WO2006006073A1

WO2006006073A1 - Single channel sensor

Info

Publication number: WO2006006073A1
Application number: PCT/IB2005/002155
Authority: WO
Inventors: Bruce Cornell
Original assignee: Ambri Limited
Priority date: 2004-07-05
Filing date: 2005-07-05
Publication date: 2006-01-19

Abstract

The present invention provides a method of detecting binding events between receptor molecules and analyte molecules. This method includes obtaining a membrane-based biosensor, contacting the receptor molecules with an analyte sample containing analyte molecules, measuring the change in current through the membrane over a course of time, and resolving the current measurement into an integer number of concurrent conducting channels. The present invention also provides a method of detecting the presence of analyte molecules of interest in a sample. The present invention further provides a biosensor that is capable of detecting discrete binding events between receptor molecules and analyte molecules, and/or detecting the presence of analyte molecules of interest in a sample.

Description

SINGLE CHANNEL SENSOR

BACKGROUND OF THE INVENTION

1. Field of the Invention

[0001] The present invention relates to methods of detecting binding events between receptor and analyte molecules. The present invention also relates to methods of detecting the presence of analyte molecules of interest. The present invention further relates to a biosensor for the measurement of single ion channel currents across high-impedance lipid bilayer membranes tethered to metal electrodes.

2. Description of the Related Art

[0002] The cornerstone of modern electrophysiology is the patch-clamp technique developed by Mueller (Ann. N. Y. Acad. Sci. 1975, 274: 247-264) and Nehler and Sakmann (Nature 1975, 260: 799-802) due to its exquisite sensitivity in measuring single ion channel currents across lipid-based membranes. Andersen et al. (Biophys. J. 1983, 41: 119-133, 135-146 and 147-165) subsequently adapted this technique for use in lipid bilayers at potentials of up to 500 mV. More recently, biosensors based on ion channels or ionophores contained within lipid membranes that are deposited onto metal electrodes and where the ion channels are switched in the presence of analyte molecules have been described in patent applications AU89/00352, WO90/08783, WO 92/17788, WO 93/21528, WO 94/07593 and U.S. Pat. Nos. 5,204,239 and 6,432,629, the disclosures of which are incorporated herein by reference. As is disclosed in these applications, ionophores such as gramicidin ion-channels can be co- dispersed with amphiphilic molecules, thereby forming lipid membranes with altered properties in relation to the permeability of ions. Various methods of gating the ion-channels, use of chemisorbed arrays of amphiphilic molecules attached to an electrode surface, and means of producing lipid membranes incorporating ionophores on the chemisorbed amphiphilic molecules are also disclosed in the prior art.

[0003] It will be appreciated by one of ordinary skill in the art that the ability to measure single ion channel currents would confer significant advantages in terms of improved efficiency, sensitivity and robustness of the data so obtained. The essential advantage of a single ion channel sensors lies in the ability to resolve signal from noise or background i current, and hence to increase confidence and sensitivity in the detection of analytes by receptor molecules, especially at very low concentrations of analyte where binding of single or small multiples of analyte molecules can be required. Single channel detection refers not to the measurement of only one channel, but rather the ability to resolve a channel current into a discrete or integer number of open channels, hi practice, this means being able to identify "step changes" in signal currents, so as to discern the opening and closing of individual ion channels, one channel at a time.

[0004] The ability to resolve the membrane channel currents into current events associated with the formation and disruption of individual single channels is synonymous with "digital" sensing. Digital audio recording, as practiced with Compact Disc technology provides a major enhancement in signal-to-noise ratio by the digitization of analogue audio signals into binary pulses, and thereby affords a major advance in sensitivity over analogue recordings. Analogously, digital biosensing affords a major advance in signal-to-noise and sensitivity over analogue biosensing, by virtue of the digitization of bioelectronic signals.

[0005] Detection of single ion channel currents forms the basis of so called "patch clamp" methods for the measurement of ion channel currents in electrophysiology ("Single Channel Recording" 2^nd ED, 1995, Sakmann & Nehler; Plenum US). Bayley and Kasianowicz (Biophys J., 1999, Vol. 76: 837-845) have proposed using arrays of single channel sensors incorporating protein-engineered ionophores, and relying on the fact that analytes will vary in their interactions with such engineered ionophores, in terms of both binding affinity and binding lifetime. The proposal of these researchers is to use stochastic signal processing to generate characteristic sets of electronic parameters, or "molecular signatures" of different analytes, based on their differential binding pattern or distribution of binding lifetimes. In Braha et al. (Chem & Biol, 1997, VoI 4: 497-505), it is claimed that engineered pores have an advantage over the use of specific attached ligands, as the basis for simultaneous detection of multiple analytes in a biosensor: "strictly selective binding is not required because the single channel recordings are rich in information; and for a particular analyte the dissociation rate constant, the extent of channel block, and the voltage dependence of these parameters are distinguishing, while the frequency of partial channel block reflects the analyte concentration. A single sensor element might, therefore, be used to quantitate more than one analyte at once." However, the use of attached, highly specific receptors, such as antibodies, as ion channel ligands, is well established experimentally. The specificity and sensitivity of antibodies as receptors is unparalleled, and the proposal that engineered ion channels could demonstrate similar specificity or sensitivity by virtue of the electronic signatures associated with channel blocking and unblocking is insufficiently proven.

[0006] Additionally, Vogel et al. (Langmuir 2003, 19: 5567-5569) have reported generating gold electrodes containing chemisorbed lipid bilayers with high electrical resistance, which were capable of measuring single ion channel currents (Figure 1). The technique, however, was limited in its utility as a detection system for application in molecular diagnostics to the specific ion channel-analyte system used, since no means was suggested for how to use such a system for specific recognition and detection of analyte molecules.

[0007] Single channel recordings of [VaI¹] Gramicidin A in a lipid membrane with a resistance in excess of 100 GΩ was also reported by Goulian et al. (Biophys J., 1998) (see Figure 2). Goulian et al. also demonstrated that the dimer formation rate increased approximately five-fold as the tension increased from 0 to 4 dyn/cm (Figure 3). Tension acted to thin the membrane. Because the bilayer was effectively incompressible with respect to volume changes, an increase in applied tension resulted in thinning of the membrane, thereby increasing the dimer formation rate and the survival lifetime of the dimer gramicidin once formed. It is demonstrated that single ion channel currents can be used as a mechanical force sensor rather than a sensor to detect the presence and concentration of specific molecules of interest.

[0008] There is a need for an improved method of detecting discrete ion channel currents and of relating changes in the properties of such ion channel currents with binding events between receptor molecules associated with the ion channels and analyte molecules that form complexes with the receptor molecules, and/or an improved method of detecting the presence of analyte molecules of interest. Additionally, there is a need for biosensors that are capable of measuring single ion channel currents and detecting step changes in these single channel currents that are associated with binding events between receptor and analyte molecules with advantages of improved efficiency, sensitivity and robustness of the data obtained.

SUMMARY OF THE INVENTION

[0009] The present invention is directed to a method of detecting binding events between the receptor and analyte molecules. This method includes obtaining a membrane-based biosensor, contacting the receptor molecules of the membrane with an analyte sample containing analyte molecules, measuring the change in current through the membrane over a course of time, and resolving the current measurement into an integer number of concurrent conducting channels, hi some embodiments the measurements are processed to yield a measure of the analyte concentration and analyte species.

