Patent Application
20210331169
The present invention relates to an apparatus comprising a microfluidic chip useful for separating a particulate containing fluid into subsets of particulates. Also contemplated are methods of separating a particulate containing fluid into subsets of particulates, for examples subsets of cells. In particular, the invention relates to an on-chip microfluidic method of separating semen cells according to sex.
Most of the embodiments for the particle/cell separation described deal with a specific method of particles/cells identification based on AC impedance spectroscopy. We will therefore first review the technology of AC impedance spectroscopy while emphasising that the device and method for the particle/cell separation can also be used with other methods and devices for the particle/cell identification such as e.g. optical scattering or optical fluorescence method.
In recent years, microfluidic impedance cytometry has been further developed to count and discriminate between different kinds of cells. Multi-frequency impedance measurements can be used to determine the electrical properties of single cells in a microchip [S. Gawad, L. Schild, P. H. Renaud, Micromachined impedance spectroscopy flow cytometer for cell analysis and particle sizing, Lab. Chip. 2001 1 76-82].
Typical impedance measurement scheme described in preferred embodiments utilize detection of particles (cells) based on a lock-in amplifier measurement and AC impedance spectroscopy measurements. A lock-in amplifier is a type of amplifier that can extract a minute signal due to particle (cell) passing through the detection area from a noisy environment. Depending on the dynamic reserve of the instrument, signals up to one million times smaller than noise components and relatively close in frequency, can still be reliably detected and segregated. Lock-in amplifier is generally used to detect and measure very small AC signals, in the range of a few nanoVolts. The measured value is complex AC impedance between two electrodes separated by the fluid in the microchannel. The most important part of this amplifier is a digital low-pass filter, the part that isolates the frequency of interest and eliminates the higher frequencies and noise. In order to achieve a real-time analysis of the cells with the rate of up to many thousands of cells per second, the detection system must be very fast.
The amplifier is essentially a homodyne detector followed by low pass filter. Homodyne detector is a device that detects frequency-modulated signal by non-linear by mixing it with a signal having the same frequency as the modulation signal. Whereas conventional lock-in amplifiers use analogue frequency mixers and R-C filters for the demodulation, state of art instruments have both steps implemented by fast digital signal processing for example using Field Programmable Gate Array (FPGA). FPGA is an integrated digital circuit whose logic behaviour and component mapping is programmable. Sine and cosine demodulation is performed simultaneously, which is sometimes also referred to as dual phase demodulation.
Operation of a lock-in amplifier relies on the orthogonality of sinusoidal functions. When a sinusoidal function of frequency f1 is multiplied by another sinusoidal function of frequency f2 not equal to f1 and integrated over a time much longer than the period of the two functions, the result is zero. Instead, when f1 is equal to f2 and the two functions are in phase, the average value is equal to half of the product of the amplitudes.
A lock-in amplifier takes the input signal, multiplies it by the reference signal (either provided from the internal oscillator or an external source), and integrates it over a specified time, usually on the order of milliseconds to a few seconds. The resulting signal is a DC signal, where the contribution from any signal that is not at the same frequency as the reference signal, is attenuated close to zero. The out-of-phase component of the signal that has the same frequency as the reference signal is also attenuated (because sine functions are orthogonal to the cosine functions of the same frequency), making a lock-in a phase-sensitive detector (PSD).
Usually, lock-in amplifiers generate their own internal reference signal by a phase-locked-loop (PLL) locked to the external reference. The internal reference is VL sin(ωLt+θref), where ωy is the angular frequency of the signal and θref is the phase of the reference signal.
The lock-in amplifies the signal and then multiplies it by the lock-in reference using a phase-sensitive detector or multiplier. The output of the PSD is simply the product of two sine waves.
The PSD output is two AC signals, one at the difference frequency [ωr−ωy] and the other one at sum frequency [ωr+ωL]. Then the PSD output passes through a low pass filter and the AC signals are removed.
If ωr equals ωy, the difference frequency component in eq. (1.1) will be a DC signal. So, the filtered PSD output will be:
V
psd=½VsigVL cos(θsig−θref) (1.2)
which is a DC signal proportional to the signal amplitude.
Unlike in analogue amplifiers, in a digital lock-in the signal and reference are converted into a digital form by Analogue-to-Digital converters (ADC) and are represented by sequences of numbers. Multiplication and filtering are performed mathematically by a digital signal processing (DSP). The phase sensitive detectors (PSDs) in the digital lock-in act as linear multipliers; that is, they multiply the signal by a reference sine wave. Analogue PSDs (both square wave and linear) have many problems associated with them such as output offsets and gain error. Output offset is a problem of analogue lock-ins because the signal of interest is a DC output from the PSD, and an output offset contributes to error and zero drift. The offset problems of analogue PSDs are eliminated using the digital multiplication of the signal and the reference. Analogue sine-wave generators are susceptible to amplitude drift: especially as a function of temperature. In contrast, the digital reference sine wave has a precise fixed amplitude. This avoids a major source of gain error common to analogue lock-ins. For applications that require high precision, short detection time, digital lock-in amplifiers are preferred to the analogue ones. Flow-cytometry is one such application.
To detect and identify particles in a microfluidic channel using the approach of AC impedance spectroscopy, the analogue input is a modulated signal that passes through a particle-containing fluid. It is applied as AC voltage at excitation frequency f0 between detection electrodes normally positioned on both sides of a common microfluidic channel in the detection area of the channel. The voltage is applied to one of the two opposite electrodes and the signal is taken from the second electrode and sent to the lock-in amplifier. The signal is a complex function of the impedance of the cell, impedance of the elements of the cell, impedance of the fluid surrounding the cell, dielectric properties of the cell and its elements, dielectric properties of the fluid carrying the cell, the shape of the electrodes, size of the cell compared to the size of the electrodes, position of the sell with respect to the electrodes and orientation of the cell with respect to the electrodes. Dielectric properties of the cell and the fluid as well as impedance of the cell and the fluid are complex functions with real and imaginary components. All these also depend of the excitation frequency f0.
Complete analytical description of the system in terms of the first principles is too complex and is far beyond the scope of this document. However, the complete analytical description is also unnecessary for the detection and identification of cells. The detection of cells is based on changes in the signal send to the lock-in amplifier resulting from the individual passing cells. The identification of the single particles (cells) is based on a principle: identical particles (cells) positioned in the same way with respect to the electrodes, will produce identical signals.
The signal induced is normally very small and it comes against significant noise background and therefore lock-in detection is used to demodulate the signal, extract it from the noise and measure it. It is important to note that the signal from the particles (cells) comes at frequency f0, the same as the excitation frequency. After multiplying the modulated signal by a sinusoid with frequency f1, we obtain a sum of sinusoids with different frequencies, i.e. a combination of modulation frequency and reference frequency, as described above. By applying the correct type of filter is possible to isolate the sinusoid with the frequency f0 we are interested in. After the demodulation, the signal must be sent via low-pass filter. There are number of options for the low-pass filter such as: FIR (Finite length Impulse Response) filters and IIR (Infinite length Impulse Response) filters. IIR filters contain feedback, whereas FIR filters do not. As a result the nth output sample, y[n], of an FIR filter is a function of only a finite number of samples of input sequence, whereas the nth sample of an IIR filter may depend upon an infinite number of samples. FIR filters are mostly implemented directly using a relation similar to the convolution sum, although they can also be implemented using discrete Fourier transforms (DFTs).
Paper [T. Sun and H. Morgan, Single-cell microfluidic impedance cytometry: a review, Microfluid. Nanofluid, 2010, 8, 423-443] describes several methods where cells flow between miniature electrodes which have an AC field applied across them. As the cell passes between the electrodes, the current path is disturbed and the change in current gives a change in the impedance signal associated with a single cell. Usually, impedance measurements at the frequency of (1-5 MHz) give information on the cell membrane capacitance whilst much higher frequencies (>10 MHz) probe the internal properties of the cell. Two or more frequencies can be applied simultaneously to differentiate different types of cells.
