CN111474106A - Method and system for determining mechanical properties of biological cells or biological cell-like particles - Google Patents

Abstract

The present application provides a method for determining a mechanical property of a biological cell or biological cell-like particle, comprising the steps of: passing a fluid comprising the biological cells or biological cell-like particles through a microfluidic channel in a first direction; capturing at least a portion of the biological cells or biological cell-like particles with optical tweezers in a second direction transverse to the first direction during passage of the fluid through the microfluidic channel; collecting an image of the biological cells or biological cell-like particles through the microfluidic channel; and determining a mechanical property of the biological cell or the biological cell-like particle based on the amount of deformation of the biological cell or the biological cell-like particle captured by the optical tweezers in the image. The application also provides a system for implementing the method.

Description

Method and system for determining mechanical properties of biological cells or biological cell-like particles

RELATED APPLICATIONS

Priority of united states provisional application No. 62/918,234 entitled "An Optofluidic stretchers for biological Cells and Soft Particles" filed on 23/1/2019, the contents of which are incorporated herein by reference in their entirety.

Technical Field

The present application relates generally to the fields of biomechanics, optics, and hydrodynamics. In particular, the present application provides methods and systems for determining mechanical properties of biological cells or biological cell-like particles that combine optical tweezers technology and microfluidic technology.

Background

The mechanical properties (also called mechanical properties) of biological cells can be used as intrinsic, unmarked biomarkers that can indicate physiological and pathological changes of the cells [1 ]. Alterations in cellular deformability are associated with alterations in the basal cytoskeleton [2], which have been found to be associated with a wide range of functional changes in cells, including differentiation [3], apoptosis [4], disease transformation [5] and drug response [6 ]. In particular, there is increasing evidence that a decrease in the deformability of human Red Blood Cells (RBCs) is a symptom of various human diseases, such as malaria infection [7], diabetes [8] and sickle cell anemia [9 ]. For plasmodium falciparum malaria, recent experiments have shown that the membrane hardness of the parasitized red blood cells can increase by as much as 10-fold [10 ]. These findings suggest that characterization of the deformability of cells may provide useful information for distinguishing healthy and unhealthy cells, which in turn may be a consideration in disease diagnosis.

In order to study the pathophysiological importance of cell deformability, the mechanical properties of cells have attracted extensive research interest during the last decades. Several techniques have been developed to assess the mechanical deformability of single cells. Reported characterization techniques include micropipette aspiration [11], atomic force microscopy [12], and optical tweezers technology requiring high refractive index microbeads (e.g., silica or polystyrene microbeads) attached to both ends of a cell for mechanical loading [13 ]. These techniques involve direct physical contact between the solid surface and the cell being measured, which may alter the natural behavior of the cell and even damage the cell during the measurement process. Another practical problem is that these techniques are based on static test conditions, which means that the characterization throughput is low. In fact, high throughput assays are more valuable and can rapidly obtain characterization or screening results for a plurality of cells in order to obtain statistically relevant results based on the inherent heterogeneity of biological cells.

Applying optical tweezers directly on cells without bead attachment is an ideal method for non-contact, non-invasive and light-induced cell deformation. Standard optical tweezers comprise a highly focused gaussian laser beam capable of applying optical forces on the order of piconewtons on micro-sized transparent objects for optical trapping and manipulation [14 ]. The optical force of the optical tweezers can be decomposed into two components, i.e., the optical gradient force attracts a high refractive index object to the center of the waist of the light beam where the intensity of the optical field is highest, and the scattering force attracts the object along the longitudinal light beam propagation direction. For a highly focused laser beam, the optical gradient force component in the longitudinal direction can balance the scattering forces, forming a conventional single laser beam tweezer.

The cell can be deformed using dual optical tweezers by applying two focused laser beams propagating in parallel directly at both ends of the cell, mechanically stretching the cell in the direction of separation of the two beams. Another type of optical cell stretching involves two counter-propagating diverging beams emanating from two optical fibers, where the cell is optically trapped and stretched in the longitudinal direction of the two laser beams [16 ].

