Patent Application
20230408395
The present invention relates generally to point of care (POC) systems and more particularly to bulk fluid processing with AC-osmotic based electrokinetic fluid flows that can be controlled, processed, and automated with integrated cellular and biomolecular detections for POC diagnostics.
The widespread nature of the COVID-19 pandemic has demonstrated the importance of POC diagnostics that allow an end-to-end biosensing capability from bio-sample to diagnostic information. Bridging this sample-to-information paradigm requires addressing two broad and equally important objectives that need to satisfy sensitivity and specificity requirements necessary for POC applications while also being low cost.
The first objective is to automate the first step in the biosensing process that involves extracting and processing the relevant bio-samples (cells/molecules) from the sample selectively. Such sample preparation can involve several steps including dissolution and mixing with several reagents, dilution, and filtering, all of which are critical to the robustness of the assay chemistry, sensing sensitivity, and specificity.
The second step involves the detection of desired substances (cells and molecules of interest) in the processed sample. In the past decade, there have been significant efforts in enabling such low-cost sensing devices. These include quantitative platforms using complementary metal-oxide-semiconductor (CMOS) based integrated circuit technology. Several modalities of bio-molecular sensing including fluorescence-based, magnetic-based, label-free sensing have been demonstrated in prior works across both nucleic acid and protein-based assays. For cells, cytometry has also been demonstrated with magnetic-labels or in a label-free manner.
However, sample fluid preparation is still typically done either manually or with an array of pressure-driven microfluidic channels, connected through a set of tubes to syringe pumps. As a result, while the sensing interface is miniaturized, the rest of the POC system can still be bulky and expensive, thereby severely liming its range of application. In the case of ingestible-based electronics for in-vivo sensing, such pressure driven flow is even more impractical, given the ultra-miniaturized nature of the entire sensing system. Therefore, electronically driven flow becomes attractive to consider for ultra-compact biosensing applications, given its scalability and compatibility with the chip-scale sensing interfaces. While electrically driven droplets and molecular and cell manipulation techniques, such as electro-wetting, electrophoresis, and dielectrophoresis, have been demonstrated in singular systems, these systems do not have the capability to process bulk bio-sample fluids that is required for POC systems.
According to various embodiments, a microfluidic bio-sensing system is disclosed. The system includes at least one semiconductor chip configured to control at least one of electrokinetic fluid flow, cell manipulation and sensing, and bio-molecular sensing by utilizing at least one plurality of electrodes in a microfluidic channel. The semiconductor chip can be configured to control voltage magnitudes on either side of the electrodes to synthesize an electric field in the microfluidic channel for electrokinetic fluid flow. The semiconductor chip can be further configured to control voltage magnitudes on either side of the microfluidic channel to synthesize an electric field in the microfluidic channel for cell focusing. The semiconductor chip can be also configured to create an asymmetric excitation on the plurality of electrodes for cell separation. A top portion of the plurality of electrodes can be designated for voltage excitation and a bottom portion of the plurality of electrodes can be connected to a receiver for cell sensing. Further, a first portion of the plurality of electrodes can be designated for voltage excitation and a second portion of the plurality of electrodes can be connected to a receiver for bio-molecular sensing, where some electrodes of the plurality of electrodes are activated with probe molecules for binding with molecules of interest.
According to various embodiments, a scalable bio-sensing system is disclosed. The system includes an array of microfluidic controllers, where each controller includes at least one semiconductor chip configured to control at least one of electrokinetic fluid flow, cell manipulation and sensing, and bio-molecular sensing by utilizing at least one plurality of electrodes in a microfluidic channel. A semiconductor chip of at least one of the microfluidic controllers is configured to control voltage magnitudes on either side of the plurality of electrodes to synthesize an electric field in the microfluidic channel for electrokinetic fluid flow. A semiconductor chip of at least one of the microfluidic controllers is configured to control voltage magnitudes on either side of the microfluidic channel to synthesize an electric field in the microfluidic channel for cell focusing. A semiconductor chip of at least one of the microfluidic controllers is configured to create an asymmetric excitation on the plurality of electrodes for cell separation. In at least one of the microfluidic controllers, a top portion of the plurality of electrodes is designated for voltage excitation and a bottom portion of the plurality of electrodes is connected to a receiver for cell sensing. In at least one of the microfluidic controllers, a first portion of the plurality of electrodes is designated for voltage excitation and a second portion of the plurality electrodes is connected to a receiver for bio-molecular sensing, where some electrodes of the plurality of electrodes are activated with probe molecules for binding with molecules of interest.