The inventors have also investigated impedance flow spectroscopy method of cell detection where two pairs of electrodes are used similar to configuration described in [“Microfluidic impedance cytometer for platelet analysis”, Mikael Evander et. al, Lab on a chip, volume 13, 2013], each pair having an excitation and measurement electrodes. An AC voltage at radio frequency from 100 KHz-100 MHz is applied to an excitation electrodes and an electrical current is measured by the measurement electrodes. The electrical current being measured is then amplified and converted into an output voltage. The output signal is then demodulated to remove excitation frequency and to recover impedance magnitude and phase. As a cell passes through the pair of excitation and measurement electrodes, impedance magnitude and phase change, thus recording the information about the cell properties. Additional pair of electrodes ensures measurement is differential thus eliminating parasitic electromagnetic noise. Typically, the impedance measurements are taken at low frequency: 100 KHZ-2 MHz to acquire information about the cell size and at high frequency 2-100 MHz to acquire information about the cytoplasm and internal properties of the cell.
Impedance flow cytometry can readily detect a cell, and the original technique was developed by Coulter for this. When it comes to more challenging task of identifying the sub-populations of cells within the sample fluid, the performance of the impedance cytometry is much less convincing due to large spread in the data points corresponding to each cell. Hydrodynamic focusing is particularly important for the detection of cells and particles on a chip utilizing impedance measurements. Indeed, for identification of cells (or particles) it is necessary to arrange these in such a flow that they pass in front of the detection system one by one. This “one cell-by-one cell” principle is fundamental for the successful cell identification: one needs to avoid the situation of multiple cells passing through the detection system at once as it could prevent the identification. It is common to use hydrodynamic focusing to achieve this. Hydrodynamic focusing is based on injection of the sample fluid into the laminar flow of sheath fluid. The two flows then merge into to a single channel, usually of a reduced cross-section. This reduces the cross-sections of both, the sheath fluid part of the flow and also the sample liquid flow, and thus achieves the desired reduction in the cross-section of the sample fluid flow. To control the cross-section of the sample fluid, one could change the flow rates of the sample fluid and sheath fluid. In relation to the electrical impedance-based cytometry, hydrodynamic focusing reduces the width of the conductive sample stream to the appropriate size of the cells, increasing the percentage resistance change in the conductive path when a cell passes by.
In recent years there is increasing body of work on the use of hydrodynamic focusing in microfluidic chips and microchannels. For example, the Japanese patent laid-open No 2003-107099 discloses a “fractionation microchip having a channel for introducing a particulate-containing solution, and a sheath flow forming channel arranged on at least one lateral side of the of the introducing channel. The particulate fractionation microchip disclosed in Patent 2003-107099, is so designed that fluid laminar flows are formed by a “trifurcated channel” having a channel for introducing a particulate-containing solution and two sheath flow-forming channels. In essence this is a 2D hydrodynamic focusing on a chip. Similar approach is described in [R. Rodriguez-Trujillo, C. Mills, J. Samitier, G. Gomila, Microfluid. Nanofluid, 3 171 (2007)] and [P. Walsh, E. Walsh, M. Davies, Int. J. Heat Fluid Flow 28 44 (2007)].
There are also solutions for integration of 3-D hydrodynamic focusing with a conventional type microfluidic chip described in [“Three-dimensional hydrodynamic focusing in a microfluidic Coulter counter”, R. Scott, P. Sethu, C. K. Harnett, Rev. Sci. Instruments 79 046104 (2008)]. A similar approach is described in [“Universally applicable three-dimensional hydrodynamic microfluidic flow focusing” Yu-Jui Chiu, S. H. Cho, Z. Mei, V. Lien, T. F. Wu, Y. H. Lo, Lab Chip 2013 13 1803]. That study deals with three-dimensional hydrodynamic focusing where the sample channel and the two sheath channels having a greater height than the sample channel, join at the junction before the main channel which has the same height as the sheath channel. The merging of channels of different heights produces flow confinement both in the lateral and transverse directions, resulting in 3D focused flow. In the patent WO 2008/125081 A1 Theisen Janko et al provides the method for focusing fluid in microfluidic channel structure and the implementation of such a microfluidic structure to achieve hydrodynamic focusing of fluid. In the patent US 2009/0283148 A1, Shinoda et. al. [US 2009/0283148 A1, “Microchip and channel structure for the same”, Masataka Shinoda, May 4 2009] teaches the method of three dimensional hydrodynamic focusing where the microtube is inserted into the microchip to providing the sample flow.
Despite these efforts, integration of 3D hydrodynamic focusing with microfluidic chip is not simple and the performance of such on-chip 3D focusing has limited capability. To reduce the variation of the impedance flow cytometry measurements, it is desirable to be able to direct the sample flow through a well-defined point in between the electrodes, e.g. the centre of the channel. This reduces the spread in the data points from a single population of cells of type of particles in the flow. It may also be desirable to align all the cells (particles) in the same way with respect to the direction of the electric field created by the electrodes. Cells often do not have an overall spherical shape but are rather elongated, ellipsoidal or discoid in shape. The signal from the cell in electrical impedance cytometry device depends on the orientation of the elongated axis of the cell with respect to the electrodes.
Having reviewed the methods and technologies for identification of particles/cell, especially those based on AC impedance spectroscopy, we are now moving to the issue of particles/cells separation. Once the particles/cell are identified it is often necessary to separate different sub-sets of particles/cell into different streams or different containers. One method is based on ejecting the flow of liquid containing particles/cell via a termination at the end of the microfluidic channel so that it breaks into minute droplets, one drop containing one particle/cell. This can be done by applying high voltage to the end of the channel. For this, a high-voltage generator is connected to the end of the microfluidic channel. This is also sometimes referred to in the industry as electrospray. The shape of the microfluidic channel termination is essential for the process of splitting the continuous jet emerging from the channel into a train of droplets. Therefore, commonly a nozzle is attached at the end of the channel to introduce control and reproducibility into the process of splitting the continuous jet into a train of droplets. The droplets are charged by the very same high voltage generator connected to the microfluidic channel. Then the droplets pass through an area of an electric field that is generated by control electrodes that are parallel to the pathways of the passing droplets. The electric field is perpendicular to the pathways of the passing droplets. The droplet trajectories ejected from the nozzle into the space in between the control electrodes, can be altered by applying a control voltage to the control electrodes. The control voltage is applied in the form of a pulse and this needs to be coordinated with the output of the device measuring and identifying the cells (e.g. fluorescence detector) so that particles/cells of the desired sub-set of particles cells move into a different trajectory and arrive to a required destination container. There are many embodiments of technology utilising this principle with various methods of forming drops/aerosols containing particles/cells. For further details one can refer to [D. Parks, B. Sahaf, O. Perez, M. Roederer, L. A. Herzenberg, Clinical Chemistry, (2002) 48: 1819-1827; W. A. Bonner, H. R. Hulett, R. G. Sweet, L. A. Herzenberg, Review of Scientific Instruments, (1972) 43: 404-409].
There is one disadvantage of this technology. It is convenient to have separation of particles/cells achieved in-situ, in the very same microfluidic chip, as opposed to letting the cells leave the microfluidic device into the ambient and perform separation in the form of droplets in the ambient.
US patent application 2009/0283148 A1 to M. Shinoda and T. Takashimizu attempts to address this issue. The particles are packed into droplets travelling through a microfluidic channel. Along the channel, there are positioned two electrodes generating electrostatic field applying a Coulomb force acting on the droplets.