Microfluidics has gained increasing research interest as an attractive tool for studying the mechanical properties of cells. Microfluidic devices contain microfluidic channels, typically on the micrometer and millimeter scale, with inlets and outlets to connect with the external environment, where fluids and biological samples can be precisely controlled and studied [17 ]. One advantage of microfluidic devices for biological cell analysis is the high characterization throughput achieved by using fluids to continuously deliver cell samples. Another attractive feature of microfluidic devices is the small volume required for analysis, which is attractive for experimental chip applications. Other advantages of microfluidic devices are mature manufacturing processes and materials. Most microfluidic devices are fabricated using well-established lithographic methods employed in the microelectronics industry, allowing for accurate and repeatable fabrication processes. The materials used to fabricate the microfluidic channels are typically transparent, such as Polydimethylsiloxane (PDMS) or glass, which can be optically imaged for cellular analysis.

Microfluidic devices have been developed to study cell deformation using fluid dynamics induced by fluid flow. A typical way is to deform when a cell flows at high velocity through a straight narrow microfluidic channel with cross-sectional dimensions close to the size of the cell [18 ]. When interacting with a non-uniformly distributed fluid in a narrow microfluidic channel cross-section, the cell deforms under fluid stress.

The inventors of the present application have also considered the following documents in developing the various inventions of the present application:

1, C.T, L im, M.Dao, S.Suresh, C.H.Sow, and K.T.Chew in "L image formation of using laser tracks," Acta materials 52(7), 1837-.

S. Rangult-Grenier, M.T.Wei, J.J.Bai, A.Chiou, P.P.Bareil, P. L. Duval, and Y.Sheng in "Dynamic degradation of red blood cells in dual-trap optical tweezers," Opti.Express 18(10),. 10462-10472(2010), shows a bifocal forceps that can expand osmotically distended red blood cells by applying optical force directly across the cell without the need for microbeads.

3, D.S. Moura, D.C. Silva, A.J. Williams, M.A. Bezerra, A.Fontes, and R.E.de Araujo in "Automatic time evaluation of red blood cell electrophoresis" Rev.Sci.Instrum.,86(5),053702(2015) shows cell spreading using motorized stage controlled single tweezers for RBC capture in a fluidic chamber. In this work, RBC deformation was achieved by using optical tweezers to drag the captured cells through a static fluid within a fluid chamber at a defined velocity. The captured cells are deformed by the fluid stress.

4. Us patent 6067859 discloses optical cell stretching based on two counter-propagating diverging laser beams from two optical fibers for deforming the cells.

B.B. L incorn, F.Wottawah, S.Schinginger, S.Ebert, and J.Guck in "High-throughput biological assays with an optical molecular tester," Methods cell biol.83,397-423(2007) further developed their prior art in U.S. Pat. No. 6067859 by combining optical cell stretching with microfluidic channels for continuous cell sample delivery.

K.B.Roth, K.B.Neeves, J.Squier, and D.W.Marr in "High-throughput patent optical bench for mechanical characterization of blood cells," Cytometry Part A89 (4), 391-. In this method, a linear diode bar laser is used as a laser source, and a cylindrical lens is used to project a pattern into the microfluidic channel. When the cells overlap the linear beam, they elongate in the linear direction due to the outward optical stress at the ends of the cells aligned with the linear beam pattern.

Z.Yao and A.W.Poon in "Optical lattice-based cell shaping and transforming using integrated vertical multimode-interference waveguides," C L EO: Science and Innovations, STh3J.6(2018), reported a beam shaping technique using SU8 to fill vertically embedded multimode interference (MMI) waveguides for cell stretching.

O.otto, p.rosendahl, a.mietke, s.golfier, c.herold, d.klaue, s.girardio, s.palgiara, a.ekpenyong, a.jacobi, and m.wobus in "Real-time denaturing assay" on-the-fly cell mechanical patterning, "nat.methods, 12(3),199(2015), show microfluidic-based cell shape change counts that can be used for continuous cytomechanical characterization. In this work, a bottleneck microfluidic channel design was used, which deformed cells by having cross-sectional dimensions comparable to the size of the cells when the cells were flowing through the narrow part of the microfluidic channel at high speed. When interacting with a non-uniformly distributed fluid in the cross-section of a narrow microfluidic channel, cells deform under fluid stress. The characterization flux was reported to be about 100 cells/sec.

D.r.gossett, t.k.henry, s.a. L ee, y.ying, a.g. L indegren, o.o.yang, j.rao, a.t.clark, and d.di Carlo in "hydro dynamic geometry for large cell signalling," proc.natl.acad.sci.u.s.a.109(20),7630-7635(2012), demonstrates a microfluidic-based cell stretching in which cells are slowed down and deformed by fluid forces at the stagnation point of the fast extension flow of a cross-shaped microfluidic channel.

