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Resonance-Mode Electrochemical Impedance Measurements of Silicon Dioxide Supported Lipid Bilayer Formation and Ion Channel Mediated Charge Transport Anders Lundgren,† Julia Hedlund,‡ Olof Andersson,‡ Magnus Br€anden,‡,§ Angelika Kunze,§ Hans Elwing,† and Fredrik H€o€ok*,§ †
Department of Cell and Molecular Biology, University of Gothenburg, SE-405 30 Gothenburg, Sweden Stena Center 1B, Layerlab AB, SE-41292 Gothenburg, Sweden § Department of Applied Physics, Chalmers University of Technology, SE-412 96 Gothenburg, Sweden ‡
bS Supporting Information ABSTRACT: A single-chip electrochemical method based on impedance measurements in resonance mode has been employed to study lipid monolayer and bilayer formation on hydrophobic alkanethiolate and SiO2 substrates, respectively. The processes were monitored by temporally resolving changes in interfacial capacitance and resistance, revealing information about the rate of formation, coverage, and defect density (quality) of the layers at saturation. The resonance-based impedance measurements were shown to reveal significant differences in the layer formation process of bilayers made from (i) positively charged lipid 1-palmitoyl-2-oleoyl-sn-glycero-3ethylphosphocholine (POEPC), (ii) neutral lipid 1-palmitoyl-2-oleoyl-sn-glycero-3-phosphocholine (POPC) on SiO2, and (iii) monolayers made from POEPC on hydrophobic alkanethiolate substrates. The observed responses were represented with an equivalent circuit, suggesting that the differences primarily originate from the presence of a conductive aqueous layer between the lipid bilayers and the SiO2. In addition, by adding the ion channel gramicidin D to bilayers supported on SiO2, channel-mediated charge transport could be measured with high sensitivity (resolution around 1 pA).
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urface-supported and surface-tethered lipid bilayers have gained increased interest as they mimic the natural cell membrane and can host membrane-embedded proteins and peptides that typically lose their activity outside the lipid environment. Since surface-supported lipid bilayers are addressable using surface analytical techniques they also contribute importantly to the development of biosensor platforms.13 For example, interaction of suspended molecules with supported bilayers and proteins residing in supported bilayers have been successfully demonstrated using diverse techniques such as surface plasmon resonance,46 optical waveguide light-mode spectroscopy,7 and gravimetric quartz crystal microbalance.4,8,9 However, these measurement techniques are limited by being sensitive to changes in bound mass and in some cases also to biomolecular structural changes. With studies of uncharged solutes being an exception,10 these methods are therefore not appropriate for functional studies of transmembrane channels, pumps, and pores that transport ions or other charged molecules. To detect charge displacement across a supported membrane, electrical measurement methods like voltage clamp or electrochemical impedance measurements11 are required. Besides direct information on ion transport through incorporated channels or pores, the electrical methods also allow the membrane potential r 2011 American Chemical Society
to be controlled. Furthermore, impedance measurements give a relatively simple read-out of the membrane quality before insertion of channels and can therefore better than other measurement methods monitor the buildup process of lipid bilayers.12 A large portion of the scientific work focused on electrochemical measurements of supported and tethered bilayers has been devoted to the development of protocols for electrode-surface modification.2 This includes refinement of bilayer transfer or assembly methods,11,1320 including means to insert membrane proteins.2123 More recently, methods to stabilize the membranes24 and to form membranes on micro- or nanoporous substrates have been introduced.1,2529 With these types of substrates, ion channel activity has been demonstrated, in some cases reaching down to single-channel recordings for different ion channels.2729 Compared to the efforts put on substrate design, less effort has been invested into the development of comprehensive, integrated instrumentation for impedance analysis. The most common evaluation method is electrochemical impedance spectroscopy (EIS),30 where the complex impedance Received: June 9, 2011 Accepted: August 30, 2011 Published: August 30, 2011 7800
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Figure 2. (A) Cartoon and profile view of the z-LAB dual-electrode chip with different surface functionalizations (not to scale) and its position in the microfluidic channel of the measurement system. (B) Schematic of the z-LAB measurement circuit showing the principal components only. The interdigitated electrode sensor (WE1 and WE2 comprising Zc) is connected in parallel with an inductive impedance ZL. A small, typically 40 mV ac, voltage is applied over the parallel circuit in series with bias resistances R, and the total impedance Z is evaluated by measuring the resulting voltage over the sensor. A dc bias voltage can be applied at the sensor by using the potentiostat and the Ag/AgCl reference (RE) and Pt counter electrodes. Figure 1. Bode plots (phase angle α, black squares, and impedance |Z|, red circles) of a SiO2-coated dual-electrode sensor without (A) and with (B) an inductive impedance, ZL, connected in parallel. The solid and dashed lines in panel A corresponds to the fitted values. The insets are circuit representations of the dual electrodes, where each electrode interface is represented by a capacitance, C, and a resistance, Rp. The resistance of the electrolyte between the electrodes is denoted by Rs.
