GB2622574A

GB2622574A - Microfluidic cell

Info

Publication number: GB2622574A
Application number: GB2213114.8A
Authority: GB
Inventors: Alexander Warren Michael; Hallam Toby
Original assignee: Newcastle University of Upon Tyne
Current assignee: Newcastle University of Upon Tyne
Priority date: 2022-09-08
Filing date: 2022-09-08
Publication date: 2024-03-27
Also published as: GB202213114D0; WO2024052697A1

Abstract

A microfluidic device comprising a fluid channel 102 with a longitudinal axis along a fluid flow direction, a first electrode 106 comprising a first aperture 108 and a first electrode surface 110, and a second electrode 112 comprising a second aperture 114 and a second electrode surface 116 which is coplanar with the first electrode surface wherein the fluid channel extends through the first and second apertures perpendicular to the first and second electrode surfaces such that the electrodes surround or wrap around the channel. The positioning of the electrodes provides a more uniform electric field. The device may further comprise a third electrode (fig 1b). A method of determining a characteristic of a particle such as a biological particle using the device comprises generating an electric field between the first and second electrodes by application of an alternating current, passing a fluid containing the particle along the channel, determining a current variation in the channel in dependence of the particle moving through the electrical filed and determining a parameter of the particle. Ideally, particle size is determined.

Description

MICROFLUIDIC CELL

TECHNICAL FIELD

The present disclosure relates to devices and methods for analysis of particles, such as biological particles, in particular using microfluidic cells. Examples relate to devices for microfluidic analysis including novel electrode arrangements.

BACKGROUND

Microfluidic devices based on micro-electromechanical technology may be used in biological and medical applications, for example. A microfluidic device is an instrument capable of controlling small amounts of fluid moving through one or more small channels e.g. the channel may have a diameter in the sub-millimetre range. Microfluidic devices can be built on microchips which are sometimes called a lab-on-a-chip. Microfluidic devices can be used to detect and analyse particulate samples. Body fluids or cell-containing solutions may be analysed to diagnose diseases. Biological entities such as cells, cell parts, and macromolecules, may be analysed using microfluidics based on their electrical and magnetic responses, for example.

In the field of biological measurements, techniques which allow label-free single cell analysis are being developed which can provide information that would otherwise be difficult to obtain using bulk measurement methods. Microfluidic Impedance Cytometry (MIC) has been developed as a method which is capable of carrying out rapid, single cell measurements in order to allow diagnostic tests to be carried out in challenging or resource-scarce environments such as conflict regions or developing countries. Disposable microfluidic cartridges and built-in signal conditioning equipment can carry out these measurements without the use of a static lab or ancillary equipment. Such approaches hold the potential for small, portable devices which can be used for other research applications. There is also possibility for these devices to be utilised in industrial environments, such as in the testing of lubricant or detect degradation on key machinery components as real time condition monitoring devices. Such functionality can be key for equipment such as back-up diesel engines or gas turbines which are often used on nuclear sites to protect essential supplies.

The MIC technique makes use of alternating current (AC) excited electrodes embedded into a microfluidic channel that define a sensing region. Current can flow between the electrodes, the amplitude of which varies, which is due to a large extent to the dielectric properties of the material within the sensing region. A sample containing particles or cells may be suspended in a fluid solution which can then be passed through the microfluidic channel. As each cell passes through the sensing region, a change in current can be measured due to the variation in material properties i.e. electrical impedance. By measuring this signal variation, determinations about the fluid and sample can be made; for example, cell counting and sizing can be performed. Some key challenges remain such as the sensitivity of such devices to cell trajectory within the sensing region, which can subsequently result in errors in the sizing of measured cells.

The development of MIC devices has led to a strong focus on three main designs. These are i) parallel electrodes whereby a pair of electrodes are positioned opposite from each other with the microfluidic channel between them, fi) coplanar electrodes whereby a pair of electrodes are located side by side together along the length of a microfluidic channel, and iii) designs using a constriction channel where the size of the microfluidic channel is narrowed at one or more points along its length. One of the key issues of these devices is the inhomogeneity of the electric field in the sensing region, which can induce errors because cell position and size information becomes conflated. A number of studies have attempted to overcome the issue of non-homogeneous electric fields in the sensing region through the use of novel designs and signal processing techniques; however, these modifications often result in complex manufacturing processes, heavy post-processing and in some cases, reduced sensitivity.

It may be advantageous to be able to address such issues. It is an aim of the present disclosure to address one or more of the disadvantages associated with the prior art.

SUMMARY OF THE INVENTION

In an aspect there is provided a microfluidic device, comprising: a fluid channel comprising a longitudinal axis along a fluid flow direction; a first electrode comprising a first aperture and a first electrode surface; and a second electrode comprising a second aperture and a second electrode surface coplanar with the first electrode surface; wherein the fluid channel extends through the first and second apertures perpendicular to the first and second electrode surfaces. The first electrode may be configured to have an alternating voltage applied thereto.

The second electrode may be configured to be connected to ground. The fluid channel may have a constant cross section.

The first electrode surface and the second electrode surface may each have a separation therebetween of between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 100 micrometers; more preferably between 20 micrometers and 40 micrometers.

The first electrode surface and the second electrode surface may each be configured to be capacitively coupled on application to an alternating voltage to the first electrode and the second electrode being grounded.

The first electrode and the second electrode may each be configured to be in contact with a fluid passing along the fluid channel.

The first aperture and the second aperture may each have the same aperture diameter. A separation ratio, of the aperture diameter to the perpendicular distance between the first electrode surface and the second electrode surface, may be between 0.1 and 10; preferably between 0.5 and 1.5.

The fluid channel may have a channel width between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 200 micrometers; more preferably between 20 micrometers and 50 micrometers.

The fluid channel may have a circular cross section. Such a fluid channel may have a diameter between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 200 micrometers; more preferably between 20 micrometers and 50 micrometers. The fluid channel may have a geometrical shape cross section. The fluid channel may have a rectangular or square cross section.

A first electrode ratio, of an area of the first electrode surface to the area of the first aperture, may be 2 or above. A second electrode ratio, of an area of the second electrode surface to the area of the second aperture, may be 2 or above. The first aperture may be circular. The second aperture may be circular. The first aperture and second aperture may have substantially the same diameter. One or more of the first electrode and the second electrode may be conductive plates.