Summary of The Invention

In a first aspect, the present application provides a method for determining a mechanical property of a biological cell or biological cell-like particle, comprising the steps of:

passing a fluid comprising the biological cells or biological cell-like particles through a microfluidic channel in a first direction;

capturing at least a portion of the biological cells or biological cell-like particles with optical tweezers in a second direction transverse to the first direction during passage of the fluid through the microfluidic channel;

collecting an image of the biological cells or biological cell-like particles through the microfluidic channel; and

determining a mechanical property of the biological cell or biological cell-like particle based on the amount of deformation of the biological cell or biological cell-like particle captured by the optical tweezers in the image.

The term "biological cell-like particle" as used in this application refers to a particle having physical properties similar to biological cells, e.g., a soft, insulating, micron-sized particle. Examples of biological cell-like particles include, but are not limited to, phospholipid membrane vesicles.

In some embodiments, the optical refractive index of the medium of the fluid is lower than the biological cells or biological cell-like particles.

In some embodiments, the cross-section of the microfluidic channel is sufficiently large that the biological cells or biological cell-like particles do not physically contact the walls of the microfluidic channel to generate a deforming force when passing through the microfluidic channel.

In some embodiments, the optical tweezers are highly focused laser beams with a gaussian distribution.

In some embodiments, the spectral window of the optical tweezers is permeable for one or more of:

the biological cell or biological cell-like particle,

the medium of the fluid is a mixture of,

a wall of the microfluidic channel.

In some embodiments, the optical tweezers are periodically modulated to periodically capture and release the biological cells or biological cell-like particles. In some embodiments, the optical tweezers are periodically turned on and off.

In some embodiments, the modulation is implemented by a modulator.

In some embodiments, the frequency of the modulation is at a hertz level, e.g., 1-5 Hz.

In some embodiments, the method uses a plurality of optical tweezers.

In some embodiments, the plurality of optical tweezers is produced by an interferometer configured with a single laser source.

In some embodiments, the plurality of optical tweezers is generated by a plurality of laser sources.

In some embodiments, the plurality of optical tweezers is arranged in an optical tweezers array, such as a one-dimensional optical tweezers array or a two-dimensional optical tweezers array.

In some embodiments, the array of optical tweezers is produced using a multimode interference waveguide with a single laser source.

In some embodiments, the array of optical tweezers is produced using a spatial light modulator with a single laser source.

In some embodiments, the focal point of the optical tweezers is set such that cells captured by the optical tweezers are away from the walls of the microfluidic channel.

In some embodiments, the images are acquired using a microscope objective and a digital camera, such as a long working distance microscope objective and a CCD digital camera.

In some embodiments, the image is acquired in a third direction transverse to the first direction.

In some embodiments, images are collected before, while, and after the biological cells or biological cell-like particles are captured by the optical tweezers.

In some embodiments, the dimension (e.g., path length) of the biological cells or biological cell-like particles captured by the optical tweezers in the image along and perpendicular to the first direction is measured and the ratio of the two is used to characterize the amount of deformation.

In a second aspect, the present application provides a system for determining a mechanical property of a biological cell or biological cell-like particle, comprising:

a microfluidic device comprising a microfluidic channel for passing a fluid comprising the biological cells or biological cell-like particles in a first direction;

optical tweezers providing means for capturing at least a part of said biological cells or biological cell-like particles with optical tweezers in a second direction transverse to said first direction during passage of said fluid through said microfluidic channel;

an image acquisition device for acquiring images of the biological cells or biological cell-like particles passing through the microfluidic channel; and

and the analysis device is used for measuring the mechanical property of the biological cells or the biological cell-like particles based on the deformation quantity of the biological cells or the biological cell-like particles captured by the optical tweezers in the image.

In some embodiments, the optical refractive index of the medium of the fluid is lower than the biological cells or biological cell-like particles.

In some embodiments, the cross-section of the microfluidic channel is sufficiently large that the biological cells or biological cell-like particles do not physically contact the walls of the microfluidic channel to generate a deforming force when passing through the microfluidic channel.

In some embodiments, the optical tweezers are highly focused laser beams with a gaussian distribution.

In some embodiments, the spectral window of the optical tweezers is permeable for one or more of:

the biological cell or biological cell-like particle,

the medium of the fluid is a mixture of,

a wall of the microfluidic channel.