[0010] The present invention is also directed to a method of detecting the presence of analyte molecules of interest in a sample. This method includes obtaining a membrane-based biosensor, contacting receptor molecules of the membrane with a sample suspected of containing the analyte molecules of interest, and detecting the change in current through the membrane. In some embodiments the measurements are processed to yield a measure of the analyte concentration and analyte species.

[0011] hi the above methods, the membrane-based biosensor comprises a lipid bilayer membrane that includes a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs; receptor molecules attached to at least a portion of the second HMSMs; and a sensing electrode. The first HMSMs are provided in a first layer of the bilayer membrane and the second HMSMs are provided in a second layer of the bilayer membrane. The first layer of the bilayer membrane is tethered to the sensing electrode. The second HMSMs are capable of lateral movement within the membrane, whereas the first HMSMs are prevented from lateral movement in the membrane. When the first HMSMs and the second HMSMs are aligned, conducting channels are formed within the membrane, hi the present invention, 50 or fewer concurrent conducting channels are formed per sensing electrode.

[0012] Alternatively, the above methods can be modified by using an array of sensing electrodes, wherein the first layer of the bilayer membrane is tethered to the array of the sensing electrodes, hi some embodiments each of the sensing electrodes in the array may be electrically connected so as to permit the independent measurement of the current in each electrode in the array between the sensing electrode and a counter electrode. Either the sensing electrode or the counter electrode may be employed as the common electrode in these configurations.

[0013] The present invention further provides a biosensor that is capable of detecting discrete binding events between the receptor and analyte molecules, and/or detecting the presence of analyte molecules of interest. This biosensor comprises a lipid bilayer membrane and at least one sensing electrode, wherein the lipid bilayer membrane is tethered to the at least one sensing electrode.

[0014] hi this biosensor, the membrane includes a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs, and receptor molecules attached to at least a portion of the second HMSMs. The first HMSMs are provided in a first layer of the bilayer membrane, whereas the second HMSMs are provided in a second layer of the bilayer membrane. The second HMSMs are capable of lateral movement within the membrane, whereas the first HMSMs are prevented from lateral movement in the membrane. When the first HMSMs and the second HMSMs are aligned, conducting channels are formed within the membrane. In the present invention, 50 or fewer concurrent conducting channels are formed per sensing electrode. In some embodiments of the present invention, the interfacial area, the thickness of the lipid bilayer membrane and density of the first and second HMSMs are selected in such a way that step changes in current through the membrane due to formation of conducting channels can be detected.

[0015] The biosensor of the present invention also relates to a biosensor for measurement of single ion channel currents in electrode arrays possessing small dimensions relative to the diffusion distance of analytes within a sample solution and the interpretation signals detected in these arrays in terms of the analyte concentration and species.

[0016] The foregoing and other advantages of the present invention will be apparent to one of ordinary skill in the art, in view of the following detailed description of the preferred embodiment of the present invention, taken in conjunction with the accompanying drawings.

BRIEF DESCRIPTION OF THE DRAWINGS

[0017] Features of the present invention as well as a preferred mode of use, further objectives, and advantages thereof, will best be understood by reference to the following detailed description of an illustrative embodiment when read in conjunction with the accompanying drawings, wherein:

[0018] Figure 1 shows a trace of current pulses arising from the spontaneous formation and disruption of channels within a lipid bilayer membrane according to the prior art (Vogel et al. Langmuir 2003). [0019] Figure 2 shows single channel recordings of [Val^Gramicidin A in a lipid membrane with a resistance in excess of 100 GΩ according to the prior art (Goulian et al., Biophys J. 1998, 74: 328 - 337). The inset is an expanded view of events in which two channels were open simultaneously.

[0020] Figure 3 shows a plot of the logarithm of the formation rate of [Gly^Gramicidin A as a function of tension according to the prior art (Goulian et al., Biophys J. 1998, 74: 328 - 337).

[0021] Figure 4 shows molecular sensing achieved using the Tip Sensing Electrode (TSE) in accordance with an embodiment of the present invention. The tip diameter is approximately 1 micrometer in diameter and the membrane impedance is approximately 10⁹ to 10¹² ohms.

[0022] Figures 5a-5c shows detection of single channel currents using the TSE in accordance with an embodiment of the present invention.

[0023] Figure 6 shows silicon chip microelectrode comprised in a membrane-based biosensor in accordance with an embodiment of the present invention.

[0024] Figure 7 shows membrane impedance measured at frequencies from 2 to 1000 Hz, for a membrane area of approximately 0.1 cm² as a function of membrane lipid chain length in accordance with an embodiment of the present invention.

[0025] Figures 8a-8c show individually resolved ion channel dimers obtained from a bilayer lipid membrane in accordance with an embodiment of the present invention.

[0026] Figures 9A-9E shows a simulation of the effect of channel density on the ability of the sensor to resolve current pulses in accordance with an embodiment of the present invention.

[0027] Figure 10 shows a theoretical plot of the number of electrodes in an array of 1000 gated as a function of analyte concentration in accordance with an embodiment of the present invention.

DETAILED DESCRIPTION OF PREFERRED EMBODIMENTS

[0028] In order to provide a clear and consistent understanding of the specification and claims, including the scope given to such terms, the following definitions are provided: [0029] As used herein, the term "impedance" shall refer to complex proportionality constant between the voltage applied across a charge conducting element and the resultant charge movement through the element.

[0030] As used herein, the term "current" shall refer to a flux of ions, as measured in units of amperage or amps. Current can be either unidirectional ("direct current") or changing in direction with a given frequency ("alternating current"). In the present invention, typically alternating current is applied in combination with a direct current bias.

[0031] As used herein, the term "admittance" shall refer to complex proportionality constant between the current through a charge conducting element and the voltage applied across the element.

[0032] As used herein, the term "amphiphilic molecule" shall refer to a molecule having a hydrophilic head portion and one or more hydrophobic tails.

[0033] As used herein, the term "analyte molecule" shall refer to a target molecule found in a sample under testing.

[0034] As used herein, the term "receptor molecule" shall refer to a molecule that contains a recognition moiety that can bind with some specificity to a desired analyte molecule (i.e., target molecule). The binding relationship between the receptor and analyte molecules includes enzymes and substrates, antibodies and antigens, chelators and metal, cell surface receptors and receptor ligands, oligomer DNA probes and single and double-stranded DNA, aptimers and aptimer target sequences.

[0035] As used herein, the terms "ionophores" and "ion channels" shall interchangeably refer to natural or synthetic substances that promote the passage of ions through natural lipid membranes or artificial membranes. Ionophores can form ion-conducting pores in membranes.

[0036] As used herein, the term "discrete binding events" shall refer to binding events that can be resolved from earlier and later events permitting each event to be counted individually and that can be the basis of an assessment of the concentration and/or type of analyte molecule. [0037] As used herein, the terms "interfacial area" and "surface area" are interchangeable and shall refer to the area of contact between the first and second layers of the bilayer membranes.

[0038] As used herein, the terms "lateral movement" and "diffusion" are interchangeable and shall refer to the translational and two-dimensional movement of molecular species within the plane of a membrane under the action of Brownian motion.

[0039] As used herein, the term "density" shall refer to the number of half membrane spanning monomers (HMSMs) at the interfacial area of the lipid bilayer membrane.

[0040] As used herein, the terms "conductive surface" and "electrode" are interchangeable and shall refer to an electronically conducting solid surface, to which the membrane is tethered and in whose surface electronic currents flow to balance image charges from ions and other charge species in solution. A conductive surface is usually a metal, a semiconductor (such as silicon), a conductive ceramic, or a conductive polymer.