One or more of the first electrode and the second electrode may comprise at least one of: graphene, a metallic solid film, a metallic conductive polymer; and a metallic nanoparticle film. The microfluidic device may further comprise a non-conductive polymeric layer between the first electrode surface and the second electrode surface.

The first electrode may be located in a first direction along the longitudinal axis away from the second electrode, the first electrode surface facing the second electrode surface. The second electrode may be a second electrode plate comprising the second electrode surface and an opposite second electrode surface coplanar with and opposite to the second electrode surface, the second electrode plate configured to be connected to ground. The microfluidic device may further comprise a third electrode located in a second direction opposite the first direction along the longitudinal axis, the third electrode comprising a third aperture and a third electrode surface coplanar with and facing the opposite second electrode surface, the second electrode plate thereby located between the first electrode surface and third electrode surface. In response to a first alternating voltage applied to the first electrode, a first electric field may be generated between the first electrode surface and the second electrode surface, and in response to a second alternating voltage applied to the third electrode, a second electric field may be generated between the opposite second electrode surface and the third electrode surface.

The microfluidic device may further comprise a first non-conductive polymeric layer between the first electrode surface and the second electrode surface, and a second non-conductive polymeric layer between the opposite second electrode surface and the third electrode surface.

The first alternating voltage may be substantially equal in magnitude and opposite in polarity to the second alternating voltage.

The separation between the first electrode surface and the second electrode surface may be substantially the same as the separation between the opposite second electrode surface and the third electrode surface.

The microfluidic device may further comprise a voltage source configured to apply an alternating voltage to the first electrode with the second electrode connected to ground, thereby generating an electric field between the first electrode surface and the second electrode surface.

The microfluidic device may further comprise a voltage source configured to apply a first alternating voltage to the first electrode with the second electrode connected to ground, thereby generating a first electric field between the first electrode surface and the second electrode surface; and a second alternating voltage to the third electrode with the second electrode connected to ground, thereby generating a second electric field between the opposite second electrode surface and the third electrode surface.

The alternating voltage applied to the first electrode may be a first alternating voltage. The first alternating voltage may be substantially equal in magnitude and opposite in polarity to the second alternating voltage The voltage source may be configured to apply an alternating voltage between 0.1V AC and 50V AC; preferably between 0.5V AC and 10V AC; more preferably approximately 1V AC. The voltage source may be configured to apply an alternating voltage of a frequency between 10 kHz and 50 MHz; preferably between 250 kHz and 2 MHz; more preferably between 500 kHz and 1 MHz. The voltage source may be configured to apply a plurality of alternating voltages, each alternating voltage of the plurality being of a different frequency. The plurality of alternating voltages may be applied in a frequency sweep from a lower frequency to an upper frequency, and/or vice versa, in some examples. The plurality of alternating voltages may be applied simultaneously in some examples.

In an aspect there is provided a method of determining a characteristic of a particle using a microfluidic device as disclosed herein, the method comprising: applying an alternating voltage to the first electrode and to the second electrode to thereby generate an electric field between the first electrode surface and the second electrode surface; passing a fluid containing the particle along the fluid channel in the fluid flow direction; determining a current variation in the fluid channel in dependence on the particle moving through the electric field; and determining a parameter of the particle in dependence on the current variation. Determining the parameter of the particle may comprise determining a size of the particle.

Within the scope of this application it is expressly intended that the various aspects, embodiments, examples and alternatives set out in the preceding paragraphs, in the claims and/or in the following description and drawings, and in particular the individual features thereof, may be taken independently or in any combination. That is, all embodiments and/or features of any embodiment can be combined in any way and/or combination, unless such features are incompatible. The applicant reserves the right to change any originally filed claim or file any new claim accordingly, including the right to amend any originally filed claim to depend from and/or incorporate any feature of any other claim although not originally claimed in that manner.

BRIEF DESCRIPTION OF THE DRAWINGS

One or more examples will now be described, by way of example only, with reference to the accompanying drawings, in which: Figures 1A and 1B show example microfluidic channels and electrodes according to examples disclosed herein; Figures 2A-2F show example geometries of coplanar (A), parallel (C) and new "sandwich" (E) electrodes as disclosed herein, with corresponding modelled electric field lines, according to examples disclosed herein; Figures 3A-3D show two-electrode parametric design studies indicating (A) sensitivity to a 10pm Red Blood Cell (RBC), (B), (C) variable device dimensions and (D) associated design optimisation results, according to examples disclosed herein; Figures 4A-4F show results from a two-electrode device accuracy study according to examples disclosed herein -Figures 4A, 48, 4C illustrate histograms of frequency distributions of measured responses of coplanar, parallel and sandwich device designs, respectively, and box and whisker diagrams of the same data is displayed in Figures 4D, 4E and 4F; Figures 5A-5F show an example disclosed herein comprising a grounded central electrode with differential current measurement. Figure 5A is a schematic representation of an example device, Figure 5B shows a plot of device signal response throughout device optimisation studies, Figure 5C shows initial variation in measured signal due to cell position variation and Figure 5D shows initial variation in potential due to a 10pm cell passing through the sensing region. Figure 5E shows improvement in measured signal variation achieved using an optimised device while Figure 5F indicates the accuracy of a device as disclosed herein is improved; and Figure 6 shows an example method according to examples disclosed herein.

DETAILED DESCRIPTION

Microfluidic Impedance Cytometry (MIC) has been developed as a method which is capable of carrying out rapid, single cell measurements. MIC uses alternating current (AC) excited electrodes embedded into a microfluidic channel that define a sensing region. Current can flow between the electrodes, the amplitude of which varies, which is due to a large extent to the dielectric properties of the material within the sensing region. A sample containing particles or cells may be suspended in a fluid solution which can then be passed through the microfluidic channel. As each cell passes through the sensing region, a change in current can be measured due to the variation in material properties i.e. electrical impedance. By measuring this signal variation, determinations about the fluid and sample can be made.

The development of MIC devices has led to a strong focus on three main designs -parallel electrodes, coplanar electrodes, and designs using a constriction channel. Inhomogeneity of the electric field in the sensing region of these designs can induce errors because cell position and size information becomes conflated. Modifications to try and address the errors often result in complex manufacturing processes, heavy post-processing and in some cases, reduced sensitivity.

