a family of electronic sensing components, change their electric resistance depending on the physic-chemical processes occurring at its active surface, including charge induced field effect [3], formation or dissolution of self-assembling monolayers [4], swelling of a conductive layer [5], molecular or ionic adsorption-desorption, redox changes [6, 7], etc. Typical chemiresistor is an inexpensive, sensitive, small, robust, and low power device commonly used for detection in gas or liquid phases [8, 9]. The metal layer itself is generally not considered as a suitable sensing element for chemiresistors since its ability to conduct DC electric current is given by all electrons available in the bulk metal and the contribution of any surface events is negligible. However, as the thickness of the metal layer decreases, the contribution of the electrons close to the surface to the total conductivity of the material increases and their mobility and scattering depends on the surface interactions. While this effect has been theoretically described over 60 years ago [10], only few papers have investigated the effect experimentally [11] and to our knowledge there is no work describing a practical application of a simple thin metal sensing probe except to our previous report on the free floating nanoelectrodes for electrochemiluminiscence generation [12]. The strength of the Au–S bond is about 50% of the C–C bond (170–210 kJ/mol) [13, 14] and, once formed, the release
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of the attached compounds in water or acetate buffer environment (pH = 5–7) is slow, on the order of days [15]. The full coverage of the 40 nm gold film sensor by the monolayer of the thiol compound leads to the increase of resistance on the order of 4% largely independent on the type of organic residue connected to the SH group [16, 17]. This resistivity change of the layer is; however, influenced by the number of sulphur atoms connected to the surface. In this work, we have investigated the DC electric resistance of thin (45–60 nm) Au films upon binding of organic thiol containing compounds in water solutions. The Au sensing elements were formed either as a part of a flow-through microfluidic device or as a droplet sensing area. By real-time resistance monitoring of the Au element is was possible to observe the rate of coverage during the experiment. By applying the electric potential between the working and counter electrode it was also possible to release the attached thiol containing species and reuse the Au element for further measurements.
2 Materials and methods 2.1 Reagents and equipment All reagents were of analytical grade quality. (–)-Glutathione, oxidized 98% (HPLC) (Sigma) (GSSG); (–)-Glutathione, reduced 98.0% (Sigma-Aldrich) (GSH); Gly-Pro-Glu 98% (HPLC) (CAS 32302-76-4) (GPG); Hexamethyldisilasane reagent grade, 99% (HDMS) were purchased from SigmaAldrich inc. Sylgard 184 silicone elastomer was purchased from Dow Corning GmbH (made in USA). Acetic acid 98% p.a. and Toluene p.a. from Lach-Ner, s.r.o., Czech Republic. Acetonitrile p.a. was purchased from Penta chemicals, Czech Republic. MicropositTM LOLTM 2000 Lift Off Layer (LOL 2000) and MicropositTM remover 1165 were purchased from Rohm and Haas Electronic Materials LLC. The ma-P 1225 photoresist and ma-P 1225 developer were purchased from Micro resist technology GmbH, Germany; Nanofilm glass – Borofloat 1.5 mm, 1000A˚ Cr layer, 5300A˚ AZ1518 Photoresist layer from Nanofilm Inc. USA; 3 inch 0.65 mm borofloat glass wafers from S. I. Howard Glass, Cambridge, MA, USA; Demineralized water (18 M⍀/cm) was used for all experiments. The Au layers were prepared by vacuum magnetron sputtering using Bal-Tec SCD 500 equipped with a QCM thickness monitor (Bal-Tec, Schalksm¨uhle, Germany). Clarity Lite (DataApex, Prague, Czech Republic) two channel A/D converter module and a laboratory constructed nanomicroampere adjustable constant current source was used for resistance monitoring. For improved hydrophilic properties the glass surfaces were modified using a radiofrequency plasma pen developed at the Masaryk University, Brno, Czech Republic [18].
