Decreasing the Limits of Detection and Analysis Time of Ion Current

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Decreasing the Limits of Detection and Analysis Time of Ion Current Rectification Biosensing Measurements via a Mechanically Applied Pressure Differential Anna E. P. Schibel and Eric N. Ervin* Electronic BioSciences, 421 Wakara Way, Suite 328, Salt Lake City, Utah 84108, United States S Supporting Information *

ABSTRACT: Improving on the analytical capabilities of a measurement is a fundamental challenge with all assays, particularly decreasing the limit of detection while maintaining a practical associated analysis time. Of late, ion current rectification (ICR) biosensing measurements have received a great deal of attention as an analyte-specific, label-free assay. In ICR biosensing, a nanopore coated with an analyte specific binding molecule (e.g., an antibody, an aptamer, etc.) is used to detect a target analyte based on the ability of the target analyte to alter the ICR response of the nanopore upon it binding to the aperture interior. This binding changes the local surface charge and/or size of the nanopore aperture, thus altering its ICR response in a time dependent manner. Here, we report the ability to enhance the transport of a target analyte molecule to and through the aperture of an antibody modified glass nanopore membrane (AMGNM) with the application of a mechanically applied pressure differential. We demonstrate that there is an optimal pressure that balances the flux of the target analyte through the AMGNM aperture with its ability to be bound and detected. Applying the optimal pressure differential allows for picomolar concentrations of the cleaved form of synaptosomal-associated protein 25 (cSNAP-25) to be detected within the same analysis time as micromolar concentrations detected without the use of the pressure differential. The methodology presented here significantly expands the utility of ICR biosensing measurements for detecting lowabundance biomolecules by lowering the limit of detection and reducing the associated analysis time.

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aptamer, etc.) is used to detect a specific target analyte (e.g., an antigen, nucleic acid, or peptide) as it binds to the aperture. As the analyte binds and coats the internal surface of the pore, the size and charge of the aperture are altered such that a change in the ICR response of the pore is observed.3−10,17 Recently, we reported on the dynamic and quantitative nature of the ICR biosensing measurement using the antibody modified glass nanopore membrane (AMGNM). 10 An AMGNM is a single, conical-shaped nanopore embedded in a thin glass membrane with a chemically protected exterior glass surface (to prevent external, nonspecific adsorption) and an antibody-functionalized nanopore interior. Using the AMGNM, we demonstrated the ability to quantify the concentration of a given target analyte via the rate of ICR change imparted by the target binding to the aperture surface when measuring the ICR of the pore at regular time intervals.10 In these initial studies, the limit of detection (micromolar levels) and associated analysis time (hours) were the result of a diffusion-limited system, meaning that in order for the analyte to be detected, it had to diffuse via Brownian motion into the pore before it could be bound.

urrently, there is a significant amount of interest in the development of analyte-specific assays for the detection and characterization of low-abundance biomolecules, including proteins and nucleic acids. An outstanding challenge, however, is to develop these biosensors with picomolar or better concentration sensitivities (i.e., limits of detection) while at the same time exhibiting quick analysis times (e.g., below 1 h).1,2 One type of measurement with significant potential in this arena is the ion current rectification (ICR) biosensing measurement, a nanopore-based technique that is label-free and specific to the analyte of interest.3−10 ICR is defined as an increase in ion conduction at a given polarity relative to a decrease in ion conduction for the same voltage value at the opposite polarity or a deviation from a linear ohmic current response.11−13 This effect occurs within conical-shaped pores due to the voltage dependent solution conductivity within the aperture.11 The degree to which a conical-shaped pore exhibits an asymmetrical current−voltage response is based on the orifice size, surface charge, and Debye length. 12,14 This current rectification is maximized at intermediate bulk ion concentrations, decreasing from this maximum at high concentrations due to the electrical screening of the surface charge as well as at low concentrations due to a fixed number of charge-carrying ions.15,16 In an ICR biosensing measurement, a conical, solid-state nanopore coated with an analyte-specific binding molecule (e.g., biotin, an antibody, an © XXXX American Chemical Society

Received: February 25, 2015 Accepted: June 4, 2015

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AMGNM, thus reducing the associated analysis time and decreasing the limit of detection. Utilizing the determined optimal pressure differential allows us to characterize picomolar cSNAP-25 concentrations, which is a concentration 106 orders of magnitude lower than can practically be achieved without the use of a convective flow, within a similar analysis time frame (i.e., a few hours). We speculate that this optimal pressure represents a balance between the target antigen molecular transport rate to the AMGNM aperture, its translocation velocity, and its ability to bind and be detected via the ICR biosensing measurement.