In some embodiments, the optical tweezers providing apparatus further comprises a component for periodically modulating said optical tweezers, thereby periodically capturing and releasing said biological cells or biological cell-like particles. In some embodiments, the assembly is used to periodically activate and deactivate the optical tweezers.

In some implementations, the component that periodically modulates the optical tweezers is a modulator.

In some embodiments, the frequency of the modulation is at a hertz level, e.g., 1-5 Hz.

In some embodiments, the optical tweezers providing apparatus can be used with a plurality of optical tweezers.

In some embodiments, the plurality of optical tweezers is produced by an interferometer configured with a single laser source.

In some embodiments, the plurality of optical tweezers is generated by a plurality of laser sources.

In some embodiments, the plurality of optical tweezers is arranged in an optical tweezers array, such as a one-dimensional optical tweezers array or a two-dimensional optical tweezers array.

In some embodiments, the array of optical tweezers is produced using a multimode interference waveguide with a single laser source.

In some embodiments, the array of optical tweezers is produced using a spatial light modulator with a single laser source.

In some embodiments, the focal point of the optical tweezers is set such that cells captured by the optical tweezers are away from the walls of the microfluidic channel.

In some embodiments, the image capture device comprises a microscope objective and a digital camera, such as a long working distance microscope objective and a CCD digital camera.

In some embodiments, the image acquisition is configured to acquire images in a third direction transverse to the first direction.

In some embodiments, the image acquisition is configured to acquire images of the biological cells or biological cell-like particles before, during, and after capture by the optical tweezers.

In some embodiments, the analysis device is configured to measure a dimension (e.g., path length) of the biological cells or biological cell-like particles captured by the optical tweezers in the image along the first direction and perpendicular to the first direction, and to use the ratio of the two to characterize the amount of deformation.

As non-limiting examples, the present application provides the following embodiments:

1. a method for determining a mechanical property of a biological cell or biological cell-like particle, comprising the steps of:

passing a fluid comprising the biological cells or biological cell-like particles through a microfluidic channel in a first direction;

capturing at least a portion of the biological cells or biological cell-like particles with optical tweezers in a second direction transverse to the first direction during passage of the fluid through the microfluidic channel;

collecting an image of the biological cells or biological cell-like particles through the microfluidic channel; and

determining a mechanical property of the biological cell or biological cell-like particle based on the amount of deformation of the biological cell or biological cell-like particle captured by the optical tweezers in the image.

2. The method of embodiment 1, wherein the optical refractive index of the medium of the fluid is lower than the biological cells or biological cell-like particles.

3. The method of embodiment 1 or 2, wherein the cross-section of the microfluidic channel is sufficiently large that the biological cells or biological cell-like particles do not physically contact the walls of the microfluidic channel to generate a deforming force when passing through the microfluidic channel.

4. The method of any of embodiments 1-3, wherein the optical tweezers are highly focused laser beams with a gaussian distribution.

5. The method of any of embodiments 1-4, wherein the optical tweezers' spectral window is transparent to one or more of:

the biological cell or biological cell-like particle,

the medium of the fluid is a mixture of,

a wall of the microfluidic channel.

6. The method according to any of embodiments 1-5, wherein the optical tweezers are periodically modulated, e.g. periodically switched on and off, thereby periodically capturing and releasing the biological cells or biological cell-like particles.

7. The method of embodiment 6, wherein said modulating is performed by a modulator.

8. The method of embodiment 6, wherein the frequency of modulation is at a hertz level, e.g., 1-5 Hz.

9. The method of any of embodiments 1-8, wherein a plurality of optical tweezers are used.

10. The method of embodiment 9, wherein the plurality of optical tweezers is generated by an interferometer configured with a single laser source.

11. The method of embodiment 9, wherein the plurality of optical tweezers is generated by a plurality of laser sources.

12. The method of embodiment 9, wherein the plurality of optical tweezers are arranged in an optical tweezers array, such as a one-dimensional optical tweezers array or a two-dimensional optical tweezers array.

13. The method of embodiment 12, wherein the array of optical tweezers is created using a multimode interference waveguide with a single laser source.

14. The method of embodiment 12, wherein the array of optical tweezers is generated using a spatial light modulator with a single laser source.

15. The method of any of embodiments 1-14, wherein the focal point of the optical tweezers is set such that cells captured by the optical tweezers are away from the walls of the microfluidic channel.

16. The method of any of embodiments 1-15, wherein the image is captured using a microscope objective and a digital camera, such as a long working distance microscope objective and a CCD digital camera.