2.2 Device fabrication The 3 inch borosilicate substrate was consecutively cleaned with detergent, water, piranha solution (H2 SO4 :H2 O2 /3:1), C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
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washed thoroughly with water, dried with nitrogen stream, baked at hotplate for 1 h at 170°C, then treated in 1% solution of HDMS (submersion) for 2 min and left to dry. The first spin-coated layer was the LOL 2000 spun at 3000 rpm for 30 s followed by baking for 5 min at 170°C on a hotplate. The second layer was Ma−P 1225 spin-coated for 30 s at 3000 rpm and baked at 100°C for 90 s. The wafer was exposed with PG 101 lithograph, developed in ma-P 1225 developer, washed with water and dried in a clean air stream. Next, a 3 nm indium-tin oxide adhesion layer and 45 nm of gold was sputtered on the wafer using the Bal-Tec SCD 500. Resists were then stripped with Microposit remover 1165. The resulting wafer with the Au motive was finally annealed in furnace for 240 min at 450°C in the air atmosphere to decrease porosity and reduce resistance by 50% [19]. Resistance of each resistor in the bridge was 790 ⍀ after annealing. Two electro-mechanical arrangements have been prepared. In both cases the gold surface sensor was arranged in a serpentine shape to provide desired resistance. The serpentine layout was complemented by a comb-like electrode interdigitated into the measuring electrode meanders. This electrode was tested as a cleaning electrode for electrochemical cleaning of the sensor after the measurement. The first design was arranged as a flow through device for on-line operation while the second device was prepared as an open structure allowing measurement in liquid drops pipetted on the electrode surface. The microfluidic part of the first design was casted from PDMS resin using a glass master microfabricated by photolithography on the Nanofilm glass. The resulting fluidic channels were 50 m deep. The form for casting of the PDMS part was milled in a PMMA sheet and the glass master was treated by HDMS for easy separation from the cured PDMS resin. Fused silica capillaries (100 m id × 360 m od, Polymicro Technologies, USA) were connected to the PDMS block with the LabSmith (LabSmith, Inc.; Livermore, CA, USA) ferrules type C360100 half-casted into PDMS block to ensure tight input/output without large void volumes. Completion of the chip was performed using plasma pencil [18] producing a stream of cool O3 containing Ar plasma for activation of the surface of the glass and PDMS to the Si-OH state. The time needed for sufficient treatment of both the PDMS and glass surfaces was about 10 s at which both surfaces were strongly hydrophilic with almost zero contact angle (checked visually with drop of water). In the next step, the PDMS part was placed on the glass sensor and cured for 5 h on a 100°C hotplate, leading to a permanent bond between the glass and PDMS. For easier manipulation with the activated PDMS block before pressing it onto the activated glass few drops of the methanol were applied on its surface [20]. This allowed precise adjustment of the PDMS block position before establishing the permanent contact. In the next step, ferrules with capillaries were inserted into port holes. For best mechanical endurance edges of the PDMS as well as the top sides of the ferrules were sealed by PDMS to form one compact piece. www.electrophoresis-journal.com
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The open surface device was fabricated on the same glass substrate as the previous one. Also the sensor circuit is the same with the only difference being the round shape of the resistive elements. The surface of the final device except the sensor pads was coated by hydrophobic layer of Teflon (PTFE Disp. 30, DuPont, Wilmington, DE, USA). During the coating the sensor pads were protected by a photoresist pattern and after hardening of the Teflon layer the photoresist was removed exposing the base Au sensor surface. This treatment allowed spatial confinement of the pipetted water sample solutions up to 200 L. Details of electrode structures and actual layouts of both devices including the electric connections for measurements are in Fig. 1. Each design contained four identical sensor pads on the chip surface arranged in the Wheatston bridge (see Fig. 1). Each of the sensors could be used for measurement with the remaining three serving for compensation of the temperature drift. For additional temperature stability the chip was placed on a computer processor heat sink using thermally conducting paste purchased in a local computer store. A stabilized laboratory DC power supply (Hewlett-Packard E3610A; Engelwood, CO, USA) was used to supply the Wheatstone bridge. Electric connection was provided by the 20-pin test clip from 3M (923690-16, St. Paul, MN, USA) and the resistance changes were monitored using a Clarity Lite A/D converter (C40, DataApex, Prague, Czech Republic) connected to a Windows PC. Data were monitored at 6 Hz. The measured signal was in mV and, if needed, the relative resistance change, ⌬ R, can be, for small changes, estimated according to the Eq. (1): ⌬ R = 2⌬ UR/U
(1)
Where R is the resistance of the sensor element, ⌬ U is the measured voltage change, and U the applied voltage.
3 Results and discussion The first experiments were directed toward characterizing the sensor response for different water-based solutions. Since all four sensors had practically identical resistance (790 ⍀) the measuring bridge was very well balanced and, without the sample, practically zero signal was recorded. Also the temperature drift was practically eliminated in the bridge arrangement. In the early stage of the work, we have also tested a device with a single sensing element; however, the temperature drift, even on a temperature controlled surface, made reliable measurements difficult. Since the gold sensor surface was in direct contact with the sample solution, we have tested the influence of different buffers and buffer concentrations. In this experiment, water was driven across the chip until stable baseline was reached, followed by 20 mM acetate buffer, pH = 5 for 25 min and water again. The change from low conductivity water (1 S/cm) to the highly conductive electrolyte (2 mS/cm) resulted in the signal of –2.02 mV. Practically no difference was observed when flushing the C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
Figure 1. (A) Layout of resistive elements with interdigitated comb counter electrode. Wide electrodes 250 m wide; narrow electrodes 50 m wide. Actual dimensions of the sensing element – 10 × 15 mm. (B) Detail of the low-through PDMS top including a serpentine microchannel and connectors for sample delivery. (C) Similar device for measurements in 50 L pipetted drops and electric scheme of the measurement.