of electrode-supported lipid membranes is measured at a wide range of frequencies (usually from millihertz to megahertz).18,25,26,28,31 Indeed, EIS is a very powerful method, since it provides information about several relaxation regimes in a single measurement. However, data acquisition is rather slow, which limits the time resolution to several minutes or more, and fitting procedures to quantify electrical parameters are often complicated. This makes EIS best suited for end point measurements, although occasionally, EIS has also been successfully combined with current recordings at a constant voltage to map ion channel activity.25,27,28 Alternatively, changes in impedance can be temporarily resolved by measuring at a single frequency during, for example, bilayer formation, insertion of channel complexes, or activation/deactivation of channels.17,25,31 Although the signal is prone to noise, especially in the low-frequency regime, measurements at a single frequency allow significantly faster data acquisition than obtained using conventional EIS, but limit the possibility to quantify electrical parameters. An efficient way to reduce noise without sacrificing rapid data acquisition is a resonance-mode measurement. Recently we presented a setup for time-resolved analysis of the capacitive change at the interface of a pair of microelectrodes induced by
biomolecular adsorption.32 By connecting a pair of electrodes positioned on the same glass chip in parallel with an external inductance, the electrode pair becomes part of a resonator with a resonance frequency essentially determined by the interface capacitance, as demonstrated in Figure 1. Shifts in the resonance frequency can then be recorded by mapping the resonance peak, providing a means to resolve capacitance changes of less than 0.2 pF with a time resolution of 0.25 Hz. In this work, we use an improved version of this single-chip impedance method to demonstrate real-time recordings of bilayer formation for three well-characterized model systems: (i) phospholipid monolayers formed by rupture of vesicles made from the positively charged lipid 1-palmitoyl-2-oleoyl-sn-glycero-3-ethylphosphocholine (POEPC) onto gold electrodes modified with long-chain alkanethiols, phospholipid bilayers formed by rupture of (ii) positively charged (POEPC) and (iii) neutral 1-palmitoyl-2-oleoyl-snglycero-3-phosphocholine (POPC) vesicles onto gold electrodes coated with films of SiO2. In addition to determining the surface capacitance during bilayer formation, also the amplitude of the resonance peak was measured. In this way, simultaneous determination of the total current through the resonator at zero-phase could be obtained and correlated with changes in capacitance. The measured changes in both resonance frequency (capacitance) and amplitude (current) were successfully represented with an equivalent circuit, yielding new information about the lipid film formation processes as well as lipid membrane defect density. The sensitivity of the system with respect to ion channel recording was investigated by insertion of the gramicidin ion channel into the SiO2-supported lipid bilayer. 7801
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’ EXPERIMENTAL SECTION Materials. All experiments were carried out using symmetric gold electrodes positioned on sensor chips made of glass as shown in Figure 2. Microelectrodes, with a total area of 0.002 cm2 per electrode and spaced by a 30 μm gap, were provided by Layerlab AB. Gold sensor chips used for alkanethiollipid bilayer experiments were cleaned in a 1:1:5 solution of hydrogen peroxide (25%), ammonia (30%) and Milli-Q water at 85 °C for 10 min, carefully rinsed with Milli-Q water, and then immersed in a 1 mM solution of octadecanethiol (ODT) in ethanol (99.5%) for at least 20 h. This way, the gold electrodes become modified with a self-assembled monolayer of ODT as illustrated in Figure 2A. The sensor chips used for lipid bilayer formation were modified with a thin layer of SiO2 on top of the sensor electrodes. The SiO2 covered the entire sensor area (except for the contact pads positioned outside the liquid cell) and had a nominal thickness of either 5 or 30 nm as illustrated in Figure 2A. In order to clean and hydrophilize the SiO2-modified sensor chips, they were treated in a UVozone chamber twice for 10 min with an intermediate and final Milli-Q water rinsing step. See the Supporting Information for further details regarding materials and methods used. Experimental Procedure. In all experiments Tris buffer 10 mM, pH 8.0 with 100 mM NaCl was used as running buffer. Sensor chips were mounted in the instrument, and flow speed was set to a value between 10 and 50 μL/min. When a stable baseline was reached, typically within a few minutes, 30 nm lipid vesicles (see the Supporting Information for preparation) diluted to 0.2, 0.5, or 1.0 mg/mL in running buffer were injected. Gramicidin D was diluted in running buffer containing 1% ethanol from an ethanol