17. The method of any of embodiments 1-16, wherein the image is acquired in a third direction transverse to the first direction.

18. The method of any of embodiments 1-17, wherein images of the biological cells or biological cell-like particles are collected before, while, and after capture by the optical tweezers.

19. The method of any of embodiments 1-18, wherein the dimension (e.g., diameter length) of the biological cells or biological cell-like particles captured by the optical tweezers in the image along and perpendicular to the first direction is measured and the ratio of the two is used to characterize the amount of deformation.

20. A system for determining a mechanical property of a biological cell or biological cell-like particle, comprising:

a microfluidic device comprising a microfluidic channel for passing a fluid comprising the biological cells or biological cell-like particles in a first direction;

optical tweezers providing means for capturing at least a part of said biological cells or biological cell-like particles with optical tweezers in a second direction transverse to said first direction during passage of said fluid through said microfluidic channel;

an image acquisition device for acquiring images of the biological cells or biological cell-like particles passing through the microfluidic channel; and

and the analysis device is used for measuring the mechanical property of the biological cells or the biological cell-like particles based on the deformation quantity of the biological cells or the biological cell-like particles captured by the optical tweezers in the image.

21. The system of embodiment 20, wherein the optical refractive index of the medium of the fluid is lower than the biological cells or biological cell-like particles.

22. The system of embodiment 20 or 21, wherein the cross-section of the microfluidic channel is sufficiently large that the biological cells or biological cell-like particles do not physically contact the walls of the microfluidic channel to generate a deforming force when passing through the microfluidic channel.

23. The system of any of embodiments 20-22, wherein the optical tweezers are highly focused laser beams with gaussian distributions.

24. The system of any of embodiments 20-23, wherein the optical tweezers' spectral windows are transparent to one or more of:

the biological cell or biological cell-like particle,

the medium of the fluid is a mixture of,

a wall of the microfluidic channel.

25. The system according to any of embodiments 20-24, wherein the optical tweezers providing means further comprises means for periodic modulation of said optical tweezers, e.g. for periodically switching said optical tweezers on and off, thereby periodically capturing and releasing said biological cells or biological cell-like particles.

26. The system of embodiment 25 wherein the component that periodically modulates the optical tweezers is a modulator.

27. The system of embodiment 25, wherein the frequency of said modulation is at a hertz level, e.g., 1-5 Hz.

28. The system of any of embodiments 20-27, wherein the optical tweezers providing means is capable of using a plurality of optical tweezers.

29. The system of embodiment 28, wherein the plurality of optical tweezers is generated by an interferometer configured with a single laser source.

30. The system of embodiment 28, wherein the plurality of optical tweezers is generated by a plurality of laser sources.

31. The system of embodiment 28, wherein the plurality of optical tweezers are arranged in an optical tweezers array, such as a one-dimensional optical tweezers array or a two-dimensional optical tweezers array.

32. The system of embodiment 31, wherein the array of optical tweezers is created using a multimode interference waveguide with a single laser source.

33. The system of embodiment 31, wherein the array of optical tweezers is generated using a spatial light modulator with a single laser source.

34. The system of any of embodiments 20-33, wherein the focal point of the optical tweezers is set such that a cell captured by the optical tweezers is away from the wall of the microfluidic channel.

35. The system of any of embodiments 20-34, wherein the image capture device comprises a microscope objective and a digital camera, e.g., a long working distance microscope objective and a CCD digital camera.

36. The system of any of embodiments 20-35, wherein the image acquisition is configured to acquire images in a third direction transverse to the first direction.

37. The system of any of embodiments 20-36, wherein said image acquisition is configured to acquire images of said biological cells or biological cell-like particles before, during and after capture by said optical tweezers.

38. The system according to any of embodiments 20-37, wherein the analysis device is configured to measure a dimension (e.g. diameter length) of biological cells or biological cell-like particles captured by the optical tweezers in the image along the first direction and perpendicular to the first direction, and to use the ratio of the two for characterizing the amount of deformation.

Brief description of the drawings

FIG. 1: (a) the working principle schematic diagram of optical fluid tweezers taking and dragging stretching operation is shown. Biological cells or biological cell-like particles are captured by the optical tweezers near the focused beam portion while being dragged and stretched by the microfluidic flow in the flow direction (x-direction) transverse to the beam (z-direction). The light beam is periodically blocked to allow the stretched cells or particles to escape from the optical tweezers; (b) a top view of optical tweezers and fluid dragging stretching is shown; (c) a schematic representation of the characterization protocol is shown, with harder (unhealthy) cells (cell 2) being less deformed (less change in dx/dy ratio) than normal (healthy) cells (cell 1) under the same test conditions.