system with 10 mM or 100 mM sodium acetate buffer or with 100 mM PBS solution. Small signal could be recorded when an air bubble entered the channel. An example of the sensor selectivity for two compounds differing just by the thiol group (beta-alanine does not contain –SH group) is in www.electrophoresis-journal.com
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Figure 2. Comparison of the signal response for thiolated and nonthiolated compounds measured in the flow through device with 1 mM Cysteine or beta alanine dissolved in 20 mM acetate buffer. At the start of the experiment the system was filled with 20 mM acetate buffer. The rise of the signal corresponds to the time when the sample solution reached the sensor element. Sample flow rate was 14 L/min delivered from a pressurized vial. In case of Beta-alanine there is actually resistance shift due to ionic adsorption on the order of 10−4 in comparison to cysteine, but after washing the sensor with water or acetate buffer, this deviation came back to the initial value. This effect was not observed at cysteine sample.
Figure 3. Graph of signal responses to various concentrations cysteine (left) and glutathione (right) in 20 mM acetate buffer. Different times of the signal rise correspond to different start of the sample flow. The negative onset in 1 M glutathione curve is caused by thermal shock from not well thermally equilibrated sample.
Fig. 2 where signals corresponding to 1 mM cysteine and beta alanine are plotted. To minimize side effects of electrochemical reactions or Joule heating the voltage source supplying the Wheatstone bridge was set to 0.2 V, resulting in 0.1 V voltage drop on each sensor element. Thus the dissipated heat was only 13 micro Watt per sensor. Figure 3 illustrates the change of the measured signal for series of solutions prepared in 20 mM acetate buffer at pH = 5. The signal was about two times higher in case of cysteine. This might correspond to the size of the molecule where the larger molecules of glutathione provide lower density of the SH-Au bonds due to the steric hindrance. The signal increase was fast in the first minute, slowing down later before reaching the saturation. This corresponds well with the previous observations using the quartz crystal microbalance measurements [21, 22]. Shortly, in the initial stage the rough, sparse monolayer is built very quickly, followed by C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
much slower reorganisation of the layer, forming higher packing density and uncovering new binding sites. This effect is less noticeable with decreasing concentration of GSH, probably because of the feed rate of new GSH molecules is going to be equal or lower than speed of the initial assembly. Since the molecules attached to the gold surface create a monomolecular layer the process should comply with the Langmuir isotherm. The plot in Fig. 4 shows the signal response at different concentrations of the analytes in the solution recorded five minutes after the exposure of the gold surface. The curves were fitted to the data using modified Langmuir isotherm equation y = A × (K × c)/(1 + K × c) , where c is the analyte concentration, A is a constant of proportionality between the degree of surface coverage and the resistivity change (response factor) and K is the equilibrium constant. Similar curves were obtained also at different times of the surface-analyte contact. While the fit is very good, it should be noted that the recorded signal is only a measured www.electrophoresis-journal.com
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Figure 4. Signals corresponding to the binding of glutathione (black squares) and cysteine (white squares) at different concentrations 5 min after the gold surface exposure to the respective solutions. The solid lines correspond to curve fittings according to the Langmuir isotherm.
potential difference and does not directly correspond to the absolute surface coverage of the gold layer. While this work was focused mainly on the fundamental properties of the Au based thin layer sensor, we have also tested its reusability after electrochemical redox cleaning of the gold surface. In this experiment, the used chip surface was regenerated for 13 min in a flowing acetate buffer a by series of 10 voltage pulses of ±1.8 V applied between one end of the resistor and the interdigitated comb counter electrode. The polarity of the applied pulses was periodically switched and ended at –1.8 V on the sensor to prevent forming of an oxide layer on the active element. Pulse duration was 1 min with 20 s intervals in between. After this procedure the sensor resistance has returned to its original value and was ready for next measurement. All the curves shown in Fig. 3 were obtained on the same sensor element with the regeneration steps in between measurements. It is important to stress that the use of the same sensor was necessary for absolute signal comparisons shown in Fig. 3. This relates to variations in the thickness of the Au layer resulting in variations of the sensor response as disclosed by comparing the maximum signals recorded in Figs. 2 and 3 for the same concentration of cysteine. Of course much better batch to batch repeatability and uniformity of the Au layer thickness on the substrate could be obtained by using a semiconductor grade sputtering machine instead of the inexpensive one for electron microscope specimen preparation available for this work. All the above experiments were performed using the flow through device shown in the upper part of the Fig. 1. The second tested device was designed for measurements in stationary droplets. While we could monitor the signals as in the previous device, the water evaporation during the extended measuring periods resulted in a significant signal drift. We have attempted to minimize the solvent evaporation effects by enclosing the device in a plastic box filled with C 2015 WILEY-VCH Verlag GmbH & Co. KGaA, Weinheim
water saturated air. Unfortunately, the manipulation of small sample volumes (50–150 L) proved to be difficult and the evaporation still caused both temperature drift and concentration changes. On the other hand, the open structure provides a good access to the Au surface. Thus we are currently testing this device for selective sample preconcentration of thiol containing compounds from complex samples.