stock solution kept in the freezer. In order to test the functionality of incorporated gramicidin D, a Tris buffer 10 mM pH 8.0 containing 100 mM KCl was injected before and after injections of the NaCl-containing buffer with gramicidin D. Electrochemical Measurements. As detailed in the Supporting Information, the sensor chip modified with 30 nm SiO2 was characterized by EIS (Autolab PGSTAT20, Ecochemie, Utrecht, The Netherlands). A representative spectrum is presented in the Bode plot in Figure 1A, confirming the expected capacitive dominance over a wide frequency range. The model circuitry presented in the inset of Figure 1A was fitted to the data using the Autolab FRA 4.9.005 software, and a good fit was obtained for C = 196 nF/cm2 and Rs = 728 Ω. No finite value for Rp could be determined using this frequency regime. By connecting the sensor electrodes in parallel with an external inductance, the electrode pair becomes part of a resonator as demonstrated in the Bode plot in Figure 1B. The circuit resonates at a resonance frequency f0, which is manifested by the phase drop from 90° to 0° and a positive peak in the impedance. Under the present conditions (see the Supporting Information) a satisfactory value of the capacitance can be calculated from the position of the resonance peak using the relation f0 = 1/(2π(LC)1/2) which is valid for an ideal RCL resonator. Since the formation of an isolating film on top of the electrode interfaces will lower the interface capacitance, this is expected to generate a positive shift in the resonance frequency. In contrast, the amplitude of the resonance peak will primarily be sensitive to changes in the interface resistance. Hence, introduction of conducting pores in the film would lead to dampening of the peak. The resonanceenhanced surface impedance (RESI) measurements were acquired
Figure 3. Change in capacitance and peak amplitude upon introduction of vesicles. (A) POEPC vesicles on ODT SAM. (B) POEPC vesicles on 30 nm SiO2. (C) POPC vesicles on 5 nm SiO2. (D and E) Magnified view of panels B and C highlighting the first 8 min after introduction of vesicles. The indicated time, tlag, defines the time between start of injection and start of bilayer formation.
using a setup from Layerlab AB (z-LAB, Layerlab AB, Sweden; see the Supporting Information for further details). An overview of the measurement chamber and the measurement circuit is shown in Figure 2. The resonance peak is detected by voltage division between the chip impedance and the bias resistance R, measuring the potential drop V over the chip. By using a set of signals U with frequencies close to the resonance peak, the z-LAB 7802
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Analytical Chemistry peak position, which varies between 2 and 5 kHz depending on surface coating, can be determined to within 10 000 nF/cm2), direct ionic contact between the gold electrode and the electrolyte is considered negligible. A common feature observed for all capacitance curves in Figure 3 is a sharp decrease in capacitance a few minutes after introduction of the vesicle suspension, being attributed to the onset of lipid monolayer and lipid bilayer formation on Au and SiO2, respectively. The steepest capacitance decrease and fastest saturation were observed for POPC on 5 nm SiO2, for which the capacitance level typically saturated within 1015 min after
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injection. Also, the POEPC lipid layer formed on the ODT surface saturated within 15 min. However, the time scale for the POEPC bilayer assembly varied significantly, from 15 min up to a few hours. As discussed further below, this relatively large variation is attributed to a high sensitivity of the measurements to the defect density in the lipid bilayer, which is expected to depend strongly on the influence of substrate properties on the electrostatic attraction between positively charged lipids and negatively charged SiO2. To evaluate the quality of the mono- and bilayers, the capacitance of the lipid films can be estimated from the measured change in capacitance. This gives Clipids values of 2070 ( 40 nF/ cm2 for the POEPC monolayer and 1290 ( 220 and 1160 ( 300 nF/cm2 for the POEPC and POPC bilayers, respectively. The obtained values are close to what is usually measured with EIS, where the capacitance for a defect-free bilayer is reported to be within the range of 5001000 nF/cm2 (ref 2) with the capacitance of a lipid monolayer being roughly twice that of a bilayer. The lowest measured capacitance for POEPC and POPC bilayers in this study was 1070 and 990 nF/cm2, respectively. We therefore assume that 1.0 μF/cm2 can be used as a reference value for essentially defect-free bilayers in this system. An additional difference between POEPC and POEPC bilayer formation is observed during the first minutes of vesicle injection (Figure 3D). For POEPC, the rapid decrease in capacitance starts directly after the vesicles reach the sensor