Here, in an effort to better understand the underlying detection mechanism of the ICR biosensing measurement, we examine the effect of a driving force (i.e., a mechanically applied pressure differential) on the transport of our target analyte to and through the aperture of the AMGNM, a strategy that has been previously unexplored in ICR biosensing experiments. The mechanical pressure differential applied across the AMGNM orifice generates pressure-driven solution flow into and through the AMGNM, as schematically depicted in Figure 1A. This increases the molecular transport rate of the analyte

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EXPERIMENTAL SECTION Chemicals and Materials. NaCl (Sigma), ethylenediaminetetraacetic acid (Sigma), 2-mercaptoethylamine-HCl (Thermo Scientific), sulfosuccinimidyl 4-[N-maleimidomethyl]cyclohexane-1-carboxylate (Thermo Scientific), BupH phosphate buffered saline packs (Fisher Scientific), 3cyanopropyldimethylchlorosilane (Gelest), and 3-aminopropyldimethylethoxysilane (Gelest) were used as received. Acetonitrile (Sigma) was stored over a 3 Å molecular sieve. Monoclonal antibody (IgG) specific to cSNAP-25 (Antibodies-Online) was received as a cell supernatant from a murine host, stored at −20 °C when not in use, and purified (Antibody Clean-up Kit, Pierce) prior to use. cSNAP-25 (AntibodiesOnline) was received as a lyophilized powder, diluted to 1 mg/ mL in H2O, and stored at −20 °C when not in use. The cSNAP-25 used in these studies is a synthetic peptide comprising amino acids 184−197 of the actual synaptosomalassociated protein 25, with a cysteine residue addition to the Nterminal (Cys-Lys-Ala-Asp-Ser-Asn-Lys-Thr-Arg-Ile-Asp-GluAla-Asn-Glu), and exists as a water-soluble ∼1.6 kDa peptide with a net −1 charge at pH 7 due to four acidic residues (Asp and Glu) relative to three basic residues (Arg and Lys). Interference species used for control experiments include tris(2-carboxyethyl)phosphine (Sigma), botulinum toxin type A light chain (List Laboratories), and monoclonal antibody specific to botulinum toxin type A light chain (R&D Antibodies), which were used as received, and uncleaved synaptosomal-associated protein 25 (Abnova) was filtered via centrifugation prior to use. All aqueous solutions were prepared using H2O (18 MΩ cm) from a Barnstead E-pure water purification system. Antibody-Modified Glass Nanopore Membrane (AMGNM) Fabrication. The complete fabrication of the AMGNM has previously been described in detail.10 The AMGNMs utilized in these studies had their outer glass surface chemically protected with 3-cyanopropyldimethylchlorosilane, in order to limit nonspecific binding,21 while their interior pore surface was functionalized with the monoclonal antibody specific to cSNAP-25, using standard thiol-based bioconjugate techniques.22−24 This approach is expected to yield an antibody surface coverage of 50−200 ng/cm2,23,24 although the actual coverage is likely in the lower end of this range due to the confined geometry of the conical pore. Nevertheless, the resulting surface coverage is sufficient for biosensor applications, as demonstrated via the detection of cSNAP-25.10 Each resulting AMGNM contained a single, conical-shaped, antibody-coated nanopore that was approximately 50 μm in length before opening up to the internal diameter of the glass capillary (0.75 mm i.d. and 1.65 mm o.d., Dagan Corp.) in which it was fabricated, resulting in aperture dimensions of 50 to 250 nm in radius, with a half-cone angle of ∼10°.25,26 While the larger

Figure 1. (A) Schematic drawing depicting the antibody modified glass nanopore membrane (AMGNM), not drawn to scale. The pore interior is coated with an antibody and while the exterior glass surface is protected with a nonreactive cyano-silane coating (depicted in green). A negative pressure differential is applied across the nanopore that causes volumetric flow through the aperture, resulting in transport of antigen molecules from bulk solution to the pore interior, where they are bound and detected. (B) Illustration of a binding curve resulting from the detection of target antigen as a function of time, where the slope of the curve is representative of the antigen detection rate.