According to various embodiments, a microfluidic bio-sensing system is disclosed. The system includes at least one semiconductor chip configured to control electrokinetic fluid flow, cell manipulation and sensing, and bio-molecular sensing by utilizing pluralities of electrodes in a microfluidic channel. The semiconductor chip is configured to control voltage magnitudes on either side of the electrodes to synthesize an electric field in the microfluidic channel for electrokinetic fluid flow. The semiconductor chip is further configured to control voltage magnitudes on either side of the microfluidic channel to synthesize an electric field in the microfluidic channel for cell focusing. The semiconductor chip is also configured to create an asymmetric excitation on the plurality of electrodes for cell separation. A top portion of the plurality of electrodes is designated for voltage excitation and a bottom portion of the plurality of electrodes is connected to a receiver for cell sensing. Further, a first portion of the plurality of electrodes is designated for voltage excitation and a second portion of the plurality of electrodes is connected to a receiver for bio-molecular sensing, where some electrodes of the plurality of electrodes are activated with probe molecules for binding with molecules of interest.
Various other features and advantages will be made apparent from the following detailed description and the drawings.
In order for the advantages of the invention to be readily understood, a more particular description of the invention briefly described above will be rendered by reference to specific embodiments that are illustrated in the appended drawings. Understanding that these drawings depict only exemplary embodiments of the invention and are not, therefore, to be considered to be limiting its scope, the invention will be described and explained with additional specificity and detail through the use of the accompanying drawings, in which:
The importance of point-of-care (POC) biomolecular diagnostics capable of rapid analysis has become abundantly evident after the outbreak of the Covid-19 pandemic. While sensing interfaces for both protein and nucleic-acid based assays have been demonstrated with chip-scale systems, sample fluid preparation has often been a major bottleneck in enabling end-to-end diagnostics. Typically, such sample handling is either done manually, or through a complex array of microfluidic channels, the flow being precisely controlled via pressure in complex tubing through bulky syringe pumps. Miniaturization of an end-to-end system requires addressing the front-end sample processing, without which, the goal for low-cost POC diagnostics remain elusive.
As such, addressed herein are bulk fluid processing with AC-osmotic based electrokinetic fluid flows that can be fully controlled, processed, and automated by complementary metal-oxide-semiconductor (CMOS) integrated circuits (ICs), allowing large scalability. Here, bulk fluid flow controls are combined with bio-molecular sensing, cell manipulation, cytometry, and separation, all of which are controlled with silicon chips for an all-in-one biosensing device. Shown herein are CMOS controlled pneumatic-free bulk fluid flow with fluid velocities reaching up to 160 μm/s within a microfluidic channel of 100×50 μm2 of cross-sectional area. Dynamic engineering of electric fields is incorporated with electrode arrays to precisely control and electronically focus cell flows for robust cytometry and subsequent separation. A 16-array impedance spectroscopy receiver is also incorporated for cell and label-free protein sensing. The massive scalability of CMOS-driven microfluidics, manipulation, and sensing can lead to a new design space and a new class of miniaturized sensing technologies.
Generally disclosed herein are embodiments for a system and method involving a CMOS-based bio-sensing approach that combines the functionalities of bulk fluid processing and cellular/bio-molecular sensing capabilities in a single handheld platform, as shown generally by
Electrokinetics: AC Osmotic Bulk Fluid Flow with CMOS ICs
Pneumatic-based microfluidic flow with syringe pumps is challenging to scale up due to the complexity and increasing number of pneumatic pumps and microfluidic tubing required for sample preparation in biosensing systems. Electrokinetics can provide a scalable solution through controlling of ionic fluids with electric fields. Here, AC osmosis allows scalable electrolyte motion with voltage levels supported by CMOS ICs.
Fundamentals of Electro-Osmotic Flow
Electro-osmosis provides one of the most popular non-pressure driven flows in microfluidics. The basic principle of the nonlinear flow generated by electro-osmosis relies on a charged electrode that creates non-neutral ionic layers at the electrode-electrolyte interface. Within the interface, a cloud of charged particles creates a diffusion layer, which under the influence of an external electric field, drags the polarizable water molecule tangential to the electrode surface causing the bulk fluid motion.