surface, which takes 3 min at a flow speed of 10 μL/min. For POPC, the response was instead delayed by 24 min. During this lag phase, there is only a moderate ( Rb+ > K+ > Na+ > Li+.43 Gramicidin D is frequently used as a model system for studying fundamental aspects of ion transfer,11,1618,25,28 which can be related to more complex integral ion channels that span cellular membranes and regulate the ion transfer in nerve signaling, muscle contraction, etc. Figure 5 shows changes in capacitance and total current versus time upon a 10 min injection of 100 ng/mL gramicidin D in Tris buffer containing 100 mM sodium chloride. During gramicidin D addition, there is a constant increase in the current reaching approximately 130 pA (ΔINa+) after rinsing. In contrast, there is only a small and transient change in capacitance, illustrating that gramicidin D binding at this low coverage does not induce a measurable response. The very small jump in the current visible at 16 min, i.e., a few minutes before the gramicidin incorporation, is an artifact caused by the injection valve as the gramicidin sample was loaded. To verify that the change in current was mediated by the gramicidin channels, we examined whether a different response was obtained by exchanging the sodium in the buffer for potassium, for which gramicidin D has about 3 times higher conductivity.43 The Tris buffer with 100 mM potassium chloride was also injected before the gramicidin insertion, thereby
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acquiring a baseline for this buffer. Compared to sodium, potassium induced a larger capacitance and higher current (Figure 5), which is attributed to its smaller hydrodynamic radius and thus higher capacity to penetrate small defects in the lipid bilayer and the underlying SiO2 (see above).44 We stress, however, that gramicidin insertion is not expected to induce defects in the bilayer at these low concentrations. This was observed, however, at orders of magnitude larger gramicidin concentrations, which resulted in substantial and irreversible changes in the capacitance. The second injection of the potassium-containing buffer after the gramicidin D incorporation yields a significantly higher current response, but with no detectable difference in the capacitance response. The current response for potassium injection after gramicidin D was estimated to 340 pA, giving the ratio (ΔIK+)/(ΔINa+) = 2.6. This value is within the expected range and equal to the value previously obtained for SiO2supported bilayers.18 Gramicidin D injection on POPC bilayers was similar to that on POEPC bilayers (not shown), and the lowest concentration tested in this study (1 ng/mL) provided a current response of approximately 20 pA using sodium buffer (not shown), corresponding to a signal-to-noise ratio (S/N) of 10. As control, gramicidin injection was also tested on ODT-supported lipid monolayers and on unmodified SiO2 electrodes, with no persistent effects observed.
’ CONCLUSIONS The formation of surface-supported lipid monolayers and lipid bilayers could be measured and characterized based on their respective capacitance employing a single-chip setup designed for resonance mode electrochemical impedance measurements. The formation of lipid bilayers on SiO2 gave rise to a transient decrease in the peak voltage during initial bilayer formation, which was followed by an increase in peak voltage during the late stages of bilayer formation. We suggest that this dip, which was observed neither upon lipid monolayer formation nor in the capacitance response, is due to the thin layer of conducting solvent in between the SiO2 and the bilayer. It was also demonstrated that, in contrast to the capacitance, the damping of the peak voltage, which is dominated by this conductive channel, is also very sensitive to the amount of defects in the bilayer. This feature was utilized by translating the resonance peak into changes in total current upon insertion of gramicidin D ion channels into preformed lipid bilayers. These results showed a high S/N ratio even at low (1 ng/mL) concentrations of gramicidin D despite the presence of defects in the supported lipid bilayers and even with relatively thick SiO2 (30 nm) substrates. These features of the system, combined with the fact that the single-chip EIS system can be directly combined with complementary surface-sensitive methods, make us believe that the principle will contribute to fill a gap in the existing repertoire of surface analytical tools. ’ ASSOCIATED CONTENT
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Supporting Information. Additional information as noted in text. This material is available free of charge via the Internet at http://pubs.acs.org.
’ ACKNOWLEDGMENT The authors gratefully acknowledge Dr. Fredrik Bj€ orefors at Link€oping University for assistance with the EIS experiments. Patrik Wallin is gratefully acknowledged for his assistance with 7805
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