through the aperture, thus increasing it detection rate, as determined from the slope of the AMGNM binding site saturation curve, as illustrated in Figure 1B. We demonstrate that the limit of detection and analysis time of an ICR biosensing measurement can be altered with mechanical pressure differentials using AMGNMs coated with a monoclonal antibody specific to the cleaved form of synaptosomalassociated protein 25 (cSNAP-25), the cleavage product of the botulinum toxin type A light chain in neuronal cells.18−20 By mapping out a profile of the detection rate of cSNAP-25 as a function of the applied pressure, we are able to determine the optimal pressure for cSNAP-25 detection or the pressure that allows us to detect cSNAP-25 at the fastest possible rate via the B

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the initial ICR baseline level. Through our experience, AMGNMs can be reused approximately 11 times before the initial baseline ICR response can no longer be restored, as previously described.10 The applied voltage, experimental cell temperature, and data acquisition were all controlled via an inhouse written program in LabVIEW (National Instruments), and data analysis was performed using Igor Pro 6 (WaveMetrics).

aperture (approaching >200 nm) pores are not actually nanopores but more correctly identified as “mesopores,” all AMGNMs discussed herein are referred to as nanopores. Previous studies by Lan et al. have shown measurable ICR for glass nanopores with radii ranging from 30 to 250 nm.27,28 Similarly, in this study AMGNMs ranging from 50 to 250 nm in radius yielded a sufficiently measurable ICR response for the detection of cSNAP-25, due the surface charges and Debye length at the pore walls still affecting the voltage dependent solution conductivity within the nanopore, even though the Debye length represents only a fraction of the aperture crosssectional area. All references to the pore radius refer to the tip opening of the AMGNM aperture. Fully fabricated anti-cSNAP25 AMGNMs were stored dry at 4 °C, in between experiments. Ion Current Rectification Measurements. All ICR measurements reported herein were recorded using an aqueous 150 mM NaCl and 100 mM sodium phosphate (pH 7.2) buffered electrolyte solution. NaCl (150 mM, Debye length ∼1 nm)29 was utilized as it allows for the ICR response of the AMGNM to be characterized and is relevant to the typical biological conditions employed to assess cSNAP-25 concentrations after exposure to botulinum toxin type A.30,31 The electrolyte strength was therefore not altered in attempts to optimize the ICR response of the AMGNM. Ag/AgCl electrodes were prepared by electroplating a thin layer of AgCl on a 0.25 mm diameter Ag wire.10 The AMGNM was filled with the buffered electrolyte solution, and the internal Ag/AgCl electrode was placed inside the AMGNM capillary. The AMGNM was then inserted into a custom-made polycarbonate experimental cell that contained a second, external Ag/AgCl electrode and the same buffered electrolyte used inside the AMGNM. The back of the AMGNM capillary was closed off with a custom built capillary holder, creating an airtight seal over the AMGNM interior to control the pressure within the AMGNM. This holder was connected to a 5 mL syringe that was used to apply pressure differentials (up to ±350 mmHg), as monitored with an in-line analog pressure gauge relative to the exterior of the AMGNM, similar to setups previously described.21,32 ICR measurements were performed using a custom built, high impedance, low noise amplifier to apply a voltage sweep across the nanopore, relative to the internal electrode, and record the resulting current response. This system was used in conjunction with its associated temperature control instrumentation to maintain a constant temperature (20 ± 1 °C) for all experiments. A voltage of ±1 V was used for all antigen binding characterization studies, and all current values were measured to two decimal places. Prior to antigen detection, the baseline ICR response of the AMGNM was measured to verify stability and establish an initial baseline ICR level for that particular AMGNM at the specified pressure differential (0 mmHg to −270 mmHg), which was then held constant throughout the measurement. The cSNAP-25 antigen (1 pM to 100 μM) was then introduced into the experimental cell and ICR measurements were made at 5 to 60 min intervals, determined from the apparent binding rate, until a stable saturation ICR level was reached. In order to detect higher antigen concentrations (above 100 μM), the measurement interval would need to be decreased. The voltage bias was always held at 0 mV between successive ICR measurements. After antigen detection, AMGNMs were flushed/rinsed with buffer under applied voltage and positive pressure (relative to the external solution) in order to remove the antigen from the antibody and restore