There are a number of publications and inventions related to the use of electrophoretic, dielectrophoretic or electroosmotic forces for separation of cells. We shall therefore explain the terms electrophoresis, dielectrophoresis and electroosmosis and forces resulting from these. In electrophoresis the particles/cells are charged by their interaction with the surrounding fluid and the charged particles move along or against the direction of electrostatic field. The force is proportional to the charge of the particle/cell. Examples of systems utilising electrophoretic forces can be found in [J. Voldman, Annual Review of Biomedical Engineering, (2006), 8: 425-454. Pub Med 16038291;
In contrast to electrophoresis, where cells move in a uniform electric field due to their surface charge, dielectrophoresis refers to the movement of particles (cells) in a non-uniform electric field due to their polarizability. For movement in response to a dielectrophoretic force, particles (cells) do not need to possess a surface charge. The surface charges are produced by the field by disproportionation (pulling) the opposite charges to different parts of the cell. Once exposed to a gradient field, cells migrate either towards or away from the region of strongest field intensity depending on the electrical permeability (permittivity) of the cell and the fluid. Cells with a higher permeability than the fluid are attracted toward the field maxima, which is known as positive dielectrophoresis. The opposite is true for negative dielectrophoresis. The magnitude of the dielectrophoretic force is dependent on the size and properties of the cell, fluid, and the parameters of the electric field, which is useful for sorting cells by size and dielectric properties. Review of the phenomenon for the separation of particles/cells can be found in [J. Voldman, Annual Review of Biomedical Engineering, (2006), 8: 425-454. Pub Med 16038291]. We stress that the dielectrophoresis is based on gradient-, i.e. non uniform field. Usually papers described using AC gradient field, i.e. AC field with amplitude that varies strongly in space. The use of AC field is preferable as DC field may be screened by a conducting liquid surrounding the cells or by the cells themselves.
Dielectrophoresis force was used by a number of authors for the separation of cells and other microscopic objects. Review [M. P. Hughes, Strategies for dielectrophoretic separation in laboratory on chip format, Electrophoresis 2002, 23, 2569-2582] describes strategies for the separation of bioparticles such as cells, viruses and proteins in miniaturised laboratory-on-a-chip format. The separation principle in based on difference in polarisation of particles. The review describes various electrodes such as stepped electrodes, “Christmas tree” electrodes, spiral electrodes creating controlled gradient (non-uniform) electric field. Publication [B. H. Lapizco-Encinas, B. A. Simmons, E. B. Cummings, Y. Fintchenko, Dielectrophoretic concentration and separation of live and dead bacteria in an array of insulators, Anal Chem 2004, 76, 1571-1579] describes the use of dielectrophoresis for separation of dead and alive cells held within a body of an insulator. The paper [N. Lewpiriyawong, K. Kandaswamy, C. Yang, V. Ivanov, R. Stocker, Microfluidic characterization and continuous separation of cells and particles using conducting poly(dimethyl siloxane) electrode induced alternating current-dielectrophoresis, Anal. Chem. 2011, 83, 9579-9585] describes separation of live yeast cells from the dead ones in a microfluidic channel under the influence of the dielectrophoretic forces in gradient electric field along the channel. The principle is based on having a different value of the dielectrophoretic force acting on live and dead cells. Other publications [H. Shafiee, M. B. Sano, E. A. Henslee, J. L. Caldwell, R. V. Davalos, Selective isolation of live/dead cells using contactless dielectrophoresis (cDEP), Lab Chip, 2010, 10, 438-445] and [H. Li, R. Bashir Dielectrophoretic separation and manipulation of live and heat treated cell of Listeria on microfabricated devices with integrated electrodes, Sensors and Actuators B 86 (2002) 215-221] describe a similar approach. U.S. Pat. No. 7,294,249 to S. Gawad, M. Wuethrich, P. Renaud, describes microfluidic component and method for sorting particles in a fluid utilising such a dielectrophoretic force. For this there are electrodes made protruding over the micorfludic channel generating gradient electric field applying dielectrophoretic force on the particle as they flow through the channel.
A number of publications use the approach described earlier in relation to the electrophoresis, i.e. encapsulating single cells into emulsified droplets and sorting he droplets using dielectrophoretic force [J. J. Agresti, E. Antipov, A. R. Abate, K. Ahn, A. C. Rowat, J. C. Beret, M. Marquez, A. M. Klibanov, A. D. Griffitths, D. A. Weitz, Proceedings of the National Academy of Sciences of the United States of America, 2010; 107: 4004-4009; L. Mazutis, J. Gilbert, W. L. Ung, D. A. Weitz, A. D. Griffiths, J. A. Heyman, Nature Protocols, 2013, 8: 870-891].
Electroosmotic sorting of two different types of E. coli cells was demonstrated in [A. Y. Fu, C. Spence, A. Scherer, F. H. Arnold, S. R. Quake, Nature Biotechnology, 1999; 17: 1109-1111]. In this work fluorescence was used to identify the cells and the difference was that one type of E. coli expressed a particular protein while the other one did not. The separation was demonstrated in the format of two microchannels forming a T-junction. There were three terminals for making electric contacts at the three distal ends of the microchannels forming the T-junction. A DC voltage was applied between the three terminals forming potential difference between the terminals and therefore resulting in electroosmotic pumping. High separation speeds are described, but passing electrical current along the full length of the microfluidic channel is not optimal for living cells and limits the application of the methodology. In addition, high voltages are required due to the length of the electrical field which would lead to bubble formation along the separation electrodes. It is an object of the invention to overcome at least one of the above-referenced problems. Similar approach was demonstrated in [P. S. Dittrich, P. Schwille, Analytical Chemistry, 2003; 75: 5767-5774] for separation of Escherichia Coli, again based on fluorescence detection. Silicon microfluidic chip was used and the connection of the voltages to drive electroosmosis was similar to the one in the publication cited previously. The voltage was applied along the channels and the switching was demonstrated in the DC mode as opposed to cell sorting at high pace.
All these publications and inventions describing the use of electrophoretic or dielectrophoretic force for the separation of particles/cell pursue a different paradigm to the one described in this invention. Their approach is based on using the force that selectively interacts with one sub-set of particles/cell but not with the others, e.g. it has significant interaction with live cells but little or no interaction with dead cells. Therefore the separation is done by the selective mechanism of effect of the stimulus on cells: some of them are affected by the stimulus while others are not.
We would like to develop a separation method where the interaction stimulus, once it is switched on, acts on the entire flow of particles/cells including ALL the sub-sets of particles/cells in the flow not just one sub-set. Furthermore, we would like to be able to switch this interaction on and off as required to guide the cells into the correct destination locations at high repletion rate. The rate of activating/deactivating the interaction should be as high as 100-10000 interactions per second or even higher. In this sense our requirements are similar to the ones delivered by the electrostatic separation of cells packed in droplets as described above but with the added restriction that the separation is done in the on-chip format.
There are some other methods are based on applying a force or pressure pulse to the flow, or fragmenting the flow into droplets and applying a force to defect the droplets. These methods are significantly outside the scope of the present invention. Therefore we will only give a brief review of these. There are methods based on acoustic streaming forces generated from non-standing surface acoustic waves to redirect fluid flow from one outlet to an adjacent outlet. An example of these can be found in [T. Franke, S. Braunmuller, L. Schmid, A. Wixforth, D. A. Weitz, Lab Chip. 2010; 10: 789-794]. There are methods based on separation of cells based on standing acoustic waves [S. C. Lin, X. Mao, T. J. Huang, Lab. Chip. 2011; 11: 1280-1285;
S. C. Lin, X. Mao, T. J. Huang, Lab. Chip. 2012; 12: 2766-2770].
There are methods based on optical manipulation [A. Jonas, P. Zemanek, Electrophoresis, 2008, 29: 4813-4851]. There are methods based on generating magnetic force. We will not review these as they are not related to the invention and are limited to magnetic particles. There are methods based on generating a pressure pulse in the channel by a mechanical actuator, e.g. by a piezoactuator.