The principle for the electrodes that create the desired osmotic flow, therefore, focuses on engineering the optimal electric fields that generate the nonlinear drag to allow net fluid flow in one direction. When a charged electrode surface comes in contact to an ionic solution, an electrode double layer forms.
and q is the charge of a single electron, ∈ is the dielectric constant of the medium, kB is Boltzmann's constant, T is temperature in kelvin, ni represents the number of charges for each ion type in a unit volume and zi is the valency of each ion. As an example, in 0.001×PBS solution, the debye length contributed by the dominated ions, Na+ and Cl− (≈0.137 mM for each ion), alone is roughly around 20-30 nm.
In summary, the various principles of the electro-osmosis can be itemized as follows:
(1) Ions in the diffusion layer attract the polarizable water molecules. Therefore, by moving the ions in the diffusion layer, there will be a drag effect on the water molecules that eventually lead to bulk fluid motion.
(2) Manipulation of the ions in the diffusion layer can be realized by creating tangential electric fields close to the electrode surface.
(3) Presence of ions is necessary for such fluid motion. Many bio-samples (blood/saliva/sweat) are ionic, and therefore, such techniques are applicable.
AC Electro-Osmotic Flow
While the previous description provides a general guidance of nonlinear electro-osmotic flow, AC osmosis provides one possible mechanism to induce the electric fields, as shown by
AC osmosis, on the other hand, relies on laterally positioned asymmetric electrode pairs, lowering the voltage requirements that is able to be handled by CMOS ICs. As shown in
Electrode Configuration and Bulk Fluid Flow
As shown in
where Ao=CDL,S(RN−R1)/(R1+RN)(CDL,L+CDL,S), Q=√{square root over (R1RN)}/(R1+RN) (CDL,L+CDL,s) (quality factor), and ω0=(CDL,L−CDL,s)/(CDL,SCDL,L√{square root over (R1RN)}). In addition, the absolute voltage at VN is larger than V1, since RN>R1. This yields a net electric field pointing from the far edge of large electrode towards small electrode during positive AC cycle, as shown in
The velocity of ion (vion) in the diffusion layer can be evaluated from the tangential electric field
(where ΔL is the separation distance between the two locations) as
where μ is the ionic mobility, ∈ is the dielectric constant of the fluid, η is the fluid's dynamic viscosity, ζ is the zeta potential at the double layer (shown in
Given the nature of the electrolyte and ionic concentration, the configuration objective behind choosing the width and gap for the AC osmosis electrodes becomes clear: maximize conversion of electrical energy to kinetic energy of the fluid. This can be achieved by maximizing ΔV at the center frequency ω0, as shown in Equation (1).
Cell Electronic Focusing, Sensing, and Separation
Precise control of the flow of cells is critical for high sensitivity cytometry. In prior works, hydrodynamic force has been mostly deployed to allow focusing of cell motion within a narrow streamline flow, thereby, enhancing overall specificity and sensitivity. This requires additional pneumatic-based microfluidic pumps.
An alternative way of manipulating cells in the microfluidic system is through dielectrophoresis (DEP). DEP relies on creating a net force on a suspended polarizable particle (such as cells) in a non-uniform electric field.
With the ability to precisely engineer the optimal fields with an array of electrodes controlled by custom configured CMOS circuits, the need for any hydrodynamic focusing can be eliminated and replaced with electronic focusing, cell manipulation, and sensing. The force exerting on a dielectric particle in a nonuniform field can be evaluated as:
where P=pV is the induced dipole under an external electric field, V is the volume of the particle, Erms is the root-mean square (RMS) intensity of the electrical field, and fcm is the Clausius-Mossotti factor which quantifies the dipole moment of the particle relative to the solution medium. This factor can be quantified as follows:
where ∈p and ∈m represents the complex permittivity of the particle and the fluid medium. Therefore, with an array of electrodes with independent drive capability, one can engineer the nature of the electric fields within the channel, creating precise positioning of the cells, which allow for high precision cytometry and subsequent separation capabilities, as shown by
Given that both ∈p and ∈m depend on frequency, the direction of the DEP forces can change depending on the frequency of operation, offering another degree of freedom to manipulate cells. For example, when a red blood cell (RBC) is present in isotonic buffer of 8.5% sucrose+0.3% dextrose (≈60 mS/m−1), the inversion of DEP force occurs at around 1 MHz, as shown by
For DEP to operate effectively, the required electric field intensity is reported to be on the order of 105-106 V/m. This indicates that under 5V of voltage excitation, the separation of the DEP electrode within the microfluidic channel is around 50 μm.