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RESULTS AND DISCUSSION ICR Biosensing Measurement Using the AMGNM. The degree to which the current is rectified by a given nanopore, i.e., its ICR response, is quantified via its rectification ratio (RRV), which is calculated by dividing the current amplitude at a given positive voltage by the current amplitude at the equal, but opposite voltage (RRV = |i+V|/|i−V|).10 Prior to the introduction of the target analyte, the AMGNM has a relatively neutral surface due to the minimal net charge of the immunoglobulin G (IgG) antibody coating the AMGNM aperture surface,10 yielding a nearly ohmic current as a function of voltage response. Once the slightly negatively charged cSNAP-25 target analyte is introduced into solution, it begins to bind to and coat the interior of the aperture, inducing a net negative charge and nominally reducing the AMGNM aperture diameter (by 7% at the most, assuming the cSNAP-25 molecule resides is its fully extended form, as estimated from the hydrodynamic radius of each residue),33 which results in an increased ICR response. A plot of the current versus voltage, Figure 2A, shows a clear transition from the nearly ohmic response of the AMGNM in the absence of antigen (solid black trace) to a more rectified response, after the AMGNM was exposed to 5 μM cSNAP-25 (dotted red trace); RR1 V increases from 1.02 to 1.35 after the addition of cSNAP-25. The ability to utilize the change in the RRV as a function of time (RRtV) to quantify the presence of the target analyte is shown in Figure 2B. Here, the characterization of 0.5 μM, 5.0 μM, and 50.0 μM cSNAP-25 is shown, along with the baseline AMGNM response in the absence of antigen (i.e., a blank control), for a diffusive system (0 mmHg). The data in Figure 2B are normalized in order to compare the rate of rectification change across different AMGNMs irrespective of the actual rectification ratio values, which may vary between experiments. The % AMGNM saturation is calculated by the following expression:10 %AMGNM saturation =

(RRtV − RRtV= 0) t=0 (RRSS V − RR V )

× 100 (1)

where RRt=0 V is the initial RRV measured before cSNAP-25 is introduced to the AMGNM, and RRSS V is the final steady state RRV value reached upon the complete saturation of the AMGNM with cSNAP-25. Among different AMGMNs over the course of this work, the change in RRV, or from RRt=0 V to RRSS V , was found to typically vary anywhere from 0.2 to 0.5, which is a significant change relative to the baseline stability of the AMGNM. As is depicted in Figure 2B, the AMGNM baseline is highly stable; here the maximum ΔRR1 V variation was found to be ±0.01 in the absence of target antigen. In cases where sigmoidal binding curves were not observed and the ΔRRV was less than 0.02, such as in the case of our control experiments, these data were normalized using a ΔRR1 V of 0.35, the average rectification change imparted by cSNAP-25 binding to the AMGNM found during the course of this work. C

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the concentration of introduced antigen; larger B values indicate a steeper curve, faster binding, and higher concentration of cSNAP-25. C has the units of time and is inversely proportional to the concentration of the antigen; larger values indicate slower binding and thus lower concentrations of cSNAP-25. Through our experience, B is easier to compare between experiments than C, as it allows for any variations in sample introduction (i.e., the reaction start time) to be excluded from the analysis. For the representative curves shown in Figure 2B, the steepness coefficient (B) of each curve was determined to be 4, 7, and 20, for 0.5 μM, 5.0 μM, and 50.0 μM cSNAP-25, respectively. In this case, from a time-sale based perspective, the measurement is limited to the detection of nanomolar to micromolar analyte concentrations. For example, it takes >6 h to detect 0.5 μM cSNAP-25 when the transport of molecules into the AMGNM is limited by diffusion (in the absence of a pressure differential, 0 mmHg). It should be noted that the applied voltage across the AMGNM is set to zero between individual RR1 V measurements and that cSNAP-25 is only slightly charged, as described in the Experimental Section. Electrophoretic migration of cSNAP-25 is therefore unlikely to contribute to its transport to and through the aperture of the AMGNM. Applied Pressure Differentials with the AMGNM. As a means of increasing the rate of antigen detection via the AMGNM, and thus decreasing the limit of detection for the same measurement time (i.e., characterizing low cSNAP-25 concentrations more quickly), a pressure differential was implemented to enhance the transport of the target antigen to the AMGNM orifice. A negative pressure was manually applied to the back of the AMGNM to create a pressure gradient across the aperture. As depicted in Figure 1A, low pressure then exists inside the AMGNM relative to the higher pressure outside the AMGNM, creating volumetric flow from bulk solution into the AMGNM. A comparison of the detection rate for a given antigen concentration in both the absence and presence of a pressure differential is depicted in Figure 3. Here, the detection of 500