[0041] As used herein, the term "Tip Sensing Electrode (TSE)" shall refer to (a) a three dimensional surface with a needle point that is coated with or comprises a conductive material such as a metal (typically gold or palladium) that acts as both an electrical conductor and a binding surface to which the first layer components of a lipid bilayer membrane is tethered. The needle point can be in a range of sizes but is preferably of order 1 micron diameter and 10 microns long. In one embodiment, TSE refers to a support for an aqueous liquid being either a well or an absorbent pad on which the second layer components of a lipid bilayer membrane is dried. When an aqueous sample solution is added to the well or to the absorbent pad, the sample solution hydrates the second layer lipid components, transferring a fraction of those components to the gas-water interface of the hydrating aqueous sample solution. When the solution contacts the neighboring tip, the system comprising the needle point tip, the tethered first layer components, the hydrated second layer components and the hydrating aqueous sample solution spontaneously causes the establishment of a lipid bilayer membrane whose area defines the interfacial area.

[0042] As used herein, the term "conducting channel" shall refer to a pair of aligned HMSM ionophores which form a dimer. [0043] As used herein, the term "concurrent conducting channels" shall refer to a number of conducting channels which are aligned in the biosensor membrane at any given instant in time. In some embodiments the number is 50 or fewer, and in other embodiments the number is 10 or fewer.

[0044] As used herein, the term "channel lifetime" shall refer to the average time during which two HMSMs are aligned and remained together as an ionic dimer in a single conducting channel. Li some embodiments, channel lifetime is related to the channel length in relation to the membrane thickness and membrane fluidity.

[0045] As used herein, the term "gating" shall refer to the modulation of ion flux through ion channels. This modulation can be a turning off ("gating closed"), turning on ("gating open"), or more broadly, a change in pattern, rate, or width of current pulses. For dimeric ion channels, this occurs principally by means of breaking or dissociation of the monomeric half- channels. In transmembrane ion channels, this can occur by steric blocking or changes in conformation of the ion channel, due to interaction of other molecules with the channels. For both dimeric and transmembrane channels, "gating events" may be evident by means of resolving the distribution of channels lifetimes or repetition rates of individual ion channel current events. In the case of the ion channel switch, gating occurs when analytes bind to receptors associated with second HMSMs, resulting in a modulation of the ion flux facilitated by the channel, which permits the determination of the analyte species and/or concentration. The modulation is usually a reduction leading to gating the channel closed, but in competitive binding assays, the initial position of the channel is gated closed with a competitive analyte, so displacement of the competitive analyte results in an opening of the channel, also referred to as "gating the channel open".

[0046] As used herein, the term "temporal frequency distribution" shall refer to percentage of time during which any given number of conducting channels are open. This ranges from 1 conducting channel open for 1% of the time to 50 conducting channels open for 95% of the time. In some embodiments, 1 conducting channel opens for 40% of the time, 5 conducting channels open for 90% of the time, and 10 conducting channels open for 50% of the time. The temporal frequency distribution may be usefully characterized or represented by a statistical histogram, which uses a bar chart to indicate, for each integer number of concurrent conducting channels, the percentage of time that number of channels is open. [0047] As used herein, the term "event counting statistical analysis" shall refer to the method capable of detecting and resolving the individual current pulses arising from conductive dimers, wherein changes in formation and disruption of the conductive dimers arising from the binding of analyte molecules to receptor molecules provide a basis for measuring changes in the rate and conductive properties of the conductive dimers. Such changes can be determined by counting changes in the properties and the number of conductive channels formed within the sensor membrane, and the relation of such changes to the analyte concentration can be analyzed using statistical analyses as though the events were particles and interpreted as in related particle counting sensors such as Geiger counters and Radioimmunoassays .

[0048] hi the present invention, the phrase "inner half of the bilayer membrane" is used interchangeably with "the bottom half of the bilayer membrane", "first layer of the bilayer membrane" and "inner leaflet of the bilayer membrane".

[0049] hi the present invention, the phrase "outer half of the bilayer membrane" is used interchangeability with "the upper half of the bilayer membrane", "second layer of the bilayer membrane" and "outer leaflet of the bilayer membrane".

[0050] The present invention is directed to a method of detecting binding events between the receptor molecules and analyte molecules. This method includes obtaining a membrane- based biosensor, contacting the receptor molecules with an analyte sample containing analyte molecules, measuring the change in current through the membrane over a course of time, and resolving'the current measurement into an integer number of concurrent conducting channels.

[0051] The present invention is also directed to a method of detecting the presence of analyte molecules of interest in a sample. This method includes obtaining a membrane-based biosensor, contacting receptor molecules with a sample suspected of containing the analyte molecules of interest, and detecting the change in current through the membrane.

[0052] In the methods described herein, a membrane-based biosensor is first obtained. The biosensors suitable for this invention comprise a lipid bilayer membrane which includes a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs; receptor molecules attached to at least a portion of the second HMSMs; and a sensing electrode. The first HMSMs are provided in a first layer of the bilayer membrane and the second HMSMs are provided in a second layer of the bilayer membrane. The first layer of the bilayer membrane is tethered to the sensing electrode. The second HMSMs are capable of lateral movement within the membrane, whereas the first HMSMs are prevented from lateral movement in the membrane. When the first HMSMs and the second HMSMs are aligned, conducting channels are formed within the membrane. Preferably, 50 or fewer concurrent conducting channels are formed per sensing electrode. Figure 4 illustrates one embodiment of the invention.

[0053] Preferably, the density of the second HMSMs (i.e., number of HMSMs at the interfacial area) in the second layer of the bilayer membrane is less than the density of the first HMSMs (i.e., number of HMSMs at the interfacial area) in the first layer of the bilayer membrane. More preferably, the ratio of the HMSMs in the first layer of the bilayer membrane to the HMSMs in the second layer of the bilayer membrane is greater than 10, ensuring that the rate limiting conduction across the membrane is the second layer of the lipid bilayer membrane.

[0054] As a primary means of limiting the number of conducting channels formed within the membrane to 50 or fewer, it has been found that limiting the dimension of the interfacial area between the first and second layers of the bilayer membranes is particularly preferred. In general at a constant channel density, the smaller the interfacial area, the fewer channels are contributing to the conduction and the greater the ability to resolve individual channel conduction. Preferably, the interfacial area is between 1 and 10000 square microns, more preferably, less than 100 square microns, further preferably, less than 10 square microns, or still further preferably, less than 1 square micron. Membrane impedance is inversely proportional to membrane surface area, so the impedance of very small membranes are very high, for example on the order of 1 gigaohm or higher for 100 square micron membrane. Since achieving effective (e.g. non-leaking) electrical seals becomes more challenging for higher impedances, the main practical requirement for using small interfacial area membranes to resolve single channels is the need to achieve effective electrical seals.

[0055] A number of embodiments for achieving membrane electrical seals on the order of 1 gigaohm or higher include but are not limited to the following: a) elimination or suppression of leaks around the edge of the membrane interface by means of a hydrophobic solid surround or well edge, to which the hydrophobic tails of the membrane lipids adsorb strongly; and b) elimination or suppression of such leaks by means of creating an air-liquid- lipid interface around the edge of the membrane interface. Some of this means are illustrated in more detail in the examples incorporated herein.