FIGS. 2 (a) - (b) show schematic diagrams of (a) a cross-sectional view and (b) a top view of a fabricated microfluidic chip with an inlet and an outlet for continuous cell or particle delivery and collection, and (c) a schematic diagram of an experimental setup, MO: microscope objective, L ED: light emitting diode.

FIG. 3: (a) showing the extraction trace of multiple permeable expanded rabbit RBCs flowing from the field of view; (b) a scatter plot showing the cell velocity as a function of x position in the area indicated by black lines in (a), the position of the optical tweezers being indicated by a grey background bar, the total number of cells N and the number of captured cells N 'being labeled (N, N'); (c) a scattergram showing the change in cell velocity with x position in the area indicated by a white line in (a) is shown.

FIG. 4: (a) - (e) optical micrographs showing representative cells passing through the region of interest; (f) - (h) shows the trend of the variation of the x position of the cells in the region of interest with the extracted (f) dx, (g) dy and (h) dx/dy ratios indicated by the black boxes in fig. 3(a), the positions of the optical tweezers being represented on the three graphs by the gray background bars. (i) - (k) shows the extracted (i) dx, (j) dy and (k) dx/dy ratios as the trend of the change in the x position of the cell in the region indicated by the white line frame in FIG. 3 (a).

FIG. 5: (a) - (e) shows the dx/dy ratio of the extracted untreated (healthy) cells as the trend of the cells in x position at different flow rates from about 3.0. mu.l/hour to about 1.0. mu.l/hour; (f) - (j) shows the dx/dy ratio of the extracted glutaraldehyde-treated cells (unhealthy) as the trend of the cells in x position at different flow rates of about 3.0 μ l/hr to about 1.0 μ l/hr; the position of the optical tweezers is indicated by the grey background bars. The total number of cells N and the number of captured cells N 'are labeled as (N, N').

FIG. 6: a comparison of the maximum dx/dy ratio of untreated (left-sided column) and chemically treated (right-sided column) RBCs at different flow rates is shown. The total number of cells N and the number of captured cells N 'are labeled as (N, N').

FIG. 7: a process schematic is shown for an embodiment of two spatially multiplexed parallel optical tweezers and a fluid drag process based on the use of two sufficiently spaced optical tweezers in a microfluidic channel.

FIG. 8: a schematic of an experimental setup is shown for an embodiment of a spatially multiplexed parallel optical tweezers and fluid drag process based on the use of two sufficiently spaced optical tweezers in a microfluidic channel, where two separate laser sources are used to generate the two light beams.

FIG. 9: a schematic of an experimental setup is shown based on an embodiment of spatially multiplexed parallel optical tweezers and a fluid drag process using two sufficiently spaced optical tweezers in a microfluidic channel, where an interferometer configuration comprising two beam splitters and two mirrors is employed to generate two displaced beams from one laser source.

FIG. 10: a process schematic is shown based on an embodiment of a parallel optical tweezers and fluid drag process spatially multiplexed in a microfluidic channel using an array of optical tweezers in (a) one dimension and (b) two dimensions.

FIG. 11: a schematic of an experimental setup based on an embodiment of parallel optical tweezers and fluid drag process spatially multiplexed in a microfluidic channel using a one-dimensional array of optical tweezers formed by interference between two beams is shown, where the interferometer configuration comprises two beam splitters and two mirrors, generating two overlapping beams from a single laser source.

FIG. 12 is a schematic diagram of an experimental setup showing another embodiment based on spatially multiplexed parallel optical tweezers and fluid dragging process in a microfluidic channel using an array of optical tweezers created by shaping a laser beam using MMI waveguides.an inset schematically shows the 3 × 3 optical lattice generated from the output end facet of the waveguides as an example.

FIG. 13: a schematic of an experimental setup is shown for another embodiment based on an array of optical tweezers generated by shaping a spread laser beam using a spatial light modulator, spatially multiplexed parallel optical tweezers in a microfluidic channel and a fluid dragging process.