Examples disclosed herein aim to address the shortcomings in the art by using a new MIC electrode arrangement which provides a more uniform electric field in the sensing region. Examples disclosed herein comprise a novel M IC design based on a sandwich of electrodes which can address weaknesses of known parallel and coplanar designs, by creating a homogenous field distribution throughout the sensing region. Examples disclosed herein can achieve a high sensitivity (for example, defined as percentage change in measured signal) while decreasing the induced error by offset cell trajectory in the sensing region, without the requirement for complex alignment procedures. Examples have been quantified as discussed below, by comparing the new electrode designs with known designs through studies involving numerical simulations. These have been carried out without additional analysis techniques such as cell focussing and signal diagnostics, which are typically used to enhance the sensitivity and selectivity of these devices, but can also reduce their effectiveness as a point of care device.

Figure la shows an example of a microfluidic device 100 comprising a microfluidic channel 102 with two electrodes 106, 112. Fluid can flow in at the inlet 104 of the channel 102, through an electric field generated by the electrodes 106, 112 in the channel portion 120 between the electrodes 106, 112, and flow out of the channel 102 at the outlet 105. Figure lb shows an example of a microfluidic device 100 comprising a microfluidic channel 102 with three electrodes 106, 112, 122. Fluid can flow in at the inlet 104 of the channel 102, through an electric field generated by the first two electrodes 106, 112 in the first channel portion 120a, through an electric field generated by the second two electrodes 112, 122 in the second channel portion 120b, and out of the channel 102 at the outlet 105.

Both examples are of a microfluidic device 100 comprising: a fluid channel 102 comprising a longitudinal axis along a fluid flow direction; a first electrode 106 comprising a first aperture 108 and a first electrode surface 110; and a second electrode 112 comprising a second aperture 114 and a second electrode surface 116 coplanar with the first electrode surface 110; wherein the fluid channel 102 extends through the first and second apertures 108, 114 perpendicular to the first and second electrode surfaces 110, 116. The first electrode 106 may be configured to have an alternating voltage applied thereto. The second electrode 112 may be configured to be connected to ground. The fluid channel 102 may have a constant cross section as shown.

The example of Figure 1B may be called a three-electrode sandwich design (see also Figures 5A-5F), and comprises a first electrode 106 which is located in a first direction along the longitudinal axis away from the second electrode 112. The first electrode surface 110 is facing the second electrode surface 116. The second electrode 112 is a second electrode plate 112 comprising the second electrode surface 116 and an opposite second electrode surface 118 coplanar with and opposite to the second electrode surface 116. The second electrode plate 112 is configured to be connected to ground. The microfluidic device 100 in this example comprises a third electrode 122 located in a second direction opposite the first direction along the longitudinal axis. The third electrode 122 comprises a third aperture 124 and a third electrode surface 126 coplanar with and facing the opposite second electrode surface 118. The second electrode plate 112 is thereby located between the first electrode surface 110 and third electrode surface 126.

The microfluidic device 100 may further comprise a voltage source (not shown) configured to apply an alternating voltage to the first electrode 106 with the second electrode 112 connected to ground, thereby generating an electric field between the first electrode surface 110 and the second electrode surface 116. In examples with a third electrode as in Figure 1B, the microfluidic device 100 may comprise a further voltage source (not shown), or may use the same voltage source, which is configured to apply a second alternating voltage to the third electrode 122 with the second electrode 112 connected to ground, thereby generating a second electric field between the opposite second electrode surface 118 and the third electrode surface 126 Thus, in response to a first alternating voltage applied to the first electrode 106, a first electric field may be generated between the first electrode surface 110 and the second electrode surface 116. In response to a second alternating voltage applied to the third electrode 122, a second electric field may be generated between the opposite second electrode surface 118 and the third electrode surface 126. The first alternating voltage may be substantially equal in magnitude and opposite in polarity to the second alternating voltage.

The voltage source or sources may (each) be configured to apply an alternating voltage between 0.1V AC and 50V AC; preferably between 0.5V AC and 10V AC; more preferably approximately 1V AC. The voltage source or sources may (each) be configured to apply an alternating voltage of a frequency between 10 kHz and 50 MHz; preferably between 250 kHz and 2 MHz; more preferably between 500 kHz and 1 MHz. The voltage source or sources may (each) be configured to apply a plurality of alternating voltages, each alternating voltage of the plurality being of a different frequency. The plurality of alternating voltages may be applied in a voltage sweep from a lower frequency to an upper frequency, and/or vice versa.

In some examples, the microfluidic device 100 may further comprising a first non-conductive polymeric layer between the first electrode surface 110 and the second electrode surface 116.

In examples with a third electrode as in Figure 1B, there may be a second non-conductive polymeric layer between the opposite second electrode surface 118 and the third electrode surface 126. The separation between the first electrode surface 110 and the second electrode surface 116 may be substantially the same as the separation between the opposite second electrode surface 118 and the third electrode surface 126.

Figures 2A-2F show examples of different electrode geometries 140, 160, 100 and corresponding electric field line distributions 202, 222, 242. The arrows 142, 162, 182 marked indicate the inlet fluid flow direction along the x axis. In order to appropriately assess the performance of the sandwich electrode device 100 and provide a suitable comparison to the existing coplanar 140 and parallel electrode 160 designs, models have been developed and in silico experiments have been performed using the commercial finite element software COMSOL Multiphysics®. The geometries of the parallel, coplanar and sandwich electrode device design can be observed in Figures 2A, 2C and 2E respectively. From these models, analysis of electric field line plots 202, 222, 242 has been carried out, as shown in Figures 2B, and 2F, to examine variations between the three electrode geometry designs. The electric field lines shown in Figures 2B, 2D and 2F are shown as viewed along the y-axis as shown (i.e. in the x-z plane).

As considerations when creating the models, as the sensing mechanism for the M IC devices 100, 140, 160 consists of a current flow through the sensing region 204; 224; 244 which is affected by the dielectric properties within the domain, the associated AC/DC module has been implemented in order to solve Maxwell's equations for the electric field, current and potential distributions within the modelled domains. Quadratic tetrahedral elements have been adopted and the mesh designed to create a higher element density in regions of higher electric field gradient. The variation in current flow between excitation and sensing electrodes 146, 152; 166, 172; 106, 112 due to the variation in conductivity and relative permittivity in the sensing region 204; 224; 244 has been modelled based on steady state calculations using a parametric sweep of the positions of the particle initially released at the centre of the entry plane and travelling along a direct trajectory. The meshing of the FEM model was carried out using a quadratic tetrahedral mesh in all domains and in order to ensure the mesh was sufficiently refined, mesh convergence studies were carried out until the readings converged to indicate minimal residual error.