Patent application EP 2 508253 A1 deals with separation of molecules in microfluidic channels. There is a branching point positioned along a channel and a set of electrodes positioned in the vicinity of the branching point. According to the inventors, the workings of EP 2 508253 A1 is based on the fact that different types of molecules respond differently to an electric stimulus that is persistently applied to the electrodes. Typically they describe the stimulus as AC voltage applied to the electrodes. In paragraph 054 the document describes the parameters of the field: the frequency is “about several MHz” and the electric field is “several MV/m”. Creating an electrical field at the frequency of several MHz and the amplitude of several MV/m in conventional laboratory conditions and applying this field to a delicate microfluidic chip is highly challenging at best. This method and device cannot be applied for the guidance for the problem addressed by our invention as AC voltage/field at the frequency of several MHz is not able of guiding the cells within the channels. The voltage of “several MV/m” is close to the catastrophic breakdown voltage of a good dielectric. This voltage would likely lyse (kill) cells and destroy the cell membrane. Furthermore, such a voltage applied to interior of the channel would likely destroy electrodes even if they are made of a noble metal. US Patent Application US2017/0128941 A1 to Sadri et. al. deals with the apparatus and method for the microfluidic manipulation, dispensing and/or sorting of particles such as cells and/or beads. The apparatus describes a channel network with a single or multiple inlet channel and multiple outlet channels so that one inlet channel branches into a several outlet channels. There is also a detector located upstream from the branching point. There is further pressure source that could be activated if the particle meets one or several criteria and that can guide the particle into the required outlet channel. The invention describes a number of actuators capable of sending the cells into the destination outlet channel. According to the inventors these may include (i) a parallel slide diverter: a system of micro slides that can slide in- and out-of the channel to keep it open and closed when required; (ii) a parallel channel vacuum diverter actuated by a flexible membrane that allows vacuum to be introduced on demand into a channel and change the direction of the particles flow; (iii) parallel channel paddle diverter that is controlled by a rotary paddle; (iv) parallel channel blocking diverter that can block the channel by a pressure-activated flexible membrane; (v) dielectrophoresis diverter that can apply a gradient electric field to the cells and cause their movement; (vi) cross channel valve diverter that is based on microvalve or a number of microvalves; (vii) cross channel rotary diverter that is a type of a rotary valve; (viii) multichannel pressure diverter; (ix) offset channel valve diverter. The dielectrophoresis diverter (v) is based on embedding two electrodes into the walls of the inlet channel that are not placed in electric contact with the interior of the channel and create a gradient electric field perpendicular to the axis of the channel. It is difficult to create electric fields in the case of certain cell containing liquids by applying electrodes outside the channel as the liquid in the channel is conducting and therefore screens the electric field.
It is an object of the invention to overcome at least one of the above-referenced problems.
The present invention addresses the problems of the prior art and provides an on-chip microfluidic cell sorter that employs electroosmotic separation of cells without damage to the cells or the need for high voltages known to produce bubbles at the electrodes. Another problem with the current approach is that cells are subjected to electric current all along the channel, i.e. the cells are subjected to current for extended length of time, not just to a short pulse of current. The extensive exposure to electric current affects the cells condition, e.g. viability of sperm cells. The invention employs separation electrodes in the separation zone of the microfluidic channels to generate an electric current/field that is limited to the separation zone. Unlike Fu et al, which teach an electrical field/current that extends along the whole of the microfluidic channel, the present invention employs separation electrodes that are disposed in the sidewall of the microfluidic channel in the separation zone, and thus produces an electrical field/current that is generally orthogonal or tangential to the direction of flow of the fluid in the microfluidic cannel. Therefore, each cell is subject only to a single short pulse of current. For selection of a single sub-set of cells out of a set of cells one could even entirely avoid the need for current pulse as will be explained below. Unlike the method of US2017/128941, which employs electrodes embedded into the wall of the channel without making contact with the fluid in the channel, and creates a dielectrophoretic force across the channel which is prone to the being screened by the fluid in the channel if it is conducting, the present invention overcomes this problem by employing separation electrodes in the separation zone of the microfluidic channels that are in contact with the fluid in the channel and drive current through the liquid. This allows greater amounts of energy to be transferred to the cells, and the current flowing into the channel causes unequal flow of liquid between different secondary microfluidic channels. US2017/12894 also mentions that the microfluidic channels could have metal walls (Paragraph 40). However, as the metal walls of the channel would screen any electrical field produced by the electrodes perpendicular to the axis of the channel, this would zero the electric field in the channel.
According to a first aspect of the present invention, there is provided an apparatus for separation of particulates in a fluid into subsets of particulates, the apparatus comprising a microfluidic chip comprising a microfluidic channel having a fluid inlet for receipt of a stream of particulate containing fluid, a detection zone disposed in the microfluidic channel and comprising a sensor configured to detect changes in the microfluidic channel corresponding to anisotropic particles passing the sensor, and a separation zone distal of the detection zone in which the microfluidic channel divides into at least two secondary microfluidic channels. The separation zone typically comprises a pair of separation electrodes including a first separation electrode disposed in electrical contact with an interior of the microfluidic channel and a second separation electrode disposed in electrical contact with an interior of the microfluidic channel or one of the secondary microfluidic channels. The pair of separation electrodes are typically configured during use to pass a pulse of current, typically direct current, between the electrodes and across the microfluidic channel in the separation zone.
In one embodiment, the apparatus comprises a voltage/current generator electrically connected to at least one or more of separation electrodes and configured to supply a voltage/current pulse to these/this separation electrodes. In one embodiment the voltage/current pulse is a direct current (DC) pulse. In one embodiment, the pair of separation electrodes are configured to be electrically (ohmically) coupled to each other via the fluid in the channel during the pulse of current through the microfluidic channel.
In one embodiment, the apparatus comprises a voltage/current generator electrically connected to at least one or more separation electrodes and configured to supply a different voltage/current pulse to each of the pair of separation electrodes.
In one embodiment, the voltage/current generator is configured to apply a voltage pulse to the separation electrode in the range of 0.01-50 V of either positive or negative polarity.
In one embodiment, the voltage/current generator is configured to apply a voltage pulse to the separation electrode with the duration in the range of 1 microsecond to 50 milliseconds.
In one embodiment, the voltage/current generator is electrically connected to one of the pair of separation electrodes and the other of the pair of electrodes is connected to a fixed potential.
In one embodiment, the apparatus includes a processor operatively connected to the sensor and the or each voltage/current generator and configured to receive signals from the sensor in the detection zone and actuate the or each voltage/current generator in response to signals received from the sensor.
In one embodiment, both of the separation electrodes are disposed in electrical contact with the microfluidic channel.
In one embodiment, the separation electrodes are disposed on opposite sides the microfluidic channel.
In one embodiment, the separation electrodes are longitudinally aligned so that during use the electrical field generated by the electrodes is substantially orthogonal to a longitudinal axis of the microfluidic channel.
In one embodiment, the separation electrodes are longitudinally offset so that during use the electrical field generated by the electrodes is tangential to a longitudinal axis of the microfluidic channel.
In one embodiment, one of the pair of separation electrodes is disposed in electrical contact with the microfluidic channel and another of the pair of separation electrodes is disposed in electrical contact with one of the secondary microfluidic channels.
In one embodiment, the apparatus includes two pairs of separation electrodes.