Detailed description of the preferred embodiments

As mentioned above, optical tweezers technology and microfluidics technology, respectively, have been used to evaluate mechanical properties of biological cells, but the inventors of the present application believe that they each potentially suffer from certain drawbacks. For example, most optical force-only based cell stretching techniques rely on static conditions during testing to minimize the effect of fluid stress on cell deformation, which is detrimental to high throughput characterization. The method based on the microfluidics technology usually requires high-speed movement of cells in a microfluidics channel, and can only deform the cells in an extremely short time scale of 1 mu s or 1ms, so that the measured cell elasticity cannot be compared with the static cell elasticity; in addition, the requirements on the imaging system are high, the exposure time is required to reach 1 mu s, and the frame rate reaches 4000 fps. Accordingly, the inventors of the present application have made extensive studies and research to provide a method and system for determining mechanical properties of biological cells or biological cell-like particles that combines optical tweezers technology and microfluidic technology, and the inventions of the present application are capable of overcoming at least in part the above disadvantages.

Without being bound by any particular theory, the working principle of the inventions of the present application includes: based on optical tweezers technology, optical tweezers are used to capture cells or particles in a fluid flow near a focused light beam, and the captured cells or particles are dragged and stretched by the fluid flow transverse to the light beam to generate deformation, thereby evaluating the mechanical properties of the cells or particles according to the deformation amount. In some cases, the light beam may be periodically (e.g., at a frequency of Hz level) blocked to allow for temporary capture of cells or particles and to allow stretched cells or particles to escape from the optical tweezers. The fluid flow enables continuous delivery of cells or particles into the optical tweezers while entraining cells or particles after capture evaluation. In some cases, it may be advantageous to have the cells or particles captured by the optical tweezers at a location remote from the various walls (including the base and sidewalls) of the microfluidic channel, so as to avoid physical contact between the cells and the solid surface of the microfluidic channel. Some inventions of the present application enable non-contact, continuous characterization (high throughput) of mechanical properties of a variety of biological cells or particles. Some of the inventions of the present application can be implemented in a variety of ways, including, but not limited to, using a spatially multiplexed version of multiple optical tweezers or optical tweezer arrays in a microfluidic flow to achieve higher characterization throughput. Some of the inventions of the present application are applicable to the characterization of mechanical properties of biological cells and particles, and are not limited by the original cell or particle shape or cell type.

In contrast to the previously described methods for characterizing the mechanical properties of cells based solely on optical techniques, the various inventions of the present application can present one or more advantages of integrated microfluidic techniques, such as (i) the use of fluid forces to stretch and deform the captured cells or particles, (ii) the ability of fluids to continuously transport cells or particles to optical tweezers, and (iii) the ability of fluids to carry cells or particles away. Thus, some inventions of the present application allow for continuous fluid flow during measurements, and thus dynamic test conditions can achieve greater characterization throughput than optical-only based techniques that rely on static test conditions.

Compared to the aforementioned methods for characterizing the mechanical properties of cells based solely on microfluidic technology, the inventions of the present application can embody one or more advantages of integrated optical tweezers technology, e.g., (i) optical tweezers can generate a sufficiently large optical trapping force on the pN level to perform optical tweezers trapping of cells or particles in a fluid flow, thereby enabling the fluid resistance to stretch the cells or particles; (ii) by deforming the cells over a longer time scale (e.g., 0.1s), the measured cell elasticity can be compared to the static cell elasticity; (iii) reducing the requirements on the imaging system. In some cases of the present application, operating the optical tweezers at low power (tens of mW or less) at a wavelength of 1064nm, which is permeable to biological cells, may avoid any optical or thermal damage to the cells.

Various exemplary embodiments of the present application are described in detail below with reference to the accompanying drawings. It is to be understood that the embodiments described are not intended to limit the invention of the present application in any respect, and that the technical features of the embodiments may also be combined within the scope understandable to the person skilled in the art.

Fig. 1 (a) schematically illustrates the working principle of optical fluid optical tweezers and fluid drag stretching according to one embodiment of the present application. The figures show osmotically swollen red blood cells as an example. The optical tweezers and fluid drag process involves trapping cells near the focused gaussian beam portion using optical tweezers, and the microfluidic flow drags and stretches the trapped cells in a flow direction (x-direction) transverse to the beam (z-direction). Under dynamic equilibrium, the stretched cells experience balanced optical gradients and fluid forces.