4 Concluding remarks The experiments demonstrate that a thiol-specific sensor, capable to monitor concentration range of six orders of magnitude, can be based on the simple thin Au layer resistance measurements. Response of the sensor is fast in the first stage of coverage followed by slower resistivity rise in the saturation phase. After each measurement the sensor can be easily regenerated by series of voltage pulses. Such a device allows performing, robust, and simple analyses and preconcentration of thiol containing species from diluted samples and releasing them for further analyses. During the experiments the sensor exhibited resistance to interference from higher concentrations of nonthiolated species in the sample. While desorption of the molecules from the sensor surface could not be measured using the DC powered Wheatstone bridge, it could be achieved by monitoring the resistance in the alternating current mode. This will be included in the next generation design. Given the cumulative nature of the sensor’s response we did not evaluate the LOD; however, for a 14 L/min flow rate and 30 min measurement time we can estimate the LOD to be at 10−10 M for cysteine. The authors wish to thank Dr. Miloˇs Kl´ıma from the Masaryk University in Brno, for the plasma pen surface treatments. This work was supported by GACR P20612G014 and institutional funding RVO: 68081715. Part of the work was
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realized in CEITEC - Central European Institute of Technology with research infrastructure supported by the project Czech Republic.1.05/1.1.00/02.0068 financed from European Regional Development Fund. The authors have declared no conflict of interest.
5 References [1] Homola, J., Vaisocherova, H., Dostalek, J., Piliarik, M., Methods 2005, 37, 26–36. [2] Homola, J., Yee, S. S., Myszka, D., Optical Biosensors: Today and Tomorrow, 2nd edition, Elsevier Science, 2008, pp. 185–242. [3] Wang, F., Gu, H. W., Swager, T. M., J. Am. Chem. Soc. 2008, 130, 5392–5393. [4] Fried, G. A., Zhang, Y. M., Bohn, P. W., Thin Solid Films 2001, 401, 171–178. [5] Steinecker, W. H., Rowe, M. P., Zellers, E. T., Anal. Chem. 2007, 79, 4977–4986.
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[10] Turkevich, J., Stevenson, P. C., Hillier, J., Disc.Faraday Soc. 1951, 11, 55–75. [11] Juskova, P., Foret, F., J. Sep. Sci. 2011, 34, 2779– 2789. [12] Juskova, P., Neuzil, P., Manz, A., Foret, F. Lab Chip 2013, 13, 781–784. [13] Nuzzo, R. G., Fusco, F. A., Allara, D. L., J. Am. Chem. Soc. 1987, 109, 2358–2368 ¨ [14] Gronbeck, H., Curioni, A., Andreoni, W., J. Am. Chem. Soc. 2000, 122, 3839–3842. [15] Flynn, N. T., Tran, T. N. T., Cima, M. J., Langer, R., Langmuir 2003, 19, 10909–10915. [16] Zhang, Y., Terrill, R. H., Bohn, P. W., J. Am. Chem. Soc. 1998, 120, 9969–9970. [17] Ossendorfova, N., Pradac, J., Koryta, J., J. Electroanal.Chem. Interfacial Electrochem. 1970, 28, 311–316. ˇ J., Kl´ıma, M., Slav´ıcek, ˇ ˇ ´ L., Surf. Coat. [18] Janca, P., Zaj´ıckov a, Technol. 1999, 116, 547–551. [19] Main, C., Lenihan, J. M. A., J. Phys. E 1978, 11, 1123.
[6] Mansurov, G. e. a., Dokl. Akad. Nauk. SSSR 1977, 236, 153.
[20] Herold, K. E., Rasooly, A., Fabrication and Microfluidics, Volume 1, Caister Academic Press, Norfolk, UK 2009.
[7] Muller, L., Mansurov, G. N., Petrii, O. A., J. Electroanal. Chem. 1979, 96, 159–164.
[21] Hepel, M., Tewksbury, E., J. Electroanal. Chem. 2003, 552, 291–305.
[8] Cai, Q. Y., Zellers, E. T., Anal. Chem. 2002, 74, 3533–3539.
[22] Bieri, M., Burgi, T., Phys. Chem. Chem. Phys. 2006, 8, 513–520.
[9] Podesva, P., Foret, F., Curr. Anal. Chem. 2013, 9, 642–652.
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