In one embodiment, the apparatus includes two electrodes in the microfluidic channel and an electrode in each of the two secondary microfluidic channels. In one embodiment, the electrodes in each of the two secondary microfluidic channels are disposed on an inner sidewall of the secondary microfluidic channels (i.e.
In one embodiment, the sensor in the detection zone is configured to measure AC impedance change between a set of detection electrodes positioned in the microfluidic channel corresponding to particulates obstructing the electric path between the detection electrodes.
In one embodiment, the set of detection electrodes comprises two pairs of detection electrodes.
In one embodiment, the or each set of detection electrodes comprises an energising electrode and a signal electrode electrically coupled to a lock-in amplifier.
In one embodiment, the sensor is configured to make optical measurements of the light scattered by the particulates passing along the detection zone of the microfluidic channel.
In one embodiment, the sensor is configured to make optical measurements of the fluorescence signal from the particulates passing along the common microfluidic channel.
In one embodiment, a cross-section of microfluidic channel perpendicular to the flow direction has a width of 20-500 μm and a height of 20-500 μm.
In one embodiment, the linear velocity of the fluid in the middle of the common microfluidic channel is 10 mm/sec.
In one embodiment, the apparatus includes a hydrodynamic focusing mechanism configured to hydrodynamically focus the stream of particulate containing fluid to provide a focused sample fluid within a guidance/sheath fluid.
In one embodiment, the apparatus includes an on-chip hydrodynamic focusing mechanism comprising a sample channel and a guidance fluid channel that merge at an oblique angle to form the microfluidic channel.
In one embodiment, the apparatus is further equipped with means for positioning of particles/cells within the cross-section of the microfluidic channel.
In one embodiment, the cross-sectional area of the secondary microfluidic channels is greater than the cross sectional area of the microfluidic channel.
In another aspect, the invention provides a method to separate particulates in a particulate containing fluid, which method employs an apparatus according to the invention. The method suitably comprises the steps of passing a particulate containing fluid along the microfluidic channel, employing the sensor to detect a change in the microfluidic channel corresponding to a particulate passing the sensor in the detection zone, and actuation of the separating electrodes in the separation zone in response to the change detected in the microfluidic channel to pass a pulse of current through the microfluidic channel and deflect the particulate into one of the secondary microfluidic channels.
In one embodiment, the current amplitude is configured to change the pathway of the cells between the microfluidic channel and the secondary microfluidic channels.
The method of the invention may be employed to separate a particulate containing fluid into at least first and second subpopulations of cells (for example X and Y subpopulations of semen cells). The separation electrodes may be configured to pass a single pulse of current through the separation zone of the microfluidic channel (i.e. a pulse corresponding to one of the subpopulations), which separates one of the subpopulations of cells. In this embodiment, the microfluidic channel can suitably bifurcate to provide a side secondary channel configured to receive deflected cells (i.e. one subpopulation), and a main secondary channel configured to receive non-deflected cells (i.e. the other sub-population or sub-populations). Thus, the main secondary channel which is nearly co-linear with the microfluidic channel, and a side secondary channel which diverts away from the main channel. In another embodiment, the separation electrodes may be configured to pass more than one pulse of current through the separation zone of the microfluidic channel in a sequential manner corresponding to the signals received in the detection zone (i.e. a first pulse corresponding to one of the subpopulations, and a second pulse corresponding to another of the subpopulations). In this embodiment, the secondary channel may take the bifurcated form shown in
In one embodiment, the invention is for separating a first and second subpopulations of particulates from the particulate containing fluid, including the steps of:
detecting a first change in the microfluidic channel corresponding to a particulate of the first subpopulation passing the sensor in the detection zone;
actuation of the separating electrodes in the separation zone in response to the first change detected in the microfluidic channel to pass a first pulse of current through the microfluidic channel and deflect the particulate of the first subpopulation into a first of the secondary microfluidic channels;
detecting a second change in the microfluidic channel corresponding to a particulate of the second subpopulation passing the sensor in the detection zone; and
actuation of the separating electrodes in the separation zone in response to the second change detected in the microfluidic channel to pass a second pulse of current through the microfluidic channel and deflect the particulate of the second subpopulation into a second of the secondary microfluidic channels.
In one embodiment, the invention is for separating a first and second subpopulation of particulates from the particulate containing fluid, and in which the secondary microfluidic channels include a side secondary channel configured to receive deflected particulates pass and a main secondary channel configured to receive un-deflected particulates, the method including the steps of:
detecting a first change in the microfluidic channel corresponding to a particulate of the first subpopulation passing the sensor in the detection zone;
actuation of the separating electrodes in the separation zone in response to the first change detected in the microfluidic channel to pass a first pulse of current through the microfluidic channel and deflect the particulate of the first subpopulation into the side secondary microfluidic channels; and allowing the un-deflected second subpopulation of cells pass into the main secondary channel.
In one embodiment, the particulates are anisotropic particulates.
In one embodiment, the particulates are cells, and in which the method separates the cells according to a phenotype (for example, cell type, cell status (dead or alive), cell sex, cell health etc.).
In one embodiment, the particulates are semen cells, and in which the method separates the cells according to a sex (i.e. into X and Y populations).
In one embodiment, a first voltage is applied to the first separation electrode and a second voltage is applied to the second separation electrode.
In one embodiment, a first voltage is applied to one of the separation electrode pair and the other of the separation electrode pair is connected to a fixed potential.
In one embodiment, a first voltage and second voltage are each, independently, positive or negative polarity.
In one embodiment the voltage pulses applied to the separation electrode/electrodes are of a substantially non-rectangular shape such as bell-shaped or pulses with smeared limiting fronts of the pulses signifying the start and the end of the pulse.
In one embodiment the separation zone is separated from the detection zone by means of shielding electrodes connected to a fixed potential or ground potential.
In one embodiment, the apparatus comprises two pairs of separating electrodes including two electrodes in the microfluidic channel and an electrode in each of the two secondary microfluidic channels, wherein the two electrodes in the microfluidic channel are connected to a fixed potential and one of the electrodes in the secondary microfluidic channel is energised with a positive voltage and one of the electrodes in the secondary microfluidic channel is energised with a negative voltage.
In one embodiment, the apparatus comprises two pairs of detection electrodes, and in which the method includes a step of determining the speed of particulate fluid in the microfluidic channel by detecting the time it takes a particulate to pass from a first pair of detection electrodes to a second pair of detection electrode.
In one embodiment, the determined speed of the particulate containing fluid is correlated with a distance between the first pair of detection electrodes and the separation electrodes to calculate a time delay between (a) detecting a change in the microfluidic channel associated with a particulate passing the detection electrode and (b) actuation of the separating electrodes to deflect the particulate into one of the secondary microfluidic channels.
Other aspects and preferred embodiments of the invention are defined and described in the other claims set out below.
All publications, patents, patent applications and other references mentioned herein are hereby incorporated by reference in their entireties for all purposes as if each individual publication, patent or patent application were specifically and individually indicated to be incorporated by reference and the content thereof recited in full.
Where used herein and unless specifically indicated otherwise, the following terms are intended to have the following meanings in addition to any broader (or narrower) meanings the terms might enjoy in the art:
This invention deals with instruments for counting, identification and separation of particles in particulate-containing fluids such as e.g. flow cytometry or particle flow analyser. For many common envisaged applications in this invention, the particles in the fluid are cells, and particulate containing fluid could be a suspension of mammalian cells, semen, blood cells, viruses, bacteria or indeed any other type of isolated particles. However, the technology could also be used with numerous other types of particles: organic, inorganic, ceramic, composite, nanoparticles, etc., both solid, semi-solid or liquid (e.g. protein particles, fatty particles, blends of soft particles and their compositions, drops of one liquid in the stream of another one, etc.). The solid particles could be particles of metals, oxides, nitrides, sulphides, polymer particles and particles of numerous other inorganic and organics materials, also mixed particles containing blends and composites of materials within individual particles and various nano- and micro-particles and clusters. We shall use the term “particle” or “cell” to cover any and all of these. We shall also use the same to describe clusters of particles/cells thus forming larger particle/cell. We shall call the liquid/fluid carrying particles in microfluidic channel as sample fluid, or particle containing fluid or particulate containing fluid.