In order to model a representative biological cell passing through the devices 140, 160, 100, the cells have been modelled as spherical geometries within the device, while the cell membrane has been modelled by a thin layer electric shielding condition, specified at all outer surfaces of the cell. 1 MHz was chosen for the simulations as an ideal frequency at which to measure cell sizing data due to the type of dielectric response expected from cells. Impedance measurements of cells suspended in solution are frequency dependent and go through characteristic dispersions. At low frequencies (<100 kHz), these devices are known to exhibit a largely capacitive effect due to the formation of an Electric Double Layer (EDL) at the interface between the electrodes and the solution which acts as a large source of capacitance in the measured signal, resulting in a large phase shift.

The magnitude of the EDL can be reduced by operation of the device in sufficiently high frequencies, therefore 1 MHz was selected as a suitable frequency to allow the magnitude of the EDL to be ignored, while avoiding the high frequency cytoplasm polarization regions, which while rich in information is not related to cell size. In order to benchmark the sandwich design against coplanar and parallel electrode designs, a Finite Element Model for a coplanar, parallel and the sandwich electrode design was created and frequency domain analysis carried out at each desired position of the subject cell within the device. For these simulations Dirichlet boundary conditions were implemented at the surface of the excitation electrodes, with an excitation voltage of 1V AC at a frequency of 1 MHz; For the corresponding sensing electrode, a Dirichlet boundary condition was also implemented on its surface with a specified potential voltage of DV AC (ground). The resultant flow of current through the grounded sensing electrode is then measured, which varies as a particle or cell passes through the sensing region 204; 224; 244 due to the variation in electrical conductivity and dielectric constant in the domain.

The sizing of the microfluidic channel and the sensing region 204; 224; 244 (Figures 2B, 2D and 2F) is a compromise between throughput and the sensitivity of the devices, as this is governed by the dimensions of the sensing volume. Hence cross-sectional dimensions of the microfluidic channel can be increased in order to reduce the chances of clogging but will suffer from a reduction in sensitivity. Due to this, it is recognised that each application will have individual design requirements which will be based on the diameter of the cells to be measured and may also depend on other cells present in the sample which require counting.

The dimensions of these devices 140, 160, 100 were selected based on potential measurement of whole blood cell counts, therefore a minimum sensing region 204; 224; 244 dimension of 30 pm has been specified in order to accommodate Plasma, Erythrocyte (Red Blood Cells), Leukocytes (White Blood Cells) and Platelets with an additional margin for enlarged, cancerous cells. To compare the particle sizing ability of these designs, a sample data set of cell trajectories for all designs has been created by using a simplified device geometry and application of a Computational Fluid Dynamics module of COMSOL in order to analytically model the flow of phosphate buffer solution at a flow rate of 400 pm/s through each device (the Computational Fluid Dynamics module solves for a single phase fluid flow using Navier-Stokes and continuity equations). To ensure accuracy of these results, the mesh for these models utilised 8 boundary layers on all inner surface boundaries in order to more accurately resolve the fluid flow gradient at the no slip boundaries. These equations were solved for the steady state conditions present in a fully developed flow. Following this, the Particle Tracing module was then implemented in order to model the trajectory of particles through the device, based on random particle release positions at the inlet (the Particle Tracing module is another COMSOL module which is used to simulate particle motion and solve for forces acting upon, and hence displacement of, the particles. This allows particle diameters and mass distributions to be specified, and allows the forces upon them to be solved for. It also allows two-way coupled particle-fluid interactions to be solved for). A freeze wall boundary condition was then specified centrally between the electrodes (at the point of highest sensitivity within the device) and cell positions captured after a data set of approximately 900 cells was achieved. These cell positions have then been used within the parametric sweep studies using the stationary AC/DC solver in order to solve for the current flow within the devices at each cell position.

For the coplanar and parallel electrode geometries of Figures 2A-2B and 2C-2D respectively, the electrodes 146, 152; 166, 172 are strips lying with their long dimension in the y direction. In the electric field line distribution plots of Figures 2B and 2D, the electrodes are viewed from the electrode ends, looking along the strip length of each electrode in the y direction, perpendicular to the fluid flow direction along the x axis. For the sandwich electrode geometry of Figures 2E-2F, the electrodes 106, 112 are planar discs with the fluid channel located through the centres of the discs, with the disc planes oriented in the y-z plane, perpendicular to the direction of fluid flow along the x axis. In the electric field line distribution plot of Figure 2F the electrodes 106, 112 are viewed at the disc edges, looking parallel to the plane surfaces in the y direction, perpendicular to the fluid flow direction along the x axis.

Figure 2A illustrates a coplanar electrode geometry in which two electrodes 146, 152 are arranged side by side along the length of the microfluidic channel (i.e. arranged one after the other in the direction of flow of fluid 142 along the channel). Both electrodes 146, 152 are embedded into the bottom of the microfluidic channel, below and either side of the sensing region 204. The channel is presented here with a representative cross-section of 30pm height and 100pm width. Figure 2B shows the electric field lines 202 arising from the two coplanar electrodes 146, 152 of Figure 2A. Figures 2B, 2D and 2F showing electric field lines each show a representations of a 2D plane cut through the x-z plane of each illustrated device. As it can be seen in Figure 2B, the electric field 202 resulting from the electrode configuration in the coplanar device are nonhomogeneous and with a field intensity which decreases as one moves away from the electrodes 146, 152. The field strength variation at varying heights introduces inaccuracy (blurring) in the determination of cell size as identical cells will produce different signals, depending upon the position within the microfluidic channel. This is one of the key disadvantages in this design and has led to the widespread uptake of complex strategies for signal conditioning and cell focussing in order to achieve suitable results.