[0056] In a preferred embodiment, the lipid bilayer membrane has a thickness in the range of 10 to 40 Angstroms, more preferably, in the range of 28 to 33 Angstroms.

[0057] As will be understood by one of ordinary skill in the art, the lifetime of a conducting channel formed by the alignment of first and second HMSMs depends not only on the type of HMSMs but also on the thickness of the membrane. Relative to the length of the ion channel and the fluidity of the membrane. As the membrane thickness increases, the conducting channel lifetime will decrease. Accordingly, when selecting particular HMSMs, account should be taken of the thickness of the membrane. The thickness of the membrane in general depends on the lipid carbon chain length, the hydrocarbon volume and the area per lipid molecule at the hydrocarbon-aqueous interface. It is preferred that thickness of the membrane is selected such that the channel lifetimes are less than 1000 seconds, more preferably, between 10 to 100 seconds, between 1 to 10 seconds, or between 50 milliseconds to 1 second. Furthermore, the fluidity of the membrane depends on chemical factors such as the composition of the membrane lipids, and physical factors such as temperature of the membrane, which controls the rate of two-dimensional diffusion of lipids within the plane of the membrane.

[0058] Ionophores suitable for the present invention include an ion channel drawn from the family of ion channels comprising β barrel structures, an ion channel drawn from the family of ion channels comprising non-ribosomally formed ion channels, and an ion carrier. In some embodiments, gramicidin, preferably gramicidin A is used. Li a preferred embodiment, the receptor molecules are biotinylated, and the second HMSMs comprise biotin modified gramicidin monomers. As a result, the biotinylated receptor molecules are attached to the at least a portion of the second HMSMs via the biotin-streptavidin-biotin link.

[0059] Electrodes suitable for the present invention are conductive surfaces composed of metals, which include palladium, gold, platinum, silver or mixtures of these metals, and mixtures of these metals with dopants and other impurities or matrix materials. The individual metals are chosen not only for their conductivity, but also for their ability to bind to sulfur, as part of the lipid sulfur-based tethering moieties in the first layer of lipids. A mixture of the individual metal with silver can be used to permit electrochemical ohmic conductance in the solid metal surface. Metal conductive surface has the advantages of permitting direct current detection of individual current conduction event.

[0060] In one embodiment, gold electrodes are used and the first layer of the lipid bilayer membrane is tethered to the gold electrode through sulfur-containing groups. Tethering a lipid bilayer membrane to an electrode such as a gold electrode provides enhanced stability for a biosensor. Unlike a conventional supported lipid bilayer, a tethered system allows the system to be formulated for an extended storage. The lipid tethering also allows high detection sensitivity due to an ionic reservoir region formed between an electrode and the tethered lipids. Ion flux between the reservoir and the external compartment allows convenient electrical transduction measurement in multi-sensor array format.

[0061] The dimensions of the sensing electrode are selected to generate the required surface area. The sensing electrode may be configured in many alternate geometries, including a flat planar pad, such as a rectangle or circle, or a three-dimensional protrusion, typically a cylindrical wire or tip. When configured as a cylindrical tip, the diameter will be between 1 and 100 microns, more preferably, about 10 microns and also about 1 micron. The length will be about 10 to 100 microns, more preferably about 10 microns in length. It will be appreciated by one of ordinary skill in the art that numerous electrode geometries and aspect ratios are possible which will meet the membrane area requirement, and can be selected on the basis of specific manufacturing requirements or end application.

[0062] The membrane-based biosensor that is suitable for the present invention is the ICS Biosensor disclosed by Australian Membrane and Biotechnology Research Institute (AMBRI) (where ICS is ion channel switch) (WO 98/55853, U.S. Pat. Nos. 5,234,566, 5,401,378, 5,637,201, 5,741,409, 5,753,093, 5,879,878, 6,291,155, 5,783,054, 6,343,346, 6,316,273, 6,417,009, 6,432,629, 6,447,656, 6,451,196, 6,573,109, the disclosures of which are incorporated herein by reference). The ICS biosensor disclosed by AMBRI provides features such as tethering chemistry and reservoir that are important for the present invention.

[0063] Next, a sample solution suspected of containing the analyte molecules of interest is introduced to be in contact with the receptor molecules. When the analyte molecules bind to the receptor molecules, the pattern of current pulses changes arising from the prevention of formation of conductive channel dimers. The pattern of current pulses corresponds to temporal frequency distribution of the conducting channels. The changes in the pattern of current pulses can be detected, measured, and further titrated against the target concentration. The current measurement is taken from the circuit formed between the sensing electrode and counter electrode, and further resolved into an integer number of concurrent conducting channels, thereby detecting the binding between the analyte and receptor molecules and further determining the concentration of the analyte molecules.

[0064] Samples that will include an analyte, and are suitable for the present invention include body samples and non-body samples. Examples of body samples are blood, serum, sweat, tears, urine, saliva, throat swabs, nasopharyngeal aspirates, smears, bile, gastrointestinal secretions, lymph, and organ aspirates and biopsies. Non-body samples include any solution samples not derived from a human body, for example, culture medium, ionic aqueous solutions, saline, organic acids and buffers. A wide variety of analytes such as hormones, proteins, nucleic acids, drugs, small molecules, microorganisms, electrolytes, antigens, and antibodies can be detected or quantitated by the present invention.

[0065] The binding of the receptor molecules to the analyte molecules that can be detected through the present method includes bindings between enzymes and substrates, antibodies and antigens, chelators and metal, cell surface receptors and receptor ligands, oligimer DNA probes and single- or double-stranded DNA, aptimers and aptimer target sequences.

[0066] hi one embodiment, the second HMSMs comprise biotin modified gramicidin monomers, and the receptor molecules are biotinylated and attached to the at least a portion of second HMSMs via the biotin-streptavidin-biotin link.

[0067] Alternatively, the above-described methods of detecting binding events between receptor and analyte molecules and of detecting the presence of analyte molecules of interest in a sample can be modified by using an array of sensing electrodes, wherein the first layer of the bilayer membrane is tethered to the array of the sensing electrodes.

[0068] As will readily be appreciated by one of ordinary skill in the art, the detection of individual channel currents in conjunction with the use of electrode arrays permits a novel and powerful class of target detection which is based on event counting statistics. The combination of event counting statistical analysis and sensor arrays in some embodiments offer enhanced speed, sensitivity, robustness and reliability arising from the redundancy of the detected events which may be used to improve confidence hi diagnosing the presence and concentration of a wide range of analyte molecules. The present invention has a potential optimized detection level of over 1000 times greater than that of current ELISA detection thresholds of 1 pM, and a ten-fold decrease in assay time.

[0069] The present invention is further directed to a biosensor that is capable of detecting discrete binding events between receptor and analyte molecules as well as detecting the presence of analyte molecules of interest in a sample. This biosensor comprises a lipid bilayer membrane and at least one sensing electrode, wherein the lipid bilayer membrane is tethered to the at least one sensing electrode.

[0070] In this biosensor, the membrane includes a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs, and receptor molecules attached to at least a portion of the second HMSMs. The first HMSMs are provided in a first layer of the bilayer membrane, whereas the second HMSMs are provided in a second layer of the bilayer membrane. The second HMSMs are capable of lateral movement within the membrane, whereas the first HMSMs are prevented from lateral movement in the membrane. When the first HMSMs and the second HMSMs are aligned, conducting channels are formed within the membrane. In the present invention, 50 or fewer concurrent conducting channels are formed per sensing electrode. Still in the present invention, the thickness of the lipid bilayer membrane and density of the first and second HMSMs are selected in such a way that a change in current through the membrane due to formation of conducting channels can be detected, hi a preferred embodiment, the lipid bilayer membrane has a thickness in the range of 10 to 40 Angstroms, more preferably, in the range of 28 to 33 Angstroms.