Figure 2. (A) Current as a function of voltage response of a 230 nm radius anti-cSNAP-25 AMGNM before introducing cSNAP-25 (solid black trace) and after the introduction and binding of 5 μM cSNAP-25 to saturation (dotted red trace). (B) Representative binding curves for the % AMGNM saturation as a function of time for 0.5 μM cSNAP-25 (black diamonds), 5.0 μM cSNAP-25 (red circles), and 50 μM cSNAP-25 (green triangles). The 0% AMGNM saturation baseline value (i.e., the response of the AMGNM in the absence of the antigen) is shown for comparison (gray asterisks). Data were taken with 50− 250 nm radius AMGNMs, at 0 mmHg applied pressure.

We speculate that the sigmoidal shape to the binding curve is a result of sample diffusion (after introduction to the experimental cell) to the AMGNM aperture, followed by period of unrestricted binding, and finally, the approach to a saturated sensor (as the number of available binding sites is reduced), similar to previously described adsorption curves.34 Each data set in Figure 2B is fit to a 5-parameter nonlinear regression model:10 ⎛ ⎜ D ⎜ F (t ) = A 0 + ⎜ ⎛ ⎜ ⎜1 + t C ⎝⎝

⎞ ⎟ ⎟ E (−B) ⎞ ⎟ ⎟ ⎟ ⎠ ⎠

()

(2)

where A0 is the minimum asymptote or the initial level of rectification when no antigen is present (i.e., 0% saturation), B is the steepness coefficient of the curve indicating the rate of cSNAP-25 binding, C is the inflection point of the sigmoidal curve, D is the maximum asymptote or final level of rectification (i.e., 100% saturation), E is the asymmetry factor, and t is time. These fit parameters are depicted in the Supporting Information for a representative cSNAP-25 binding curve, Figure S1. Both B and C provide information on the kinetics or rate of the AMGNM binding reaction and are thus dependent on the antigen concentration. B is a dimensionless parameter that is directly proportional to the AMGNM binding rate and

Figure 3. Representative binding curves for the % AMGNM saturation as a function of time for 500 nM cSNAP-25 detected in the absence of a pressure differential (0 mmHg, black diamonds) and with an applied pressure differential (−90 mmHg, blue diamonds). Data taken at −90 mmHg were acquired with a 75 nm radius AMGNM, while 0 mmHg were taken with a 115 nm radius AMGNM. D

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cSNAP-25 concentration over a pressure range of 0 to −270 mmHg (data were plotted on a log−log scale for visualization purposes). All data points on this plot represent individual binding experiments in which cSNAP-25 was characterized using an AMGNM in combination with the specified pressure differential. Each pressure dependent data set in Figure 4A was then fit using the power series model log(y) = log(A) + m log(x),10 where y is the steepness coefficient (B), A is the scaling coefficient, x is the concentration of antigen, and m is the slope. A power model is used here to account for the nonlinear trend in the steepness coefficient as a function of cSNAP-25 concentration. We speculate that this occurs due to the fact that at higher antigen concentrations the rate of binding begins to approach the time limit of our measuring interval (∼5 min) for these studies. The fact that all of the data sets share the same slope (m = ∼0.4), indicates that although the use of pressure alters the binding rate, we are still able to measure a similar binding process across all pressures studied. It should be noted that for the data presented in Figure 4A, there is an average coefficient of variation of 26% in the binding rate (i.e., steepness coefficient, B) for a given pressure and concentration. This large deviation between experiments is likely due to the relatively large dispersion in AMGNM aperture radii used during this investigation (50−250 nm in radius), which will affect the molecular transport rate to the aperture as well as the time required for the target analyte to traverse to the aperture surface and be detected once it has entered the AMGNM aperture, as further described below. As is depicted in Figure 4A, a larger pressure differential increases the scaling coefficient, A, between pressures of 0 and −210 mmHg. However, at pressure differentials greater than −210 mmHg, the detection rate begins to decrease, leading to longer analysis times relative to the −210 mmHg data set. To evaluate the influence of the applied pressure differential on the cSNAP-25 binding rate, the scaling coefficient (A) for each pressure dependent data set in Figure 4A was plotted as a function of applied pressure in Figure 4B, highlighting that the scaling coefficient peaks at ∼−210 mmHg. This peak in the pressure profile reflects the pressure differential at which cSNAP-25 detection occurs at the fastest possible rate utilizing the AMGNM. At pressures up to −210 mmHg, the binding rate likely occurs faster with increasing pressure because molecules are delivered more quickly to the AMGNM aperture due to their enhanced flux, increasing the molecule/aperture encounter rate and thus the apparent rate of binding and detection. However, at pressures greater than −210 mmHg, the binding rate likely decreases because the pressure differential, which increases the molecule velocity, also serves to minimize the residence time of the molecule within the aperture, thus decreasing its likelihood of binding. It is possible that the applied pressure differential is also altering the on/off kinetics of the antigen to its antibody, thus decreasing the capture efficiency and/or increasing the off rate. Similar phenomena have previously been reported for studies in which fluid flow has been used to investigate the interaction kinetics and affinity properties of intermolecular recognition events.35−39 For example, cell adhesion properties are frequently studied as a function of shear rate (i.e., the rate of change of velocity from the surface) to probe vascular system behaviors with a simplified, in vitro model. In these adhesion experiments, cells are passed over a functionalized surface (coated with antibody, aptamer, or biomolecule such as fibronectin or collagen) in a fluidic chamber,35−38 probing