Figure 2C illustrates a parallel electrode geometry in which two electrodes 166, 172 are arranged facing each other on opposite sides of the microfluidic channel perpendicular to direction of flow of fluid 162 along the channel. Here two electrodes 166, 172 are implemented, one at the top and one at the bottom of the microfluidic channel, again presented with representative dimensions of 30pm height and 100pm width. Figure 2D shows the electric field lines 222 arising from the two parallel electrodes 166, 172 of Figure 2C. The microfluidic channels of Figures 2A and 2C in these examples have a rectangular microfluidic channel cross-section of 30pm height and 100pm width. An improvement in performance over the coplanar device of Figures 2A-2B is achieved for the parallel device shown in Figure 2C. The electric field lines 222 of the parallel device are shown in Figure 2D, and as the electrodes 166, 172 are located on opposite sides of the sensing region 224, the homogeneity of the electric field lines is increased over that of the coplanar device. However, there remains significant fringing of the electric field around the sensing region 224. It is also apparent that for electrodes with similar dimension to the coplanar device, the sensing region 224 is reduced in size. The reduction in sensing volume results in a higher sensitivity, but the cost of this improvement is that during the manufacturing process, the electrodes must be precisely aligned when they are fabricated on either side of the microfluidic channel. An additional effect of reducing the sensing region 224 in such a way, is that larger cells, which extend outside of the small sensing region 224 will interact with the non-homogeneous field and therefore reduce accuracy.

Figure 2E illustrates a sandwich electrode geometry as an example of a microfluidic device 100 as disclosed herein. In this electrode geometry, two electrode plates 106, 112 are arranged having the first electrode plate 106, comprising a first aperture and a first electrode surface, and a second electrode plate 112 comprising a second aperture and a second electrode surface parallel with the first electrode surface. The fluid channel comprises a longitudinal axis along the fluid flow direction 182 and extends through the first and second apertures of the first and second electrodes 106, 112 perpendicular to the first and second electrode surfaces.

Figure 2F shows the electric field lines 242 arising from the two sandwich electrodes 106, 112 of Figure 2E. The microfluidic channel in the example of Figure 2E is circular with a diameter of 30pm. As shown, the electrodes 106, 112 in this sandwich geometry completely surround the microfluidic channel, thereby forming a sandwich structure. In some examples, such electrodes 106, 112 may be embedded in a support matrix, such as a polydimethylsiloxane (PDMS) matrix. In this example, the channel has a circular 30pm diameter channel throughout.

Due to the electrode arrangement, the uniformity of the electric field is significantly improved over the examples of Figures 2A-2D, as shown in Figure 2F. The electrode 106, 112 edges extend away from the microfluidic channel, which removes the field fringing from the sensing region 244. This results in a much more homogeneous electric field distribution within the sensing region 244 with excellent alignment of the electric field lines, through the sensing region 244. In the studies discussed below, it is shown that this design results in high sensitivity and much greater cell sizing accuracy than that of the traditional coplanar and parallel devices. It is also shown that the cell sizing accuracy of this device 100 is much greater than that of the traditional designs.

The electric field lines 202, 222, 242 in Figures 2B, 2D and 2F indicate the overall behavioural tendencies of the three comparative MIC designs shown in Figures 2A, 20 and 2E. To quantify the difference in performance (i.e. the different in signal variation), a comparative 'single cell' sensing study has been carried out in order to directly compare the relative current variation for each electrode configuration of Figures 2A, 2C and 2E due to a 10 pm cell through each device The results are shown in Figures 3A -3D.

In the simulations, each device has been modelled as described in relation to Figures 2A-F above. The steady state electric current was solved for using a stationary study of a single 10pm cell at 100 points through the centre of the sensing region 204; 224; 244 to determine the maximum signal variation. To further explore the potential for increased sensitivity of sandwich electrode devices 100 with the variation of key geometrical parameters (such as the channel height 320), the effects of varying electrode gap 312; 314; 318 and channel dimension 316; 320 were investigated using a parametric study as shown in Figure 3D. The geometry was defined as indicated in Figures 38 and 3C, using a set of global parameters for electrode gap 312; 314; 318 and channel dimension 316; 320, allowing the device geometry to be altered and the model re-meshed in order to solve for varying devices and cell positions. The channel dimension 316 is illustrated for the parallel electrode configuration in Figure 33 (right), but also applies to the coplanar electrode configuration in Figure 3B (left).

The sensitivity has been determined by modelling the presence of a cell both outside and in the most sensitive position of the device sensing region 204; 224; 244, so the maximum unit perturbation is gathered as the sensing signal in each case. Fig. 3D demonstrates the sensitivity of all three electrode designs to variations in the device geometry such as electrode gap 312; 314; 318 (or width, in the case of the parallel electrode design) and channel dimension 316; 320 Figure 3A shows the sensitivity of each of the electrode arrangements in Figures 2A, 2C and 2E to a 10pm Red Blood Cell (RBC). It can be seen that over a width in the x-direction of approximately 50pm (i.e. in the direction of flow of the fluid along the fluid channel), a peak in real current variation as a percentage is seen of around 0.20% for a coplanar electrode arrangement 306, around 0.45% for a parallel electrode arrangement 304, and around 1.90% for the sandwich electrode arrangement 302. That is, the real current variation as measured using the sandwich electrode arrangement is around 9.5 times that measured using coplanar electrodes and around 4.2 times that measured using parallel electrodes.

Device dimensions are shown in Figure 33 (left, a coplanar arrangement as in Figure 2A and right, a parallel arrangement as in Figure 2C), and in Figure 3C (the sandwich arrangement as in Figure 2E). In the coplanar arrangement, the electrode gap 312 is defined as the separation between the two adjacent (or closest) edges of the two coplanar electrode strips. In the parallel arrangement, the electrode width 314 is defined as the perpendicular distance between the two long parallel edges of one of the electrode strips; both the electrodes in this arrangement have substantially the same width. In the sandwich arrangement of Figure 3C, the electrode gap 318 is defined as the perpendicular distance (or separation) between the two facing coplanar surfaces of the electrodes. The channel dimension 316 in the coplanar and parallel electrode examples is defined in this model as the height of the channel (in the coplanar electrode arrangement this is the dimension across the rectangular cross section channel with the electrodes at one side of the channel; in the parallel electrode arrangement this is the dimension across the channel between the electrodes, with an electrode at each side of the channel). The channel in the sandwich electrodes example is a circular cross section channel, and the channel dimension 320 is defined in this model as the channel diameter (which is the size of the aperture of each of the electrodes through which the channel passes).