[0071] As will be understood by one of ordinary skill in the art, the lifetime of a conducting channel formed by the alignment of first and second HMSMs depends not only on the type of HMSMs but also on the thickness of the membrane. Relative to the length of the ion channel and the fluidity of the membrane. As the membrane thickness increases, the conducting channel lifetime will decrease. Accordingly, when selecting particular HMSMs, account should be taken of the thickness of the membrane. The thickness of the membrane in general depends on the lipid carbon chain length, the hydrocarbon volume and the area per lipid molecule at the hydrocarbon-aqueous interface. It is preferred that thickness of the membrane is selected such that the channel lifetimes are less than 1000 seconds, more preferably, between 10 to 100 seconds, between 1 to 10 seconds, or between 50 milliseconds to 1 second. Furthermore, the fluidity of the membrane depends on chemical factors such as the composition of the membrane lipids, and physical factors such as temperature of the membrane, which controls the rate of two-dimensional diffusion of lipids within the plane of the membrane.

[0072] In some embodiments, the density of the second HMSMs (i.e., number of HMSMs at the interfacial area) in the second layer of the bilayer membrane is less than the density of the first HMSMs (i.e., number of HMSMs at the interfacial area) in the first layer of the bilayer membrane. Preferably, the ratio of the HMSMs in the first layer of the bilayer membrane to the HMSMs in the second layer of the bilayer membrane is greater than 10, ensuring that the rate limiting conduction across the membrane is the second layer of the lipid bilayer membrane.

[0073] As a primary means of limiting the number of conducting channels formed within the membrane to 50 or fewer, it has been found that limiting the dimension of the interfacial area between the first and second layers of the bilayer membranes is particularly preferred. In general at a constant channel density, the smaller the interfacial area, the fewer channels are contributing to the conduction and the greater the ability to resolve individual channel conduction. Preferably, the interfacial area is between 1 and 10000 square microns, more preferably, less than 100 square microns, further preferably, less than 10 square microns, or still further preferably, less than 1 square micron. Still preferably, the interfacial area is bounded by hydrophobic solid or by gas. The gas can be inert gas or air.

[0074] In some embodiments, the interfacial area is established at the time of introduction of a sample solution to the sensing electrode. As a consequence of the first half membrane component being stored dry and tethered on the sensing electrode tip as well as the second half membrane component being stored dry at a site separate from the sensing electrode tip, the introduction of aqueous sample solution hydrates the second half membrane component, causing it to contact the first half membrane component thereby establishing the interfacial area. In some embodiments, the aqueous sample solution may contain the analyte molecules to be detected. In other embodiments, the analyte molecules may be introduced into the sample solution following the formation of the interfacial area.

[0075] Ionophores suitable for the present invention include an ion carrier. In some embodiments, gramicidin, preferably gramicidin A is used. In a preferred embodiment, the receptor molecules are biotinylated, and the second HMSMs comprise biotin modified gramicidin monomers. As a result, the biotinylated receptor molecules are attached to the at least a portion of the second HMSMs via the biotin-streptavidin-biotin link.

[0076] Electrodes suitable for the present invention are conductive surfaces composed of metals, which include palladium, gold, platinum, silver or mixtures of these metals, and mixtures of these metals with dopants and other impurities or matrix materials. The individual metals are chosen not only for their conductivity, but also for their ability to bind to sulfur, as part of the lipid sulfur-based tethering moieties in the first layer of lipids. A mixture of the individual metal with silver can be used to permit electrochemical ohmic conductance in the solid metal surface. Metal conductive surface has the advantages of permitting direct current detection of individual current conduction event. hi one embodiment, gold electrodes are used and the first layer of the lipid bilayer membrane is tethered to the gold electrode through sulfur-containing groups.

[0077] When a biosensor comprises a plurality of sensing electrodes, the electrodes are deposited on a single substrate thus allowing for an array of the electrodes to be formed. Raw data generated by the array of the electrodes are analyzed using event counting statistical analysis, for example, pattern recognition analysis, which enables data such as pulse time distributions to be interpreted in terms of particular target populations. Such analytical technique is configurable either to enable measurement of the analyte concentration at extremely low detection thresholds (high sensitivity), or to enable detection with less sensitivity but greater data redundancy, providing for an increase in assay confidence. Li a preferred embodiment, the array of the electrodes contains a range of receptor molecules with varying properties and target affinities, which enables rapid fingerprinting of unknown analyte molecules.

[0078] It will be appreciated by one of ordinary skill in the art that charge-coupled device (CCD) technology can be used to interrogate the biosensor described herein in a serial/parallel fashion, in order to image diagnostic characteristics of analyte samples.

[0079] The present invention provides detection of discrete binding events between the receptor and analyte molecules as well as detection of the presence of analyte molecules of interest. Unlike a continuous measurement of current as performed by conventional sensors, the resolution of individual current pulses provides additional information in the form of distributions of the properties such as the channel lifetime (i.e., pulse width), the formation rate (i.e., the time between the leading edges of the current pulses), and the number of concurrent conducting channels at any given time.

[0080] As the biosensor described in the present invention is gated closed, the current not only decreases but also the shape of the distribution changes, providing additional information on the nature of the receptor-analyte interaction. Advantages of this additional analysis include: 1) The status of the sensor can be determined prior to challenge with the test sample. Particularly, if the distribution of the metrics described above are not within proscribed ranges, the sensor can be rejected or be subject to a corrective interpretation to its reading; 2) The need for on board calibrators and controls is greatly reduced since the shape of the histogram provides an absolute indication of concentration rather than a relative measurement; and 3) The changes in the pattern of the distributions allow for more subtle, pattern recognition algorithms when interpreting the nature and even distributions of target molecules with a range of affinities and binding rates.

[0081] The following examples are given for the purpose of illustrating various embodiments of the invention and are not meant to limit the present invention in any fashion. It will be obvious to one of ordinary skill in the art that various changes can be made without departing from the scope of the invention, which is not to be considered limited to what is described in the specification.

EXAMPLES

Example 1. Detection of Single Channel Currents Using the Tip Sensing Electrode

(TSE)

[0082] A Tip Sensing Electrode (TSE) comprising a gold flashed 1 μm diameter cylindrical glass microelectrode and coated with a sulfur bound monolayer that constitutes the inner leaflet of a membrane was used (Figure 4). This leaflet was formed from membrane lipids, tethered ion channels and reservoir components. The membrane outer leaflet was presented as a monolayer firm on an aqueous liquid drop and comprised lipids and ion channels. Touching and partially withdrawing the TSE from the aqueous drop spontaneously formed a bilayer membrane at the surface of the aqueous solution.

[0083] Single channel currents obtained using the TSE circuit described above are illustrated in Figure 5. Prior to the measurement of the currents, sufficient silver (Ag) was deposited on the gold electrode through a period of 15-30 minutes of DC current flow (gold electrode negative) to permit steady current of 100 pA to flow in response to a potential of 200 mV (negative on the gold electrode).