nM cSNAP-25 is shown in the absence of an applied pressure differential or at 0 mmHg (solid black diamonds) and under the influence of a −90 mmHg pressure differential (blue diamonds). With a −90 mmHg pressure differential, 500 nM cSNAP-25 is detected ∼8 times faster relative to its detection at 0 mmHg. The steepness coefficient (B) of the best fit for the representative data shown in Figure 3 is 4 and 33 for the 0 mmHg and −90 mmHg data sets, respectively. This result quickly demonstrates the general concept of using an applied pressure differential to speed up the rate of detection of the ICR biosensing measurement. However, it should be noted that while speeding up the rate of detection of the target analyte, an applied pressure differential also alters the transport of ions to and through the aperture of the AMGNM with the resulting volumetric flow and thus also alters the absolute ICR response. When a negative pressure differential is applied across a conical shaped nanopore, relative to the external solution, the induced volumetric flow transports ions from the external bulk solution into the aperture. This increased solution flow, and ion flux, reduces the ion accumulation and depletion effects measured at the aperture, resulting in a less rectified ICR response, as previously described by Lan et al.27 This effect was observed for the AMGNM, as illustrated in the Supporting Information (Figure S2). However, even under pressure, the change in the ICR of the AMGNM upon cSNAP-25 binding is still resolvable (as shown in the Supporting Information, Figure S3). Thus, the fact that pressure serves to change the initial baseline ICR of the AMGNM is additional justification for normalizing the data, as described by eq 1. To further assess the utilization of pressure as a means of increasing the rate of cSNAP-25 detection at the AMGNM, we examined a wider range of pressure differentials. Figure 4 depicts the steepness coefficient (B from eq 2) as a function of

Figure 4. (A) Plot of the steepness coefficient, B, as a function of cSNAP-25 concentration (from 1 pM to 100 μM) over an applied pressure differential range of 0 mmHg to −270 mmHg. (B) Plot of the scaling coefficients, A, as a function of the applied pressure differential (from 0 mmHg to −270 mmHg) for the detection of cSNAP-25 using the AMGNM. Data were taken with 50−250 nm radius AMGNMs. E

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Analytical Chemistry specific interactions that lead to cell capture/adhesion. From the resulting optimal shear rate (i.e., rate at which the capture efficiency is strongest), differences in binding affinity can be assessed. Additionally, solution flow is an integral part of surface plasmon resonance-based techniques. In surface plasmon resonance experiments, fluid is passed through the chamber and a target analyte is captured on a surface functionalized with a receptor molecule. To achieve efficient analyte capture, the flow must be fast enough to minimize mass transport limitations (i.e., the formation of analyte depletion regions)40 yet slow enough to maximize capture efficiency.39 Generally, for analyte interactions at a surface under solution flow conditions, there exists an optimal linear velocity that maximizes the encounter rate of the target analyte and interaction time (i.e., residence time) required for binding and subsequent detection.38,41 Here, for the first time, we have utilized fluid velocity as a means of controlling and enhancing the rate of detection of a specific target analyte within a nanopore. Lan et al. have previously characterized the pressuredependent solution flow profiles within conical nanopores,27,42 which are used here to describe the pressure-dependent antigen detection behavior we observe. Within our conically shaped AMGNMs, the volumetric flow (Q) can be estimated by27,42 Q = 3πr 3ΔP /8η cot θ