With the aforementioned functionalities, disclosed herein is a system of electrodes capable of performing cytometry, cell actuation, and AC osmosis driver on a single glass substrate. As shown in
As shown in
In between the cell focusing and separation electrodes, are four pairs of impedance sensing electrodes that are able to capture cell flow in microfluidic channel in real-time. The top four electrodes are designated for voltage excitation, where the bottom four are directly connected to the receiver inputs. To simplify the configuration, all of the DEP and impedance sensing electrodes have a width of 40 μm and separation of 30 μm, as shown later in
Impedance Spectroscopy Sensing
Built on top of the AC electro-osmosis bulk fluid driver and DEP cell actuation, is a 16-array impedance spectroscopy receiver based on a direct conversion architecture. The purpose of the impedance spectroscopy sensing is to provide the impedance measurements in real-time at the electrode-electrolyte interface. This sensing modality is employed for cytometry sensing, and for label-free protein bio-assay in this multi-modal bio-sensing platform. As shown in
The voltage source, Vx, shown in
Bio-Molecular Sensing
Bio-molecular sensing, compared to cell sensing, focuses on detecting a bio-molecular binding event, such as protein-protein interactions and complementary nucleic-acid binding activities. In labelled detection, extra labelling molecules or particles (such as fluorophores or magnetic beads) are relied on to produce a signal which is then detected by a transducer. However, for label-free detection, the bio-molecular binding event is detected directly through methods such as but not limited to impedance spectroscopy and optical resonance. Here, impedance spectroscopy is deployed, as it can be realized through compact CMOS technology without any external instrumentations and is compatible with compact POC diagnostic applications.
The sensor electrodes are first activated with probe molecules that will only bind with specific target molecules. Therefore, when a bio-molecular binding event occurs on the electrode surface, it causes a tiny impedance change which can be detected by the on-chip current sensor, resulting in a positive signal. The on-chip impedance spectroscopy, in general, can be adapted to any bio-molecular detection in a compact form, where the probe molecule can be customized to any bio-molecular sensing application, such as protein antibodies in immunoassay or nucleic-acid sequence probes for viral DNA/RNA detection in a sample.
Microfluidic Device Fabrication and Assembly
The microfluidic flow assembly is shown in
Electrode Fabrication
Here, borosilicate float glass is used as substrate, as a nonlimiting example. A silicon wafer can also be used as a substrate, glass is known to best bind with PDMS. The electrode pattern can be fabricated through any standard photolithography procedure. First, a 10 nm layer of Chrome (Cr) is deposited as an adhesive layer, followed by a 100 nm of gold (Au) layer deposition. Other adhesive metals such as but not limited to tungsten, niobium, and titanium can be used in alternative embodiments. Au is chosen as the interface between the electrode-electrolyte interface, to take advantage of its inert chemical nature with ionic liquid. The deposition can be conducted with an e-beam evaporator. Next, the metal layer was removed with photoresist, and the electrodes were formed by a lift-off process. The resulting glass-metal interface is demonstrated at the bottom of
Microfluidic Channel Fabrication and Assembly
Polydimethylsiloxane (PDMS) microfluidic channels can be fabricated through soft photolithography. Glass or SU-8 can also be used as the material for the microfluidic channel in alternative embodiments, as nonlimiting examples. Initially, a SU-8 2025 layer is spin-coated at 3000 rpm on a silicon wafer and soft-baked at 65° C. for 1 min and 95° C. for 6 min. The wafer is then exposed to UV and is post-exposure baked at 95° C. for 6 min. After that, the wafer is developed using a SU-8 developer, followed by a wash with isopropanol and drying. The mold wafer was then used for PDMS casting with curing temperature set to 80° C. for one hour. Finally, the PDMS was peeled off and punched with 1.5 mm-diameter holes for fluid inlets and outlets.