Associated design optimisation results are shown in Figure 3D showing for each of a 20pm and a 30pm channel width 320 (for the sandwich electrode configuration) or each of a 20pm and a 30pm channel gap 316 (for the coplanar and parallel electrode configurations), the signal variation obtained for: a range of electrode widths 314 between 10pm and 100pm for the parallel electrode geometry of Figure 3B; a range of electrode gaps 312 between 10pm and 100pm for the coplanar electrode geometry of Figure 3B; a range of electrode gaps 318 between 10pm and 100pm for the sandwich electrode geometry of Figure 30. Note the log 10 scale of the y axis (signal variation % axis). Over the range of electrode gaps/widths, the sandwich electrode arrangement provides a signal variation of between around 1.08% and 6.8%, peaking at about 6.8% for an electrode gap of 15pm and a channel dimension of 20pm, and peaking at about 1.75% for an electrode gap of 40pm and a channel dimension of 30pm. This is higher than seen for the coplanar and parallel electrode configurations. Over the range of electrode gaps/widths, the coplanar electrode arrangement provides a signal variation of between around 0.2% and 0.75%, peaking at about 0.75% for an electrode gap of 10pm and a channel dimension of 20pm, and peaking at about 0.3% for an electrode gap of 30pm and a channel dimension of 30pm. Over the range of electrode gaps/widths, the parallel electrode arrangement provides a signal variation of between around 0.47% and 1%, peaking at about 1% for an electrode gap of 10pm and a channel dimension of 20pm, and peaking at about 0.42% for an electrode gap of 30pm and a channel dimension of 30pm.

The results of this study is that the signal variation obtained by a cell passing through the sandwich electrode device is larger than for the other two designs. It is important to highlight that the sensing volume is larger on the traditional designs (coplanar and parallel), with an effective cross-sectional area of 3000pm2 as opposed to 706pm2 that is present in the proposed sandwich electrode device. These simulations have been repeated for traditional devices with dimensions more comparable to the proposed sandwich electrode device, which demonstrate that the sandwich electrode device still shows an increased sensitivity.

It can be seen in Figure 3D that the sandwich electrode design achieves a far greater sensitivity in all configurations (i.e. various electrode gaps and channel dimensions), confirming that the increased uniformity in the electric field results in improved sensitivity. This may be of great benefit to device practicality due to a reduction in signal conditioning and analysis requirements in order to increase the signal to noise ratio.

The observed reduction in sensitivity for the sandwich electrode arrangement, as the electrode gap is reduced to the minimum value (below around 20pm to 40 pm depending on the channel diameter), may be indicative of the current density being focused in the region immediately between the edges of the electrodes. This may in turn result in the electric field not fully permeating the central part of the sensing region, resulting in lower sensitivity. A similar effect is known as field compression which occurs when the channel height is reduced.

Figures 4A, 4B and 4C illustrate histograms of frequency distributions of measured responses of the coplanar (Figure 4A), parallel (Figure 4B) and sandwich (Figure 40) electrode devices, respectively to a sample of different sized cells passing along the fluid channel. Box and whisker diagrams of the same data is displayed in Figures 4D, 4E and 4F. for the coplanar, parallel and sandwich electrode devices, respectively. The whiskers indicate the minimum to maximum range of results. These results illustrate the effect of the increased focussing of the electric field in the sensing region in the sandwich electrode design of Figures 4C and 4F compared with the coplanar (Figures 4A and 4D) and parallel (Figures 4B and 4E) electrode designs, which results in a clear reduction in the induced error of cell population size measurement (taken as the cube root of real impedance), based on variation of the cell trajectory within the sensing region.

As well as detecting cells, high accuracy MIC should be able to unambiguously determine the size and type of the cell. Within existing designs, a challenge to obtaining accurate measurements of cell size is the variation of signal amplitude due to varying cell position in the channel. This issue arises from the inhomogeneity of the electric field within the sensing region of the device where the magnitude of the signal variation from a cell passing through the sensing region is dependent upon the position of the cell within the channel as well as the size of the cell. For example, in the coplanar design (Figure 2A), a small cell passing near the electrodes would show a similar signal variation to a larger cell travelling further away from the electrodes.

The results shown in Figures 4A-4F are obtained by using a fluidic flow model along with a particle tracing model implemented in COMSOL Multiphysics® modelling software, to create a large (706 cells) randomised set of particle traces through the sensing region of the three different electrode configurations. This was used to represent a size distribution of real cells in a sample flow through these devices. This allows the interplay between cell size and position across the devices under study to be understood. Three different cell sizes were chosen (10 pm, 8 pm and 6 pm diameter cells) to be representative of blood cell species. These cell sizes have then been used to calculate and illustrate the variation in signal due to cell distribution in the device using the three investigated electrode configurations (coplanar, parallel and sandwich).

Each of the electrode designs were simulated for the sample of cells passing through and the peak signal variations are presented as histograms in Figures 4A, 4B and 4C. In an ideal case, the signal strength for each different cell size should be tightly clustered around a single value that scales linearly with cell volume. In the coplanar electrode design (Figures 4A and 4D), in which the field inhomogeneity is the largest, it can be observed that there is a considerable overlapping of measured values for differing cell sizes, which makes cell sizing information almost impossible to extract from the raw data. For the parallel electrode device (Figures 4B and 4E) the signal strengths are better distinguished from each other according to cell size, showing less overlap for signals from each cell size. It is easier to differentiate between cell sizes with improved fidelity using the parallel configuration than the coplanar configuration. However, of concern is the large signal variation for the larger cells from the parallel electrode configuration, which would make accurate sizing and detection of larger abnormal cells extremely difficult. The sandwich design proposed herein (Figures 40 and 4F) shows a large improvement, with clear discrimination between the cell population sizes, and with large separation of the population's signals according to cell size, indicating an increased accuracy over both of the traditional coplanar and parallel designs.

In order to quantify the improved separation of cell size groups for the sandwich electrode device, Brown-Forsythe Analysis of Variance (ANOVA) tests have been carried out to analyse the variance between measured groups of cell sizes for each device, while relaxing the assumption around equal variances across the datasets. The Brown-Forsythe (B-F) test is a statistical test for the equality of group variances based on performing an Analysis of Variance (ANOVA) on a transformation of the response variable. Wien a one-way ANOVA is performed, samples are assumed to have been drawn from distributions with equal variance.

If this assumption is not valid, the resulting B-F-test is invalid. A result of particular note in ANOVA analysis is the F ratio. This indicates the deviation from the null hypothesis (i.e., a value of 1 indicates that the data is sampled from data with the same mean, the higher the value, the more significantly separated the data, see Equation 1). The P ratio is analogous to the F ratio for a standard ANOVA, but with a variation in the calculation due to the use of the Brown-Forsythe ANOVA.