[0084] Current response to a square wave excitation potential of ± 200 mV was applied between the gold electrode and an Ag/ AgCl return electrode in the sample solution. The positive excitation cycle, when the sensing electrode is held at a positive potential difference relative to the counter electrode, elicited a current transient indicating that the TSE behaved approximately as a capacitor in series with a resistive element, characterized by the current transient followed by decay to approximately zero current. The negative excitation cycle, when the sensing electrode is held at a negative potential difference relative to the counter electrode, elicited a current transient that decays at a similar rate to a steady current of approximately 100 pA (figure 5a). Current response on expanded scale is shown during positive (on gold electrode) potential cycle (Figure 5b) as well as during negative (on gold electrode) potential cycle (figure 5c). The later demonstrates current transients associated with the formation and disruption of individual conductive ion channel dimers.

Example 2. Detection of Bindings of Analyte Molecules to Receptor Molecules Using TSE

[0085] Analyte molecules (e.g. antibodies) in a sample solution are introduced to the membrane of Example 1. The binding of the analyte molecules to receptor molecules (e.g. haptens) on ion channels causes an alteration of the channel dimer lifetime, channel dimer conduction, or number of channels that can form dimers, and further changes in the pattern of current pulses arising from the formation of conductive channel dimers. Such changes are detected through the bilayer membrane formed between the tip of the cylindrical conductive surface and a gas-water interface of an aqueous solution. In order to minimize oxidation, the gas can be air, nitrogen, or a chemically inert gas such as Xenon. The inert gas can be used to prevent oxidation of the conductive surface. Example 3. Silicon Chip Microelectrode

[0086] Figure 6 shows an alternative example of a mechanism embodying the features of the membrane-based biosensor for the measurement of single channel currents and their modulation for the purpose of the detection of analytes when they interact with and bind to receptors (antibodies). It represents a single channel sensor based on a small diameter flat electrode within a hydrophobic well. Compared to the sensor using the TSE₅ this sensor is a lot smaller and has fewer concurrent channels.

[0087] This sensor comprises a mechanically rigid and molecularly smooth laminar substrate 1 of low conductivity material (e.g. high purity silicon or glass), a diffusion boundary layer 2 of typically titanium and/or tungsten of about 12.5 nm in thickness, a first conductive base layer 3 of gold of about 100 nm in thickness, an adhesion layer 4 of typically titanium of about 12.5 nm in thickness, a thick hydrophobic layer 5 of typically silicon nitride or silicon carbide of about 10000 nm in thickness, and a second conductive, hydrophilic layer 6 of typically titanium or platinum of about 500 nm in thickness.

[0088] In this biosensor, the diffusion boundary layer 2 acts to prevent the migration of subsequent metal layers into the substrate. The first conductive layer 3 contains gold that has high purity (0.99995 or higher atomic purity) and low roughness (1.5 nm or lower RMS roughness). The second conductive layer 6 acts as a counter electrode and a boundary for the outer second layer of the lipid bilayer membrane.

[0089] A well is defined within the interfacial area such that it extends from the upper surface of the gold layer 3 through the adhesion layer 4, the hydrophobic layer 5 and the second conductive layer 6 to provide an opening in the upper surface of the conductive layer 6. A membrane is located within the region of the well defined by the hydrophobic layer, which comprises a first layer and a second layer of closely packed amphiphilic molecules and a plurality of ionophores with at least a proportion of the amphiphilic molecules and ionophores of the lower first layer being connected to the upper surface of the first conductive layer 3 by means of linker groups which form an ionic reservoir between the membrane and the gold surface. [0090] The interfacial area is chosen to be in such a range that 50 or fewer conductive ion channel dimers are formed from the first layer ion channels complexing with the second layer ion channels. In some embodiments, the dimensions of the interfacial area are chosen so that approximately 1 to 10 conducting ion channel per second are formed. For the ion channel densities normally employed, approximate interfacial areas are preferably in the range 1 to 5 square microns.

[0091] Another important measure in the fabrication of this biosensor is the impedance of the hydrophobic layer 5 and the sealing impedance of the supporting membrane. The hydrophobic layer serves to secure a tight, high impedance seal at the edge of the membrane interfacial area, immobilizing lipids and preventing leakage of ions at the edge of the membrane interfacial area, m some embodiments, the impedance of the surrounding hydrophobic layer 5 is preferably greater than 10¹² ohms and the impedance of the supporting lipid membrane is greater than 10¹¹ ohms.

[0092] By way of example, the steps taken in fabricating this biosensor include: taking a support material and depositing on it in sequence the layers described above; etching the resulting multi-layer substrate to form a well or wells with the required geometry; and cleaning the etched substrate and reducing gold oxides on its surface. In some embodiments, the support material is a single crystal silicon wafer, and the electrode area is wet etched using a photolithographic patterning approach. It is preferred that the gold electrode consists of a freshly evaporated or sputtered gold electrode.

[0093] Once the well or wells are formed, a membrane can then be formed in the well or wells by forming a solution containing amphiphilic molecules, linkers and ionophores; and contacting the cleaned gold base of the well with the solution to form a first layer membrane comprising a closely packed array of amphiphilic molecules and a plurality of ionophores. The first layer membrane is connected to the electrode by means of a linker group. The ionophores are preferably gramicidin A or an analogue thereof. The ionophores may be further biotinylated to enable subsequent binding of streptavidin or analogues thereof. Upon the formation of the first layer membrane, the silicon wafer is rinsed with a suitable organic solvent, and the excess organic solvent is then removed.

[0094] A solution of lipid and a plurality of ionophores dispersed in a suitable solvent is further formed, and then contacted with the electrode containing a first layer membrane to form a second layer membrane. Preferably, the electrode is immersed in the solution immediately upon removal of the excess organic solvent. The electrode is further raised with an aqueous solution, and then removed from the solution to allow draining.

[0095] The membrane so formed extends across the interfacial area. In some embodiments, the solvent for the adsorbing solutions and for the rinsing steps is ethanol. It is preferred that the solvent is removed by rapid air dry.

[0096] Once formed, the membrane can be further functionalised in order to provide for the detection of the presence of analyte by the membrane-based biosensor. One convenient method to attach appropriate receptors to the surface of a membrane is by using streptavidin, avidin or one of the related biotin binding-proteins as a means of coupling a wide range of receptors onto a biotinylated gramicidin ion channel or membrane-spanning lipid.

[0097] The process of functionlizing the membrane may include the steps of: adding a solution of streptavidin, avidin, neutravidin, avidin or streptavidin derivative onto the surface of the membrane of a membrane-based biosensor according to the present invention in which at least a portion of the components are biotinylated; rinsing the electrode with an aqueous solution in order to remove excess streptavidin, avidin, neutravidin or other avidin or streptavidin derivative; adding a solution of a biotinylated receptor molecule so that the receptor molecule is attached to the membrane via the biotin-streptavidin-biotin link; rinsing the coated electrode with an aqueous solution; removing the electrode from the aqueous solution and allowing to drain, such that a bead of water is retained within the well of the device; and storing the electrode at reduced temperature of preferably between -2O⁰C and +5°C. m some embodiments, the biotinylated receptor molecules are introduced into the well of the device using an ink jet robot.

Example 4. Effect of Membrane Thickness on Channel Lifetime

[0098] A lipid gramicidin bilayer was prepared on a gold-coated glass electrode. The inner layer consisted of lOuM full membrane spanning lipid (MSLXXB), 12mM glyceromonophytanylether (GMPE) andlmM double-length reservoir half-membrane- spanning phytanyl lipids (DLP). The outer layer consisted of 28mM monoalkyl glycerol lipids and biotinylated gramicidin (GaXB) 100,000:1. The chain length of the outer membrane lipids ranged from 16 to 22 carbon atoms. The impedance of the membrane was measured at frequencies from 2 to 1000 Hz. As shown in Figure 7, membrane impedance decreased with chain length, indicating that the channel lifetime decreases with membrane thickness.