The emphasis in the document is on technologies that can perform counting and identification of particles and/or cells in a microfluidic chip format. The said microfluidic chip has at least one microfluidic channel where such counting and identification takes place in the detection area (zone) of the channel. The preferred embodiment deals with the method of counting and/or identification of particles/cells based on AC electrical impedance spectroscopy. For each passing particle/cell, such counting and/or identification is normally done before the separation step.
The term “separation” of particles/cells is used to describe the process of physical separation of these, e.g. moving different subsets of the entire set towards different destination points. We will use the term “identification” of particles/cells to describe the process of attaining the information on the sub-set to which the particle/cell belongs to.
Unless otherwise required by context, the use herein of the singular is to be read to include the plural and vice versa. The term “a” or “an” used in relation to an entity is to be read to refer to one or more of that entity. As such, the terms “a” (or “an”), “one or more,” and “at least one” are used interchangeably herein.
As used herein, the term “comprise,” or variations thereof such as “comprises” or “comprising,” are to be read to indicate the inclusion of any recited integer (e.g. a feature, element, characteristic, property, method/process step or limitation) or group of integers (e.g. features, element, characteristics, properties, method/process steps or limitations) but not the exclusion of any other integer or group of integers. Thus, as used herein the term “comprising” is inclusive or open-ended and does not exclude additional, unrecited integers or method/process steps.
In the context of treatment and effective amounts as defined above, the term subject (which is to be read to include “individual”, “animal”, “patient” or “mammal” where context permits) defines any subject, particularly a mammalian subject, for whom treatment is indicated. Mammalian subjects include, but are not limited to, humans, domestic animals, farm animals, zoo animals, sport animals, pet animals such as dogs, cats, guinea pigs, rabbits, rats, mice, horses, cattle, cows; primates such as apes, monkeys, orangutans, and chimpanzees; canids such as dogs and wolves; felids such as cats, lions, and tigers; equids such as horses, donkeys, and zebras; food animals such as cows, pigs, and sheep; ungulates such as deer and giraffes; and rodents such as mice, rats, hamsters and guinea pigs. In preferred embodiments, the subject is a human.
“Particulate” as applied to a particulate containing fluid means a solid body in the fluid or a semi-solid, i.e. a body with properties different to that of the fluid. Examples include particles of metals, oxides, nitrides, sulphides, polymer particles, particles of inorganic or organic materials, particles of gel, also composite particles, and mixed particles, nano-particles, microparticles, particulate complexes, cells, clusters of cells, bacteria, fungi, virus, clusters of viruses, particles of proteins. Likewise, “particulate containing fluid” means a fluid containing particulates. Examples include cell containing fluids, such as sperm containing fluid.
“Analysis” means determining a qualitative or quantitative characteristic of the particulates in the fluid, for example determining whether the particulates are a homogenous population or a heterogenous population, determining the amount or concentration of particulates, or differentiating or sorting the particulates based on differences. Thus, the term broadly covers analysis of the particulates (i.e. cells) qualitatively or quantitatively, or differentiation or sorting of the particulates based on detected impedance response differences.
“Cells” means any type of cell, including mammalian cells such as sperm, white blood cells, red blood cells, bone marrow cells, immune cells, epithelial cells, nerve cells, pulmonary cells, vascular cells, hepatic cells, kidney cells, skin cells, stem cells, or bacterial and fungal cells and hybridomas, yeast cells, cancerous cells. Generally, the particulate containing fluid contains at least two different types of particulates, for example different cell types, sperm of different sex, sub-populations of the same cell types, the same cell type having different phenotypes, dead and living cells, diseased and non-diseased cells, immature and mature cells of the same kind. The apparatus and methods of the invention may be employed to analyse and/or differentiate and/or separate these different types or phenotype of particulates/cells.
“Different phenotypes” as applied to cells means different populations of cells (i.e. hepatic cells and vascular cells), different sub-populations of the same cell type (i.e. different types of cartilage cells), different phenotypes of the same cell type (i.e. cell expressing different markers, diseased and healthy cells, transgenic and wild-type cells, stem cells at different stages of differentiation).
“X and Y population” as applied to sperm cells means male sperm and female sperm cells.
“Focused stream of particulate containing fluid” means a fluid containing particulates in the form of a focused beam of particulates positioned within a guidance stream. In one embodiment the particulates in the focused beam are focused into a single cell stream arrangement. In one embodiment, in which the particulates have an anisotropic shape, particulates in the focused beam are aligned in the same direction.
“Microfluidic chip” means a chip having at least one microfluidic channel having a cross-sectional area of less than 1 mm2 and a length of at least 1 mm. In one embodiment, the microfluidic chip has at least one microfluidic channel having a cross-sectional area of less than 0.25 mm2. In one embodiment, the microfluidic chip has at least one microfluidic channel having a cross-sectional area of less than 0.01 mm2. In one embodiment, the microfluidic chip has at least one microfluidic channel having a cross-sectional area of less than 0.0025 mm2. In one embodiment, the microfluidic chip has a plurality of microfluidic channels, for example at least 2, 3, 4, 5, 6, 7, 8, 9 or 10 microfluidic channels. In one embodiment, the microfluidic chip has at least one microfluidic channel having a length of at least 1.5 mm. In one embodiment, the microfluidic chip has at least one microfluidic channel has a length of at least 2 mm. In one embodiment, the microfluidic chip has a length of at least 3 mm. In one embodiment, the microfluidic chip comprises a plurality of layers, for example at least 2, 3, 4, 5, 6, 7, 8, 9 or 10 layers.
“AC impedance changes” should be understood to mean changes in impedance detected between the detection electrodes. The changes may include changes in amplitude, phase, or amplitude and phase of the signal.
“In electrical communication with the microfluidic channel” as applied to the electrodes means that the electrodes are in direct contact with the fluids analysed in the microfluidic channel.
“Separation zone” is a part of the microfluidic chip, distal of the detection zone, where particulates in the fluid can be separated in accordance with the results of the characterization of the particulates in the detection zone. The separation zone comprises at least one pair of electrodes, optionally two pairs of electrodes. One or more Voltage/current generators are provided and electrically coupled to the separation electrodes. The or each voltage/current generator is configured to provide a pulse of current across the microfluidic channel. Generally, at least one separation electrode is disposed in the separation zone proximal (upstream) of the bifurcation (splitting) of the microfluidic channel, and at least one separation electrode is disposed either proximal (upstream) of the bifurcation (splitting) of the microfluidic channel or distal of the bifurcation (i.e. in one of the secondary channels. Thus, the separation electrodes may be disposed on opposite sides of the microfluidic channel proximal of the bifurcation (i.e.
The term “anisotropic” refers to being not spherical in overall symmetry of particle's shape or its response to the stimulus used in the apparatus. In the simplest case, this refers to overall shape of the particle (cell). For example, if the particle is elongated, ellipsoidal, bar-shaped or disk-shaped, discoid, this is then described as anisotropic in contrast to a spherical shape particle that is being described as isotropic. However, the overall shape in its own right is insufficient to distinguish between anisotropic and isotropic particles (cells). For example, if a conducting rod (segment of wire) is embedded into an insulating sphere, this forms an anisotropic particle even if the overall shape of the particle is spherical, i.e. isotropic. The reason is that such a particle has different response to the Radio Frequency (RF) electromagnetic field depending on whether it is directed with the length of the rod along the field or perpendicular to the field. The main response to the RF field will be in this case from the metallic rod, this response will be highly anisotropic, the insulating spherical envelope will have little effect on the situation. The same applies to optical response: it will be different depending on the direction of the light incidence and the polarization with respect to the long axis of the rod, again the effect of the isotropic dielectric envelope on the optical response will not alter anisotropic response from the conducting rod. The same applies to the cells. The main contribution to RF signal response from a cell may not come from the exterior periphery of the cell but from its interior features. This depends on the structure of the cell and the RF frequency.