The molecular transport rate when characterizing 1 pM cSNAP-25 thus increases to ∼1.28 × 101 molecules/s when utilizing a −210 mmHg pressure differential, an ∼350-fold increase relative to 0 mmHg (υavg = 0). While the application of increasing pressure differentials will continue to increase the cSNAP-25 transport rate to the aperture, at −210 mmHg the optimal detection rate is experimentally achieved, likely as a result of balancing molecule delivery, transport velocity, and capture efficiency. Assuming a sensing zone length within the AMGNM of ∼12.5 μm (the distance into the 50 μm length pore at which the electric field is reduced 99.5% from its maximum value, as estimated from Figure S4 in the Supporting Information), once cSNAP-25 enters the AMGNM, it will reside within this sensing region for a minimum of ∼42 μs at −210 mmHg, based on dividing the length of the sensing region by υavg (∼0.3 m/s), before the molecule can no longer bind and be detected at the aperture. However, once the molecule has entered the aperture, it can traverse its way to the walls of the aperture via its diffusive motion, bind to its antibody, and be detected while it still resides in the sensing zone. The time that it takes the molecule to traverse its way from the center of the AMGNM aperture to the pore wall (ttrav) can be estimated from a threedimensional random walk:46 t trav = r 2/6D

(3)

where ttrav is the diffusion time, and r is the radius of the aperture. Given this simplified model, it will take ∼37 μs for the molecule to reach the aperture surface, which leaves enough time for the molecule to bind and be detected at −210 mmHg. It should be noted that this description is only an approximation and does not take into account the angle of the pore, the on/off kinetics of the binding reaction, the fact that the random motion of the molecule increases its residence time within the sensing zone, and the fact that all the molecules do not start at the center of the aperture. However, it does highlight that there will be a pressure at which the molecule residence time is too brief for detection. At a pressure differential of approximately −225 mmHg, the molecule residence time will equal the diffusion time (∼37 μs); therefore, at pressures greater than this value, molecules entering at the aperture center (a distance r from the pore wall) will not have time to traverse and bind, reducing the overall binding efficiency at the AMGNM. These approximations support the idea that the applied pressure differential can be optimized in pressure-dependent ICR biosensing studies to achieve maximum capture and detection for a given binding molecule-target analyte pair. A demonstration of the ability to characterize low concentrations of cSNAP-25 using the optimal pressure differential of −210 mmHg, as determined from Figure 4B, is shown in Figure 5. Here, the detection of cSNAP-25 at concentrations of 0 pM, 1 pM, 5 pM, and 10 pM are shown for comparison. As would be expected, the binding of 10 pM cSNAP-25 occurs the fastest with B = 8, and 1 pM binding occurs the slowest with B = 5. Simply comparing the detection times of picomolar concentrations of cSNAP-25 (collected at −210 mmHg) in Figure 5 relative to the detection times of micromolar concentrations of cSNAP-25 (collected at 0 mmHg) in Figure 2B demonstrates the ability to detect cSNAP-25 within the same time frame at concentrations at least 6 orders of magnitude lower with the use of an optimal pressure differential as compared to without.

where r is the aperture radius, ΔP is the pressure differential in Pa, η is the solution viscosity (9 × 10−4 Pa s),27,43 and θ is the half cone angle of the pore (10°).21,42 For a 150 nm radius AMGNM (the intermediate radius size utilized in this study), assuming the pressure drop occurs primarily at the aperture, the volumetric flow at −210 mmHg (the optimal pressure differential for cSNAP-25 detection at the AMGNM) will be ∼2 × 10−8 cm3/s (or ∼2 × 10−5 μL/s). The average solution velocity (υavg) at the orifice can then be calculated by dividing the volumetric flow rate Q by the cross-sectional area of the aperture (υavg = Q/πr2),42 yielding ∼0.3 m/s. From this velocity, we can estimate and compare the molecular transport rate (I) through the aperture of the AMGNM at −210 mmHg relative to a diffusion limited system (i.e., 0 mmHg). On the the assumptions that the cSNAP-25 detection process is diffusion limited at 0 mmHg, that the aperture of the AMGNM is a perfect sink, and that there is zero diffusion along the exterior surface of the AMGNM, the diffusion limited molecular transport rate through the AMGNM (Idiffusion), which is a function of the flux and cylindrical geometry of the aperture, can be estimated by44−46 Idiffusion = 4DrC