-Variance between groups Variance within group Equation (1) Device Design F* Value Coplanar 758.8 Parallel 5955 Sandwich 64655 Three Electrode 116802 Sandwich Table I -a summaty of ANOVA tests for cell sizing in MIC devices of Figures 4A-4F and Figures 5A-5F (three-electrode sandwich). ;It can be seen from the results in Table 1 that the P value for the coplanar device is the lowest of all designs, indicating the mean value for the population cell size shows much lower separation than for the sandwich design. This conflates to greater error in cell sizing in the coplanar electrode design. ;These results indicate that the sandwich device may be suitable for sizing cell populations far more accurately than by using traditional designs, reducing the requirement for additional cell focussing, the use of multi-frequency opacity measurements, or novel signal differentiation techniques. Although these techniques have been proven to be effective, they result in a more cumbersome and overall complex device. Thus, examples disclosed herein provide electrode configurations which are able to address these issues in the initial design stages, giving much higher accuracies through device electrode geometry alone. ;As indicated in Table 1, as well as the two-electrode sandwich design discussed above, a three-electrode sandwich design has also been conceived and modelled. A modification to the two-electrode sandwich design of Figures 1A, 2E and 30 is a three electrode sandwich model as shown as a schematic representation in Figure 5A. This may be used to increase the sensitivity of a sandwich electrode device by allowing a differential measurement to be carried out, with a signal of nominally zero when there are no cells passing through the sensing region. In this example as shown in Figure 5A, similarly to the example of Figure 1B, a grounded central electrode (second electrode) is located coplanar to and midway between a pair of outer electrodes (first and third electrodes). In other words, the central ground electrode is surrounded by two excited electrodes. This arrangement of electrodes allows for a differential current measurement to be made. ;Using a similar method to the cell sizing study for the two-terminal device, a parametric study for the electrode gap and the channel diameter was then carried out to determine optimal dimensions for this differential sandwich design. Figure 5A indicates the device dimensions that were varied, where the position of the outer electrode separation changes and the central electrode is kept at the midpoint. The channel diameter is also varied between 20pm and 40pm. ;Figure 5A indicates the parameters that have been adjusted during the studies. The electrode gap is defined as the separation of the facing surfaces of the two outer electrodes. The gap between the outermost electrodes is 30pm in this example. The channel diameter is the diameter across the cross section of the fluid channel co-planar to the electrode faces. The electrodes to either side of the central electrode define two separate field regions. In the example study of the three electrode sandwich design presented here, the two outer electrodes are excited at 1.0 Vac 1 MHz, consistent with the two terminal device studies shown in Figures 2E-2F and Figures 4C and 4F. The central electrode is grounded and the current passing through the outer excitation electrodes is evaluated. The resultant signal is measured as a differential signal (excitation electrode 1 real current -excitation electrode 2 real current) and results in a much higher sensitivity due to the use of this differential signal measurement. Because the gap between the outermost electrodes is 30pm, the reduction in the distance between the excitation and ground electrode results in a much higher gradient of electric field (6.53x105 Vim in initial device configuration) when compared to that of the two-electrode design (2.81x105 Vim in an optimised device configuration). This then results in a highly sensitive response in the resultant trace of a cell moving centrally through the sensing region, which can be seen in Figure 50. ;Figure 5B shows a plot of signal response of the device of Figure 5A at various electrode gaps between 10pm and 100pm and different channel widths as part of a device optimisation study. Figure 5B shows the signal variation obtained by comparing the signal for a cell outside of the sensing region and at the position within the channel which gives the highest response. As with the two-terminal device, sandwich electrode configuration achieves its highest sensitivities with the lowest sensing volume (channel diameter of 20pm). In contrast, the electrode separations illustrated in Figure 5B do not follow this trend and it is apparent that for smaller electrode gaps (<50pm), the device exhibits lower sensitivity. This is due to the differential design having two sensing regions between outer excitation electrodes, resulting in an electrode gap of at least 60pm being required between the outer electrodes in order to achieve maximum sensitivity (30pm per sensing region). ;Setting up the two separate field regions to either side of the central electrode allows for a symmetric electric field either side of the central grounded electrode to be set up. This results in a characteristic bipolar pulse being detected when a cell passes along the fluid channel past the electrodes, as shown in Figure 50. Figure 50 shows an initial variation in measured signal due to cell position variation in the channel. When compared to the two sandwich electrode devices discussed above, a differential device such as that of Figure 5A can yield higher sensitivity. Furthermore, the peak to peak time which can be obtained from the bipolar pulse plot of Figure 5C can be used to determine the cell velocity. ;Figure 5D shows an initial variation in potential due to a lOpm cell passing through the sensing region of the device of Figure 5A. It can also be observed in Figure 5D, that the simulations carried out using a population of cells in a fluidic flow demonstrate excellent separation and therefore hold promise for this device to achieve high accuracy upon implementation. Analysis of Variance testing has been carried out on this dataset and a F value of 3584 and IR2 value of 0.8569 was achieved for these results, indicating that the groups (cell sizes) are significantly separated. These values are comparable to those found in Table 1 for the parallel electrode device and show a poorer separation when compared to that of the two-electrode sandwich electrode device. ;Figure 5E shows an improvement in measured signal variation achieved using the device of Figure 5A, optimised according to the results shown in Figure 5B. Taking the optimal electrode separation of 80pm and a 30pm channel diameter (which is desired in this case in order to allow whole blood cell counts), the increased separation of the electrodes results in a significant increase in sensitivity of the device, seen in Figure 5E, while excellent separation of cell sizing signals is observed in Figure 5F, indicating high accuracy of the device of Figure 1B and 5A. Figure 5F indicates that the accuracy of the device for cell size discrimination has also been greatly improved folowing device dimension optimisation. Therefore, a differential current design as in Figure 5A may be used to obtain a model of a device having dimensions which allow for good cell size discrimiation. ;Brown-Forsythe Analysis of Variance testing has been carried out on this dataset and a F* value of 116802 was achieved for these results as shown in Table 1, again indicating that the groups (cell sizes) are significantly separated for the optimised, three electrode device of Figure 1B and 5A. This is a significant improvement over the two electrode design of Figure 1A. Optimization for three electrode devices in the coplanar configuration were performed and it was found that the sandwich electrode design is much more accurate for all designs.