Example 5. Individually Resolved Ion Channel Dimers

[0099] Figures 8a-8c illustrate individually resolved ion channel dimers obtained from a bilayer lipid membrane supported in a Teflon septum, containing a varying number of conducting biotinylated ion channel dimers (i.e., from 0 to 5) varying with time. Figure 8a illustrates the temporal frequency distribution of the conducting ion channel dimers prior to the addition of streptavidin. Figure 8b illustrates temporal frequency distribution of the conducting ion channel dimers following the addition of 1 nM streptavidin to the sample solution. The results show that there are fewer conducting dimers as streptavidin binds to the ion channels and further prevents the formation of conducting dimers. Figure 8c illustrates that the channel conduction is eliminated after a period of binding of streptavidin to the biotinylated ion channels.

[00100] Figures 9A-9E illustrate a simulation of the effect of channel density on the ability of the sensor to resolve current pulses. Figure 9A simulates two channels within the interfacial area of the membrane, with typical formation rates and lifetimes. Figure 9B simulates 5 channels, Figure 9C simulates 10 channels, Figure 9D simulates 50 channels, and Figure 9E simulates 200 channels. It is evident that for more than 50 channels per electrode under these conditions the ability to discriminate current pulses decreases to the point where there is insufficient information available for an analysis of the target concentration. The maximum number of concurrent conducting channels for the present invention is 50.

Example 6. Detection of Analyte Molecules Using An Array of Electrodes

[00101] Figure 10 illustrates a theoretical plot of the number of electrodes in an array of 1000 gated as a function of analyte concentration. This figure is derived from the following sequence of calculations: 1) The conduction per channel is ~ 10ⁿΩ. The conduction of the sealed membrane is ~100kΩ/ cm² = 10¹³Ω/μm².

2) Thus the absolute detection limit of a gramicidin in the membrane is one gramicidin in ~ 10¹³/10^π/μm²= 10²/μm² = 100/μm², ie: a 10 x lOμm electrode.

3) The distance over which a channel will diffuse in the 2^nd layer, unimpeded by the 1^st layer gramicidin ~ (2 x D x t) ⁰⁵ = (2 x 10^'8cm²/s x 180s) ^α5 = 1.9 x 10^"3cm = 19μm.

4) However if the mobile gramicidin encounters a 1^st layer gramicidin then add a gramicidin lifetime component, normally ~ 100s.

5) Thus the reasonable limit of detection of one channel is a lOμm x lOμm electrode.

6) The maximal density of membrane spanning capture species (MSL4B) ~ 10 IQ molecules/cm².

7) The maximal "on rate" for the capture antibody ~ 10⁶ M⁴S^"1.

8) Thus for a target concentration N_A, the arrival rate of target at the MSL4XB is: 10¹⁰xl0⁶x N_A molecules/cm²/s/M.

9) Thus to gate off one gramicidin in a 10 x lOμm electrode in 100s, ie: 10⁶gramicidins per cm²/100s requires 10⁶gramicidins/cm² = 10²(s) xlO¹⁰ (molecules/cm²) x 10⁶(/s/M) x NA(M) or: N_A= IpM in 100s.

10) At IpM the target analyte is spaced by ~10μm. Thus the collision rate with an array of lOμm x lOμm electrodes spaced by ~10μm will result in all electrodes being gated in 100s. For 1000 electrodes = array of 33 x 33 covering an area of ~lmm².

11) At IpM the target analyte is spaced by ~ 10 μm. This will result in 1/(1.0 x 1.0) ~ 100.0% of the electrodes being gated.

12) At 10OfM the target analyte is spaced by ~ 21.5μm. This will result in 1/(2.15 x 2.15) ~ 21.5% of the electrodes being gated.

13) At 1OfM the target analyte is spaced by ~ 46.4μm. This will result in 1/(4.64 x 4.64) ~ 4.6% of the electrodes being gated. 14) At IfM the target analyte is spaced by ~ lOOμm. This will result in 1/(10.0 x 10.0) ~ 1% of the electrodes being gated.

[00102] These calculations provide approximate estimates of the anticipated modulation of the ion channel currents that will result over the analyte concentration range of 1 pm to 1 fin (Table 1). 1% to 100% of the electrodes in a 1000 electrode array will be influenced and will elicit alterations in the pattern of current pulses. As will be apparent to one of ordinary skill in the art of digital data analysis, algorithms exist that permit the enhancement of these variations and their interpretation as indicative of the analyte species and concentration in a sample solution.

Table 1

[00103] It will be obvious to those skilled in the art that various changes may be made without departing from the scope of the invention, which is not to be considered limited to what is described in the specification.

Claims

WHAT IS CLAIMED IS:

1. A method of detecting binding events between receptor molecules and analyte molecules, comprising the steps of: a) obtaining a membrane-based biosensor which comprises :

1) a lipid bilayer membrane which includes a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs, the first HMSMs are provided in a first layer of the bilayer membrane and the second HMSMs are provided in a second layer of the bilayer membrane, wherein the second HMSMs are capable of lateral movement within the membrane and the first HMSMs are prevented from lateral movement in the membrane;

2) receptor molecules attached to at least a portion of the second HMSMs; and

3) a sensing electrode, wherein the first layer of the bilayer membrane is tethered to the sensing electrode, wherein conducting channels are formed by alignment of the first HMSMs and the second HMSMs and 50 or fewer concurrent conducting channels are formed per sensing electrode; b) contacting the receptor molecules with an analyte sample containing analyte molecules, wherein the binding of the receptor molecules to the analyte molecules causes a change in the temporal frequency distribution of the conducting channels; c) measuring the change in current through the membrane over a course of time; and d) resolving the current measurement into an integer number of concurrent conducting channels, thereby detecting the binding events between the receptor molecules and analyte molecules.

2. A method of detecting the presence of analyte molecules of interest in a sample, comprising the steps of: a) obtaining a membrane-based biosensor which comprises:

2) receptor molecules attached to at least a portion of the second HMSMs; and

3) a sensing electrode, wherein the first layer of the bilayer membrane is tethered to the sensing electrode, wherein conducting channels are formed by alignment of the first HMSMs and the second HMSMs and 50 or fewer concurrent conducting channels are formed per sensing electrode; b) contacting the receptor molecules with the sample suspected of containing the analyte molecules of interest, wherein the binding of the receptor molecules to the analyte molecules causes a change in the temporal frequency distribution of the conducting channels; c) detecting the change in current through the membrane; and d) resolving the current measurement into an integer number of concurrent conducting channels, thereby detecting the presence of the analyte molecules of interest in the sample.

3. The method of claim 1 or 2, wherein interfacial area between the first and second layers of the lipid bilayer membrane is between 1 and 10000 square microns.

4. The method of claim 3, wherein the interfacial area is less than 100 square microns.

5. The method of claim 4, wherein the interfacial area is less than 10 square microns.

6. The method of claim 5, wherein the interfacial area is less than 1 square micron.

7. The method of claim 1 or 2, wherein the density of the second HMSMs in the second layer of the bilayer membrane is less than the density of the first HMSMs in the first layer of the bilayer membrane.