Here, the microfluidic channel has a height of 50 μm and is aligned over the gold electrode device using a custom aligner that includes 3-axis micromanipulators and a microscope. Finally, the PDMS channels are covalently bonded to the glass substrate using UV/O3 cleaner for 20 min. The device is then incubated at 95° C. for 20 min on a hot plate to increase the bonding strength. Finally, the complete channel-electrode interface is shown at the top of
It should be noted that the fabrication method described herein is based on the prototype developed (glass wafer electrodes and PDMS microfluidic channel) and not intended to be limiting. For instance, the fabrication method will differ based on different materials such as SU-8 or glass-based microfluidic channels, and is generally understood by those skilled in the art. The electronic interface should work with most fabricated sensors as long as there are electrodes extending out from the substrate.
Circuits and System Implementation
This section describes an on-chip architecture for controlling AC-osmotic flow, cell flow, and sensing with integrated impedance spectroscopy. The chip is configured and fabricated in a 65-nm LP bulk CMOS process. This particular process is chosen to optimize the performance at low power operation for POC applications but is not intended to be limiting. Other process nodes can be used as well to achieve the same functionality.
AC-Osmosis and DEP Driver Architecture
To provide a driving signal for both the AC electro-osmosis and DEP electrodes, the system includes a programmable signal generator. On-chip high voltage drivers are employed for the DEP system (>5 Vpp). The signal is further boosted with off-chip drivers for swings higher than 10 Vpp for AC osmosis. While the boosting was achieved off-chip for AC-osmosis in this exemplary chip, this can be easily integrated in a longer node CMOS process with higher breakdown voltage limits.
Programmable Square Wave Function Generator:
The core of the square wave function generator includes a 20 MHz three-stage ring oscillator and a 16-bit synchronous counter. The 16-bit counter can be periodically reset through the comparators, where the resetting time can be digitally controlled through a SPI interface, as shown in
High-Voltage Level Shifter and Driver:
The output of the 1.2V square wave is first level shifted to a 3.3V reference (realized with thick-oxide devices), as shown in
The HV level-shifter, in particular, uses a stacked topology with thick oxide transistors and with self-biasing circuitry to simplify the overall design and complexity, as shown in
As shown in
Next, both Vinp and Vinn are fed into the HV driving stage together, which also includes three vertical branches of stacked transistors. The first two vertical branches of stacked transistors on the left of the HV driver are necessary to provide proper biasing voltages for the stacked transistors at the output (VDEP).
In summary, the output of the HV driver is capable of providing square wave driving signal between 5-9V. Two HV drivers generate a differential signal pair that drives the DEP focusing and separation electrodes, that is sufficient to cover the minimum electric field intensity (105 V/m) required for DEP electrodes with 50 μm of separation.
Impedance Spectroscopy Receiver Architecture
The purpose of the direct conversion receiver, as aforementioned, is to provide real-time impedance sensing for the cytometry in the microfluidic channel and for protein sensing, as shown in
Excitation Path:
The excitation signal is generated internally through the divide-by-2 I/Q clock generator, shown in
The signal is then high pass filtered through a combination of on-chip metal-oxide-metal (MOM) capacitor (9 pF) and resistor (10 MΩ poly resistor) to pin the output DC voltage to VSOL, as shown in
Trans-Impedance Amplifier (TIA):
An op-amp based TIA is employed with a RFB of 100 kΩ in feedback to achieve a bandwidth of 2 MHz. The op-amp topology employed in the TIA is similar to the driver op-amp in the excitation path. However, in this configuration, the length of the input differential pair is significantly increased to 1 μm to minimize the 1/f noise. Since the subsequent passive mixer translates the noise spectrum, the output voltage noise of the receiver path has a significant contribution from the VGA and LPF due to the low signal bandwidth (10 kHz) and presence of 1/f noise of the later stages.
Passive Mixer, Differential-To-Single-Ended VGA and LPF:
A double-balanced passive mixer is used to minimize power consumption, LO feed-through, and maximize the conversion gain, as shown in
To bias all the analog components, the chip deploys a constant-gm self-biasing circuit in order to minimize the usage of external biasing pads. As shown in
System Noise Analysis
The noise of the system is primarily contributed by a combination of the input TIA, mixers, VGA, and the low pass filters, as shown in
As an example, for an AC osmotic flow rate of 100 μm/s (to be illustrated in the measurement section), the time taken for a cell to pass through one sensing electrode of width 40 μm is ˜0.25 s. To capture this dynamic with high sample rate (to identify and classify different cells, if needed), the integration time is kept below 0.1 ms, or in effect, the integration is set around 10 kHz by the LPF. Therefore, the 1/f noise of the circuitry after the frequency translation of the mixer becomes critical.