Figure 6 shows an example method 600 of determining a characteristic of a particle using a microfluidic device as disclosed herein. The method 600 comprises: applying an alternating voltage to the first electrode and to the second electrode to thereby generate an electric field between the first electrode surface and the second electrode surface 602; passing a fluid containing the particle along the fluid channel in the fluid flow direction 604; determining a current variation in the fluid channel in dependence on the particle moving through the electric field 606; and determining a parameter of the particle in dependence on the current variation 608. Determining the parameter of the particle may comprise determining a size of the particle 610.

In summary, the performance of a new MIC device design having a sandwich electrode configured has been evaluated against that of the equivalent MIC device designs having coplanar or parallel electrodes using a finite element method. The results of the simulations demonstrate that the sandwich electrode device has an inherently higher sensitivity and is less susceptible to cell sizing measurement error due to variation of cell trajectories in the sensing region, which has been achieved by re-assessing the design from the ground up without the use of complex cell focussing or signal conditioning techniques. The electrode configuration for the sandwich design has been investigated in both a two-electrode and three-electrode configuration in order to demonstrate how this variation on MIC design can be implemented in more advanced devices to yield higher sensitivities. Beyond the improved performance, a significant benefit which should be highlighted is that of the device geometry, which would lend itself to simplified fabrication techniques rather than multi-step layer-by-layer microfabricafion methods associated with existing designs.

It will be appreciated that various changes and modifications can be made to the present disclosed examples without departing from the scope of the present application as defined by the appended claims. Whilst endeavouring in the foregoing specification to draw attention to those features believed to be of particular importance it should be understood that the Applicant claims protection in respect of any patentable feature or combination of features hereinbefore referred to and/or shown in the drawings whether or not particular emphasis has been placed thereon.

Claims

CLAIMS1. A microfluidic device, comprising: a fluid channel comprising a longitudinal axis along a fluid flow direction; a first electrode comprising a first aperture and a first electrode surface; and a second electrode comprising a second aperture and a second electrode surface coplanar with the first electrode surface; wherein the fluid channel extends through the first and second apertures perpendicular to the first and second electrode surfaces.
2. The microfluidic device of claim 1, wherein the fluid channel has a constant cross section.
3. The microfluidic device of any preceding claim, wherein the first electrode surface and the second electrode surface have a separation therebetween of between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 100 micrometers; more preferably between 20 micrometers and 40 micrometers.
4. The microfluidic device of any preceding claim, wherein the first electrode surface and the second electrode surface are configured to be capacitively coupled on application to an alternating voltage to the first electrode and the second electrode being grounded.
5. The microfluidic device of any preceding claim, wherein the first electrode and second electrode are each configured to be in contact with a fluid passing along the fluid channel.
6. The microfluidic device of any preceding claim, wherein the first aperture and the second aperture have the same aperture diameter, and wherein a separation ratio, of the aperture diameter to the perpendicular distance between the first electrode surface and the second electrode surface, is between 0.1 and 10; preferably between 0.5 and 1.5.
7. The microfluidic device of any preceding claim, wherein the fluid channel has a channel width between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 200 micrometers; more preferably between 20 micrometers and 50 micrometers
8. The microfluidic device of any preceding claim, wherein the fluid channel has a circular cross section and has a diameter between 5 micrometers and 500 micrometers; preferably between 10 micrometers and 200 micrometers; more preferably between 20 micrometers and 50 micrometers.
9. The microfluidic device of any preceding claim, wherein one or more of: a first electrode ratio, of an area of the first electrode surface to the area of the first aperture, is 2 or above; and a second electrode ratio, of an area of the second electrode surface to the area of the second aperture, is 2 or above.
10. The microfluidic device of any preceding claim, wherein one or more of the first electrode and the second electrode comprise at least one of: graphene, a metallic solid film, a metallic conductive polymer; and a metallic nanoparticle film.
11. The microfluidic device of any preceding claim, further comprising a non-conductive polymeric layer between the first electrode surface and the second electrode surface.
12. The microfluidic device of any preceding claim, wherein: the first electrode is located in a first direction along the longitudinal axis away from the second electrode, the first electrode surface facing the second electrode surface; the second electrode is a second electrode plate comprising the second electrode surface and an opposite second electrode surface coplanar with and opposite to the second electrode surface, the second electrode plate configured to be connected to ground; the microfluidic device further comprising: a third electrode located in a second direction opposite the first direction along the longitudinal axis, the third electrode comprising a third aperture and a third electrode surface coplanar with and facing the opposite second electrode surface, the second electrode plate thereby located between the first electrode surface and third electrode surface, wherein: in response to a first alternating voltage applied to the first electrode, a first electric field is generated between the first electrode surface and the second electrode surface, and in response to a second alternating voltage applied to the third electrode, a second electric field is generated between the opposite second electrode surface and the third electrode surface.
13. The microfluidic device of claim 12, further comprising a first non-conductive polymeric layer between the first electrode surface and the second electrode surface, and a second non-conductive polymeric layer between the opposite second electrode surface and the third electrode surface.
14. The microfluidic device of any preceding claim, further comprising a voltage source configured to apply an alternating voltage to the first electrode with the second electrode connected to ground, thereby generating an electric field between the first electrode surface and the second electrode surface.
15. The microfluidic device of any of claims 12 to 13, further comprising a voltage source configured to apply: a first alternating voltage to the first electrode with the second electrode connected to ground, thereby generating a first electric field between the first electrode surface and the second electrode surface; and a second alternating voltage to the third electrode with the second electrode connected to ground, thereby generating a second electric field between the opposite second electrode surface and the third electrode surface.
16. The microfluidic device of any of claims 14 to 15, wherein the voltage source is configured to apply an alternating voltage between 0.1V AC and 50V AC; preferably between 0.5V AC and 10V AC; more preferably approximately 1V AC.
17. The microfluidic device of any of claims 14 to 16, wherein the voltage source is configured to apply an alternating voltage of a frequency between 10 kHz and 50 MHz; preferably between 250 kHz and 2 MHz; more preferably between 500 kHz and 1 MHz.
18. A method of determining a characteristic of a particle using the microfluidic device of claim 1, the method comprising: applying an alternating voltage to the first electrode and to the second electrode to thereby generate an electric field between the first electrode surface and the second electrode surface; passing a fluid containing the particle along the fluid channel in the fluid flow direction; determining a current variation in the fluid channel in dependence on the particlemoving through the electric field; anddetermining a parameter of the particle in dependence on the current variation.
19. The method of claim 18, wherein determining the parameter of the particle comprises determining a size of the particle.