8. The method of claim 7, wherein the ratio of the HMSMs in the first layer of the bilayer membrane to the HMSMs in the second layer of the bilayer membrane is greater than 10.

9. The method of claim 1 or 2, wherein the lipid bilayer membrane has a thickness in the range of 10 to 40 Angstroms.

10. The method of claim 9, wherein the lipid bilayer membrane has a thickness in the range of 28 to 33 Angstroms.

11. The method of claim 1 or 2, wherein the conducting channels have lifetimes of less than 1000 seconds.

12. The method of claim 11, wherein the conducting channels have lifetimes of between 10 to 100 seconds.

13. The method of claim 11, wherein the conducting channels have lifetimes of between 1 to 10 seconds.

14. The method of claim 11, wherein the conducting channels have lifetimes of between 50 milliseconds to 1 second.

15. The method of claim 1 or 2, wherein the sensing electrode is a metal.

16. The method of claim 15, wherein the sensing electrode is palladium, gold, platinum, silver or mixtures thereof.

17. The method of claim 16, wherein the sensing electrode is gold.

18. The method of claim 17, wherein the first layer of the lipid bilayer membrane is tethered to the gold electrode by sulfur-containing groups.

19. The method of claim 1 or 2, wherein the ionophore is an ion channel drawn from the family of ion channels comprising β barrel structures, or the family of ion channels comprising non-ribosomally formed ion channels.

20. The method of claim 1 or 2, wherein the ionophore is an ion carrier.

21. The method of claim 1 or 2, wherein the ionophores are selected from the group consisting of gramicidin and valinomycin.

22. The method of claim 21, wherein the ionophores are gramicidin A.

23. The method of claim 1 or 2, wherein the binding of the receptor molecules to the analyte molecules includes bindings between enzymes and substrates, antibodies and antigens, chelators and metal, cell surface receptors and receptor ligands, oligimer DNA probes and single- or double-stranded DNA, aptimers and aptimer target sequences.

24. The method of claim 1 or 2, wherein the receptor molecules are biotinylated.

25. The method of claim 24, wherein the second HMSMs comprise biotin modified gramicidin monomers, and the biotinylated receptor molecules are attached to the at least a portion of second HMSMs via the biotin-streptavidin-biotin link.

26. A method of detecting binding events between receptor molecules and analyte molecules, comprising the steps of: a) obtaining a membrane-based biosensor which comprises:

2) receptor molecules attached to at least a portion of the second HMSMs; and

3) an array of sensing electrodes, wherein the first layer of the bilayer membrane is tethered to the array of the sensing electrodes, wherein conducting channels are formed by alignment of the first HMSMs and the second HMSMs and 50 or fewer concurrent conducting channels are formed per sensing electrode; b) contacting the receptor molecules with an analyte sample containing analyte molecules, wherein the binding of the receptor molecules to the analyte molecules causes a change in the temporal frequency distribution of the conducting channels; c) measuring the change in current through the membrane over a course of time; and d) resolving the current measurement into an integer number of concurrent conducting channels, thereby detecting the binding events between the receptor molecules and analyte molecules.

27. A method of detecting the presence of analyte molecules of interest in a sample, comprising the steps of: a) obtaining a membrane-based biosensor which comprises:

2) receptor molecules attached to at least a portion of the second HMSMs; and

3) an array of sensing electrodes, wherein the first layer of the bilayer membrane is tethered to the array of the sensing electrodes, wherein conducting channels are formed by alignment of the first HMSMs and the second HMSMs and 50 or fewer concurrent conducting channels are formed per sensing electrode; b) contacting the receptor molecules with the sample suspected of containing the analyte molecules of interest, wherein the binding of the receptor molecules to the analyte molecules causes a change in the temporal frequency distribution of the conducting channels; c) detecting the change in current through the membrane; and d) resolving the current measurement into an integer number of concurrent conducting channels, thereby detecting the presence of the analyte molecules of interest in the sample.

28. A biosensor comprising a lipid bilayer membrane and at least one sensing electrode, wherein the lipid bilayer membrane is tethered to the at least one sensing electrode, wherein the membrane comprises: a plurality of ionophores comprising first half membrane spanning monomers (HMSMs) and second HMSMs, the first HMSMs are provided in a first layer of the bilayer membrane and the second HMSMs are provided in a second layer of the bilayer membrane, wherein the second HMSMs are capable of lateral movement within the membrane and the first HMSMs are prevented from lateral movement in the membrane; and receptor molecules attached to at least a portion of the second HMSMs; wherein thickness of the lipid bilayer membrane and density of the first and second HMSMs are selected in such a way that a change in current through the membrane due to formation of conducting channels can be detected, wherein the conducting channels are formed by alignment of the first HMSMs and the second HMSMs and 50 or fewer concurrent conducting channels are formed per sensing electrode.

29. The biosensor of claim 28, wherein interfacial area between the first and second layers of the lipid bilayer membrane is between 1 and 10000 square microns.

30. The biosensor of claim 29, wherein the interfacial area is less than 100 square microns.

31. The biosensor of claim 30, wherein the interfacial area is less than 10 square microns.

32. The biosensor of claim 31, wherein the interfacial area is less than 1 square micron.

33. The biosensor of claim 28, wherein the density of the second HMSMs in the second layer of the bilayer membrane is less than the density of the first HMSMs in the first layer of the bilayer membrane.

34. The biosensor of claim 33, wherein the ratio of the HMSMs in the first layer of the bilayer membrane to the HMSMs in the second layer of the bilayer membrane is greater than 10.

35. The biosensor of claim 28, wherein the lipid bilayer membrane has a thickness in the range of 10 to 40 Angstroms.

36. The biosensor of claim 35, wherein the lipid bilayer membrane has a thickness in the range of 28 to 33 Angstroms.

37. The biosensor of claim 28, wherein the conducting channels have lifetimes of less than 1000 seconds.

38. The biosensor of claim 37, wherein the conducting channels have lifetimes of between 10 to 100 seconds.

39. The biosensor of claim 37, wherein the conducting channels have lifetimes of between 1 to 10 seconds.

40. The biosensor of claim 37, wherein the conducting channels have lifetimes of 50 milliseconds to 1 second.

41. The biosensor of claim 28, wherein the interfacial area is bounded by hydrophobic solid.

42. The biosensor of claim 28, wherein the interfacial area is bounded by gas.

43. The biosensor of claim 42, wherein the gas is inert gas or air.

44. The biosensor of claim 28, wherein the interfacial area is established at the time of introduction of a sample solution to the sensing electrode.

45. The biosensor of claim 28, wherein the at least one sensing electrode is a metal.

46. The biosensor of claim 45, wherein the at least one sensing electrode is palladium, gold, platinum, silver or mixtures thereof.

47. The biosensor of claim 46, wherein the at least one sensing electrode is gold.

48. The biosensor of claim 47, wherein the first layer of the lipid bilayer membrane is tethered to the at least one gold electrode by sulfur-containing groups.

49. The biosensor of claim 28, wherein the ionophore is an ion carrier.

50. The biosensor of claim 28, wherein the ionophore is gramicidin

51. The biosensor of claim 50, wherein the ionophores are gramicidin A.

52. The biosensor of claim 28, wherein the receptor molecules are biotinylated.

53. The biosensor of claim 52, wherein the second HMSMs comprise biotin modified gramicidin monomers, and the biotinylated receptor molecules are attached to the at least a portion of the second HMSMs via the biotin-streptavidin-biotin link.