(4)

where D is the bulk diffusion coefficient of the antigen, r is the radius of the aperture, and C is the bulk concentration of the antigen. By approximating the cSNAP-25 diffusion coefficient as ∼10−6 cm2/s,47 we can estimate the number of molecules entering a 150 nm radius aperture for an antigen concentration of 1 pM cSNAP-25 at 0 mmHg to be ∼3.61 × 10−2 molecules/ s. However, when a pressure differential of −210 mmHg is applied, the molecular transport rate through the AMGNM is increased by an independent convective component (i.e., a net cross-flow), yielding the new molecular transport rate:46 Idiffusion+convection = 4DrC + υavg C(πr 2)

(5) F

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ASSOCIATED CONTENT

S Supporting Information *

Detailed description of the five-parameter nonlinear regression model, pressure-based ICR responses, simulated electric field profile, and AMGNM interfering agent species characterization plot. The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acs.analchem.5b00757.

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AUTHOR INFORMATION

Corresponding Author

*E-mail: [email protected].

Figure 5. Representative binding curves for the % AMGNM saturation as a function of time for 1 pM cSNAP-25 (red circles), 5.0 pM cSNAP25 (blue squares), and 10 pM cSNAP-25 (black triangles), detected under an applied pressure of −210 mmHg. The 0% AMGNM saturation baseline value at −210 mmHg (i.e., the response of the AMGNM in the absence of the antigen) is shown for comparison (green diamonds). Data were taken with 50−250 nm radius AMGNMs.

Author Contributions

The authors declare no competing financial interest. Notes

The authors declare no competing financial interest.

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ACKNOWLEDGMENTS This work was supported by NIH Grant 1R43NS074610 and the CBD Program W81XWH-11-C-0085.

After determining −210 mmHg to be the optimal pressure differential for cSNAP-25 detection using the AMGNM, we also assessed the selectivity of the anti-cSNAP-25 coated AMGNM ICR biosensing measurement. Specifically, the response of the AMGNM (utilizing a pressure differential of −210 mmHg) was tested against uncleaved synaptosomalassociated protein 25, botulinum toxin type A light chain, monoclonal antibody specific to the botulinum toxin type A light chain, and tris(2-carboxyethyl)phosphine, species that would be typically encountered during botulinum characterization studies, none of which elicited a detection response (as shown in the Supporting Information, Figure S5). While we recognize that more complex solution conditions and interfering agents may nonspecifically bind to the AMGNM and that analogue peptides or proteins containing the cSNAP25 sequence and terminal −COOH (as generated from botulinum cleavage) are also likely to have affinity for the antibodies utilized, under the conditions studied to date, the AMGNM is specific to cSNAP-25.

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CONCLUSIONS The results presented here demonstrate the ability to utilize a mechanically applied pressure differential to manipulate and control the detection of a target analyte in an ICR biosensing measurement. By mapping out the dependence of the rate of analyte detection as a function of the applied pressure differential, we were able to determine the optimal pressure for detecting cSNAP-25 utilizing an AMGNM, i.e., the pressure that enables cSNAP-25 detection at the fastest possible rate. Utilizing this optimal pressure differential allowed us to characterize picomolar cSNAP-25 concentrations, 106 orders of magnitude lower than can be achieved without the use of a convective force. We speculate that this optimal pressure represents a balance between convective molecular transport, molecular diffusion, and binding of the target antigen to the AMGNM surface. The methodology presented significantly expands the applications of nanopore ICR biosensing measurements and demonstrates that these measurements can be used for the quick detection of target analytes at picomolar levels, a current and pressing challenge in quantitative and specific analyte detection.1,2 G

DOI: 10.1021/acs.analchem.5b00757 Anal. Chem. XXXX, XXX, XXX−XXX

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DOI: 10.1021/acs.analchem.5b00757 Anal. Chem. XXXX, XXX, XXX−XXX