When referring to laminar flow regime, we shall imply the flow conditions that fall under the Stokes regime (−1<Re<˜1000). Re is the Reynolds number defined as Re=ρUH/μ, where μ, U and μ, are the fluid density, the average velocity and dynamic viscosity respectively and H is the characteristic channel dimension. In some cases the effect of particle focusing may still be achieved when Reynolds number is below 1 and therefore the invention is not restricted to the situation of <Re<˜1000. Generally the range of Re values at which the focusing is achieved, also depends on the difference between the density of the liquid and the density of the particles. The greater is the difference, e.g. the heavier are the particles compared to the liquid, the greater is the effective gravity force (difference between the gravity force and the buoyance force) pulling the particles down from the locations defined by the hydrodynamic forces. Therefore, the greater is the difference between the densities, the greater should be the force acting on the particles to achieve effective focusing of the particle's trajectories.
The term “separation of particulates in a fluid into subsets of particulates” should be understood to mean separating a type of particulate within a fluid. The fluid may contain just that particulate or more probably contains at least two subpopulations of particulates. The particulates may be cells, and may be separated into subpopulations based on phenotype, for example cell type, cell health (diseased, normal), viability (dead or alive), sex (X and Y subpopulations), level of differentiation (stem cells). The method employs an electric current to deflect specific cells, provided as a pulse. In one embodiment, the method of the invention employs at least two different current pulses, each configured to deflect a different subpopulation of particulate.
The invention will now be described with reference to specific Examples. These are merely exemplary and for illustrative purposes only: they are not intended to be limiting in any way to the scope of the monopoly claimed or to the invention described. These examples constitute the best mode currently contemplated for practicing the invention.
The detection electrodes normally have electric contact with the interior of the common channel but for some embodiments, especially operating at higher end of frequency range described in this document, the detection electrodes may also be electrically insulated from the interior of the common microfluidic channel. The detection electrodes 5a, 5b, 5c, 5d are connected to the detection circuit. Some detection electrodes are connected to AC voltage or current generators and others are connected to the AC voltage or current detectors. These are not shown in
The separation area 16 shown in
We teach in this patent application that such a symmetric situation will be altered if there is a voltage/current pulse applied between the electrodes. When the voltage is applied between the separation electrodes 16a and 16b, there is current tangential to the flow direction that will introduce non-equivalence into the entrances from the common microfluidic channel and the secondary channel 17a and the same for the secondary channel 17b. This can be considered as additional fluidic impedance or pressure source inserted between the common microfluidic channel and one of the secondary channels. The non-equivalence is results from the following key contributions. Firstly, there is electroosmotic pressure arising in the channel and this pressure will alter the difference between the distribution of flows of liquid into the secondary channels 17a and 17b. Secondly, there is electrophoretic force acting on the particles/cells. This force arises from the ionisation on the particles/cells that results from their interaction with the liquid in which they are immersed. The ionised particle/cell interacts with the electric field between the electrodes 16a and 16b and this produces the force acting on the particle/cell moving predominantly into the secondary channels 17a and 17b. Thirdly, there is dielectrophoretic force. This force results from the fact that the electric field between the electrodes 16a and 16b is not homogeneous and there is a region of gradient electric field. The dielectric permittivity of the particle/cells differs from that the liquid surrounding it. Therefore, there is force acting on the particle/cell that is proportional to the gradient of the electric field and the difference in the dielectric constants of the particle/cell and that of the liquid around it. Depending on the electric permittivity of the cell and the liquid carrying the particles/cells and also on the polarity of the voltage applied between the electrodes 16a and 16b, the particles/cells will be pulled towards the secondary channel 17a or 17b.
The quantitative effects of these three contributions to the force are difficult to describe theoretically. They depend on properties of the liquid, its pH, electric properties of the particle/cell, configuration of the electrodes.
The exact balance of the forces contributing to the overall force should preferably be determined experimentally and can be performed by a person skilled in the art. However, it is important that in any event, the introduction of the electric current between the electrodes as shown in
The voltage/current pulse should be short enough and coordinate with the time of arrival of a specific cell that has been measured in the detection area 5, to the separation area 16.
One example of sets of voltages applied to the four separation electrodes is now described: electrodes 16a and 16c connected to ground potential, electrode 16d connected to positive potential +V and electrode 16b is connected to negative potential −V. Electric fields produced by this set of voltages are shown schematically in
What is common is all these configurations is that all of them create electric field and current that is significantly non-collinear with the direction of the fluid flow in the common microfluidic channels in proximity of the point of bifurcation of the common microfluidic channel into a number of secondary channels.
The voltage values applied to all these electrodes do not need to be constant in time. In a typical operation the voltages applied to the electrodes switch on and off or alter between the set values in order to achieve switching of the flow between different secondary channels. This is explained in further detail using
The voltage applied to the electrodes can be switched on and off on demand. In most preferred embodiments the voltage to the electrodes is supplied in the form of pulses of fixed shape and amplitude: one pulse or one fixed train of pulses is supplied to direct a single particle/cell. In the typical embodiment, the generator/generators supplying the voltages to electrodes are controlled from the output of the circuit detecting the particles/cells. This is shown schematically in
The positions of the electrodes are more clearly shown in
The detection electrodes 5a and 5b detect the change in AC complex impedance between them (between the excitation electrode and signal electrode) resulting from the particles/cells passing through the detection area in between the excitation electrode/electrodes and the signal electrode/electrodes.
Once a particle/cell passes in between a pair of detection electrodes, it obstructs the electric coupling between the excitation electrode and the signal electrode. This changes the voltage induced in the signal electrode. The amplitude and the phase of the voltage induced in the signal electrodes depends on the properties of particle/cell and therefore can be used as a key indicator in identifying the particle (cell) or the subset to which the particle/cell belongs. The detection of this signal is done using the method of demodulation with the help of the lock-in or phase sensitive detection as described in the state of the art section of this document.
The identification of the particle/cell position and alignment should be done in real time within the same time interval as the measurements using the demodulation of the AC impedance change described above.
In order to improve signal-to-noise ratio of the signals induced in the signal electrodes, it is beneficial to use differential amplifiers. For this, a single signal electrode is replaced by two electrodes: signal electrodes 5b and 5d. These are connected to inputs of pre-amplifiers: pre-amplifiers 13 and 13′. The outputs of the pre-amplifiers are connected to a comparator/differential amplifier 14 as shown in
The comparator/differential amplifier 14 amplifies the difference between two inputs which are the outputs of the pre-amplifiers 13 and 13′. The output from the comparator is coupled to a lock-in amplifier or demodulator/phase sensitive detector capable of measuring the signal at the frequency ω of the AC voltage source connected to the two excitation electrodes. The generator connected to the excitation electrodes is not shown in
Once the particle/cell passes in between the excitation electrode and the signal electrode, it will induce signal in the comparator at the frequency ω of the generator connected to the excitation electrodes. The typical amplitude of the output signal from the comparator is shown in
Once the particle/cell is identified/measured in the detection area, it is possible to determine the time required for it to travel from the detection area to the separation area. In
As will be appreciated by those skilled in the art of electronics, the supply of the signal from the AC voltage generator into demodulator, lock-in amplifier/phase sensitive detector as the frequency reference, is important. This is not shown for clarity of the figures.