TIA and Mixer Noise:
The output noise of the TIA around the frequency of analysis (f0≈500 kHz) gets translated down to the signal frequency at baseband after mixing. As show in
V
n,opTIA
2
where
In addition, assuming fast switching with 50% duty cycle in the double-balanced passive mixer, the output noise spectrum contributed by the mixer itself, can be approximated as 8kTRL, where RL is the equivalent output loading resistance of the mixer. Combining TIA noise, the net spot noise at the output of the receiver, due to the TIA and the mixer circuits can be evaluated to be 0.32 pV2/Hz. Within the LPF bandwidth of 10 kHz, the total noise contribution due to TIA and mixer combined can be calculated to be ≈57 μV of RMS noise voltage.
Programmable VGA and LPF Noise:
As mentioned before, the 1/f noise portion in VGA is critical in the total noise contributions. Shown in
V
n,oVGA
2
where AVGA is the gain of VGA and it is set to 20 dB. The input-referred noise spectrum density (
The receiver output noise due to the active 4th order LPF (Vn,oLPF2) can be evaluated by considering the transfer functions from each noise source within LPF to the output node, as demonstrated in
Analysis of the noise sources show that the dominant noise source in the receiver chain, as predicted, arises from the 1/f noise in the VGA. The signal acquisition path for both cell and protein sensing is co-configured with the electrode and the receiver configuration, resulting in measured SNR of 10 dB and higher for cell sensing measurements, to be described further below. While these metrics can be further optimized, the achievable levels of sensitivity will allow for demonstrating the principles of multi-modal functionalities of the presented platform.
Measurement Results
This section presents the overall packaging, electrical performance, CMOS-driven AC-osmotic flow, cell manipulation, and bio-sensor measurements according to an embodiment of the invention.
CMOS Packaging and PCB System Implementation
The overall system is shown in
The bio-sensing platform can be powered by two sets of AA batteries.
Electrical Characterization of Impedance Sensing Chip
To characterize the sensitivity of the receiver, the receiver output is connected to an off-chip low-noise pre-amplifier (for example, Stanford SR560) and subsequently to a spectrum analyzer (for example, R&S FSW) to acquire the output noise voltage spectrum. The measured input-referred current noise is illustrated in
To measure the linearity of the receiver, a 10 kΩ resistor is connected between the excitation path and the TIA input. The VGA is set to unity gain for maximum linearity. As shown in
In summary, the input-referred noise and linearity performance is compared against other state-of-the-art CMOS impedance biosensors in the table in
DEP and AC Osmosis Driving Signals
To test the high voltage driving capabilities across frequency, a 16-bit code is sent to the chip that sets the output frequency.
AC Electro-Osmotic Flow Characterization
To characterize the AC-osmotic fluid flow, the electrodes are driven and then the fluid flow is measured on the other side of the microfluidic channel. This is to observe the bead velocity and to eliminate the effect of DEP on the beads, as shown in
In summary, the fluid driving capability is measured in following order. First, the 1 μm polystyrene beads (for example, Sigma-Aldrich LB11) is mixed with 0.001×PBS. Then, the fluid mixture is injected in the microfluidic channel. Next, a 100 kHz driving frequency is turned on from the CMOS chip with varying amplitude. Finally, the fluid velocity is determined by measuring the time that the 1 μm beads takes to travel across 100 and 200 μm electrodes respectively.
In summary, a balance between the ionic concentration, frequency, and the electrode dimension is experimentally optimized and summarized in
Cell Focusing and Positioning with Engineered Fields
As described earlier, the chip architecture has the ability to engineer the electric fields within a cell flow to precisely control their position for sensing and subsequent separation. This functionality is characterized through electronic actuations and is calibrated with video from the microscope, as shown in
Controlling the balance of the excitations on the two sides of electrodes, one can control the lateral displacement of the cells. As shown in
Cytometry Measurement Using Polystyrene Beads
The cytometry impedance measurement is first performed using two types of polystyrene beads: Sigma-Aldrich Supelco 72986 (10 μm) and 74491 (20 μm). First, the beads are diluted from the stock aqueous solution into 0.001×PBS solution buffer with dilution ratio between 1-0.1%. This ratio is experimentally obtained to prevent clogging in microfluidic while maintaining the best measurement visuals under microscope.