In a typical embodiment the width of the excitation and signal electrodes is in the range of 0.05 mm to 50 mm. The typical width of the separation electrodes is also in the same range of dimensions.
The detection circuit given in
In relation to the frequency of the voltage source/sources, the signal induced in the signal electrodes by cells/particles passing in front of the electrodes is due to conductance change in the space between the excitations and sensing electrodes and also due to permittivity change in the same space. It should be appreciated that the dielectric permittivity ε(f) of a typical cell is frequency f dependent, and its conductance properties are also frequency dependent. These properties may also depend on other external factors such as pH liquid or temperature as the internal properties of the cells are affected by the conditions outside the cell. The conductance of the liquid bi-layer along the cell surface also depends on the pH of the liquid carrying the cells and its internal structure. It is preferable to choose the AC frequency f of the voltage energising the excitation electrodes such that the dielectric permittivity E at this frequency is significantly different from that of the liquid where the cells are placed. The dielectric permittivity is a complex function that has two components: the real Re ε(f) and imaginary Im ε(f) ones, both of these being functions of frequency. The same applies to conductance. It is preferable to operate the excitation electrodes at the frequency where there is significant difference between the conductance of the particles/cells and that of the liquid carrying these. Such a difference is responsible for the magnitude of the signal induced in the signal electrodes. Since cells have internal fine structure with elements of the structure having their own different dependencies of permittivity, there is generally a rather broad window of frequencies where the signal from the cells/particles can be readily detected. The frequency ω is usually selected in such a window by tuning the frequency to optimise the value of the signal induced in the signal electrodes.
The walls of the microfluidic common microfluidic channel could be made of plastic, glass or other materials, e.g. silicon. The walls of the common microfluidic channel are indicated on
As a brief note of the AC impedance detection, the signals received at each signal electrode in general are complex signals. They are characterised by amplitude and phase or otherwise by real and imaginary parts of the signal. Such complex response arises from a complex circuit in between the excitation and the signal electrode: there is resistive and capacitive coupling between the electrodes. The cells are also characterised by a dielectric constant that has real and imaginary parts. The imaginary part is responsible for the losses in the cell under the influence of the AC electric field. The liquid carrying particles/cells also has real and imaginary dielectric constant. For a given configuration of the electrodes and given characteristic of the liquid and the particles/cells, the importance of the phase contained in the signal is dependent on the excitation frequency ω. There can be a window of frequencies where the phase characteristics of the signal collected from the signal electrode should not be neglected as it contains valuable information helping to identify the particles/cells. This information can be readily collected if the signal detection is done using phase-sensitive detector or lock-in amplifier.
The microfluidic device may also include means for sustaining and control of the fluid flow in the microfluidic channel, mixing the flows of the sample fluid and the sheath fluid, regulating the flows. The flow of fluids in the channels is sustained by a pump or multiple pumps or a pressure source/sources. This could be e.g. one or several UniGo pressure driven pumps from Cellix Limited, Dublin, Ireland. The pump/pumps are not shown in figures of this document for shortness as those skilled in the art of microfluidics are familiar with this aspect. There is a flow of the particle containing fluid 3a enveloped in the flow of the guidance fluid 4a. For shortness, we may call the particles-containing fluid also the sample fluid. We shall also refer to it as the liquid carrying particles/cells. Therefore, it is understood that the liquid carrying particles/cells and the particles/cells themselves form sample fluid. For example, this could be For example, this could be TRISA based buffer which is commonly used for holding sperm cells which might contain concentration of sperm cell from 0.1-100 million sperm cells per millilitre. The particle containing fluid is the fluid that contains the particles of interest. These could be organic or inorganic particles. These could also be alive or dead cells including mammalian cells, sperm cells, yeast cells, particles of biological and non-biological origin etc. In this embodiment, the particles-containing fluid stream (sample fluid) is located at the centre of the guidance fluid 4a flow, i.e. at the centre of the cross-section of the common microfluidic channel 2. Therefore, the guidance fluid 4a performs the function of the sheath fluid focusing the flow of the particle containing fluid into a tighter flow of reduced cross-section. One could construct other embodiments where the sample fluid is guided by the guidance fluid to be positioned not at the centre of the common microfluidic channel but e.g. at a corner of rectangular cross-section of the common microfluidic channel or along its wall and there are certain benefits of such a positioning. The guidance fluid also does not need to envelope the sample fluid all around but could be enveloped e.g. on two or three sides. All these embodiments are included in the present document. To obtain the focusing of the particles-containing fluid 3a by means of the guidance fluid, the microfluidic chip 1 comprises sample microfluidic channel 3 sustaining a flow of particles containing fluid and a guidance microfluidic channel or channels 4 sustaining a flow of guidance (sheath) fluid. Typically the flow rate of the guidance fluid is substantially greater than the flow rate of the sample fluid, i.e. greater by a factor of 2 to 100.
It will be appreciated by those familiar with hydrodynamic focusing that there is no sharp physical boundary between the sample fluid flow and the guidance fluid in the common channel. These two liquids gradually intermix by diffusion and under influence of other forces, along the common microfluidic channel as they move from the point of their mergence downstream. However, over the distance of 0.1-10 millimeters one could readily consider the flow as the flow of two fluids: the sample fluid enveloped by a guidance fluid. The guidance fluid 4a may be interchangeably called the sheath fluid in this document. The linear velocity of the liquid in the sample microfluidic channel and the guidance channel is in the range 0.01 meters/sec to 10 meters/sec. Those skilled in the art appreciate that the linear velocity varies strongly across the channel and normally is greatest at the centre of the channel at least for pressure-driven flows. This compression of the sample fluid by the sheath fluid through the laminar mixing of the streams in one common microfluidic channel, is known as hydrodynamic focusing. The focused flow of the particles containing fluid gradually becomes defocused due to the diffusive movement of the particles perpendicular to the flow direction. The detection area is located at short enough distance away from the point where the sample microfluidic channel merges with the guidance channels, at such a distance where the particles containing fluid still remains focused at the centre of the common microfluidic channel 2. This distance could be in the range 20 to 2000 micrometers.
We refer to particulates/particles throughout this document. This also includes clusters of particulatesparticles. The term cluster describes a small group of particles, for example 2-20 particles linked together by binding forces. For the purpose of this document we shall treat such clusters in the same way as single particles.
Figures of this document show excitation electrodes and signal electrodes making direct contact with the interior of the common microfluidic channel. This does not have to be always the case and one could construct embodiment where some or all of these electrodes are electrically insulated from the interior of the common microfluidic channel. This embodiment could more practical for operation at very high excitation frequencies, e.g. above 10 MHz.
The cross-section of the channel does not need to be rectangular or square. One could have circular, triangular, elliptical channel or indeed channels of various other cross-sections.
It is understood that cells addressed by the above description could be any live or dead cells, and non-mammalian mammalian cells, sperm cells, etc.
The use of hydrodynamic focusing is entirely optional. One may arrange the flow of sample fluid through the common microfluidic channel without any use of the sheath fluid.
All the embodiments shown in this document have separation electrodes positioned on two vertical walls. This is done so only for the clarity of the drawings. For example, with reference to
The foregoing description details presently preferred embodiments of the present invention. Numerous modifications and variations in practice thereof are expected to occur to those skilled in the art upon consideration of these descriptions. Those modifications and variations are intended to be encompassed within the claims appended hereto.
| Number | Date | Country | Kind |
|---|---|---|---|
| 17177624.8 | Jun 2017 | EP | regional |
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/EP2018/067003 | 6/25/2018 | WO | 00 |