The diluted beads are then injected into the microfluidic channel and controlled electronically for accurate impedance measurements.
Cytometry Measurement of Yeast Cell
To demonstrate the cytometric sensing capabilities, the performance is measured with cultured yeast cells (for example, Saccharomyces cerevisiae). Similar to the previous sample preparation, the yeast cells are suspended in the same 0.001×PBS solution. The sizes of the yeast cells are reported in the range of 5-10 μm. Therefore, ensuring consistent focusing of the cell in the channel is a critical aspect in the measurement.
As shown in
Protein Measurement
Finally, protein bio-assay measurement capability is also disclosed herein. The measurement is performed on separate gold electrode platform on a silicon substrate and can be easily integrated in the same glass substrate as the AC osmosis and DEP system electrodes.
As shown in
First, the impedance gold electrodes are immersed in acetone, isopropanol (IPA), and Milli-Q water and then sonicated for 3 min. The electrodes were dried with a nitrogen stream and placed in a UV-O3 cleaner (for example, BioForce Nanoscience, USA) for 30 min. Finally, they were rinsed with IPA and dried with nitrogen stream. A mixed self-assembled monolayer (SAM) was formed ex-situ by incubating the gold electrodes overnight at room temperature with a mixed solution of alkanethiols (for example, 16-mercaptohexadecanoic acid (MHDA) and 11-mercaptoundecanol (MUOH)) in ethanol. The electrodes were then thoroughly rinsed with ethanol and dried with nitrogen stream.
Before applying the target cytolysin protein, the carboxylic group of MHDA were activated as carbodiimide esters by incubation of a mixed solution of EDC/sulfo-NHS (0.2 M/0.05 M) in MES buffer for one hour, and then rinsed with water and dried. The electrodes were incubated with a solution of antibody (50 μg/mL) in PBS buffer pH 7.4 overnight at room temperature. The electrodes were rinsed with PBS buffer to remove unbound antibody and dried with nitrogen stream. The uncovered gold surface was blocked with a 3% bovine serum albumin (BSA) to avoid non-specific bonding. The final surface chemistry is illustrated in
The protein measurements are then performed with varying target cytolysin concentrations.
Finally, the performance for the proposed system is summarized in the table in
As such, generally disclosed herein are embodiments for an approach to combine CMOS-based electrokinetic microfluidics with multi-modal cytometry and protein bio-essay sensing array platform for POC applications. Specifically, embodiments focus on realizing pneumatic-free microfluidic drivers with integrated cell actuation and impedance sensing capabilities for both cytometry and label-free protein sensing that are compatible with standard bioassay protocols. Shown herein were AC-osmotic bulk fluid flows with optimized electrode geometries and dynamic engineering of electric fields to precisely control, sense, and separate cells for cytometry. Demonstrated herein was 16-element impedance spectroscopic receivers on-chip for real-time cell sensing and protein assays. The approaches can be further optimized to yield
better performance but demonstrated overall herein are pathways toward pneumatic-free complex biosensing platforms and ultra-miniaturization for in-vitro and in-vivo applications.
It is understood that the above-described embodiments are only illustrative of the application of the principles of the present invention. The present invention may be embodied in other specific forms without departing from its spirit or essential characteristics. All changes that come within the meaning and range of equivalency of the claims are to be embraced within their scope. Thus, while the present invention has been fully described above with particularity and detail in connection with what is presently deemed to be the most practical and preferred embodiment of the invention, it will be apparent to those of ordinary skill in the art that numerous modifications may be made without departing from the principles and concepts of the invention as set forth in the claims.
This application claims priority to provisional application 63/116,226, filed Nov. 20, 2020, which is herein incorporated by reference in its entirety.
This invention was made with government support under Grant No. ECCS-1711067 awarded by the National Science Foundation. The government has certain rights in the invention.
| Filing Document | Filing Date | Country | Kind |
|---|---|---|---|
| PCT/US2021/060021 | 11/19/2021 | WO |
| Number | Date | Country | |
|---|---|---|---|
| 63116226 | Nov 2020 | US |