The embodiment shown in figure 1 (a) also blocks the light beam periodically (at horizontal frequency in Hz) to allow the stretched cells to escape from the optical tweezers and move with the fluid when the light beam is blocked. The fluid flow continuously delivers the cells to the optical tweezers and stretches the trapped cells. In the embodiment shown in figure 1 (a), to avoid physical contact between the test cell and the solid surface, the optical tweezers are focused above the substrate and away from the walls of the microfluidic channel, so that the cell is trapped at a location away from any solid surface, e.g., see figure 1 (a) a cross-sectional view and (b) a top view.

FIG. 1 (c) schematically illustrates a scheme for characterization of the mechanical properties of cells. In this embodiment, a conventional optical microscope is used with a bright field imaging system to record a top view image of the cells through the top surface of the microfluidic channel. The system includes a long working distance microscope objective and a digital charge-coupled device (CCD) camera. The recorded video and images are analyzed using an image processing program.

The deformation of a cell under a stretching force depends on its stiffness or elasticity. Under the same force, the stiffer cells (cell 2 in the schematic) have a smaller amount of deformation than the less stiff cells (cell 1 in the schematic). Thus, by measuring the amount of deformation of a cell under a certain stretching force, in the case where a disease state may affect the mechanical properties of the cell, it is possible to provide a certain indication for distinguishing healthy (elastic) cells from diseased (less elastic) cells.

To quantify the deformation of the cells, the dimensions (e.g., path length) of the cells along the flow direction (x-direction) and transverse to the flow direction (y-direction) can be measured from the top view image, labeled dx and dy, respectively, to extract the dx/dy ratio representing the cell shape. The dx/dy ratio of the osmotically swollen spheroid cells before stretching is 1, and the dx/dy ratio of the cells stretched in the flow direction exceeds 1. Considering that the cells are not perfectly spherical before stretching, the dx/dy ratio during and after stretching can be normalized in the analysis with respect to the ratio before stretching.

In the development phase of the present application, cell stiffness was qualitatively compared by comparing the dx/dy ratio between healthy and chemically treated (non-healthy cell modeling) cells under the same test conditions. Under the assumption of a linear spring model and fluid force dominated cell stretching, a more quantitative data analysis can be performed, which can estimate the cell elasticity (by measuring the change in dx under the estimated fluid drag force) to a first order approximation.

Experimental verification

The following describes a confirmatory experiment on the inventive concept of the present application.

Microfluidic channel fabrication

The microfluidic channels were fabricated in this experiment using standard soft lithography techniques. Briefly, contact lithography and Deep Reactive Ion Etching (DRIE) are used to define microfluidic channel patterns on silicon chips. The patterned silicon chip was used as a mold to transfer the pattern to a PDMS layer of approximately 3mm thickness. After peeling off the PDMS from the silicon mold, an inlet and an outlet having a diameter of about 1mm were formed at both ends of the microfluidic channel in the PDMS layer, respectively, using a punch.

The patterned PDMS layer was bonded to a thin cover glass to form microfluidic channels. The PDMS layer and the glass slide were treated with oxygen plasma prior to bonding them together. The bonded PDMS-glass interface was sufficiently stable at various flow pumping rates from 1. mu.l/hr to 3. mu.l/hr.

Figures (a) and (b) in fig. 2 schematically show a cross-sectional view and a top view of a microfluidic chip with an inlet and an outlet for cell delivery. In this experiment, PDMS microfluidic channels were used with a width of about 90 microns and a height of about 40 microns.

Experimental device

The diagram (c) in fig. 2 schematically shows the experimental setup. A continuous wave fiber laser with the wavelength of 1064nm is adopted as a light source.This wavelength is compatible with biological applications, with very little photodamage to biological cells (cells are permeable to 1064nm wavelength), and very little absorption of water (α ═ 0.61 cm)^-1) So that the laser beam does not heat the fluid medium. The linearly polarized laser power was controlled using a rotatable half-wave plate and a fixed Polarizing Beam Splitter (PBS).

The laser beam is focused into the microfluidic channel using a microscope objective lens having a high Numerical Aperture (NA) of about 0.85 in air. The focused beam waist diameter inside the fluid medium was estimated to be about 1.1 microns with a depth of focus of about 2.6 microns. The position of the beam waist relative to the channel substrate is calibrated by back-reflecting the beam divergence from the glass-fluid interface. For this experiment, the optical tweezers were placed at a distance of about 6 μm above the substrate to avoid the captured cells (theoretical diameter of about 7 μm) from contacting the substrate.

The laser is periodically blocked using a mechanical shutter to allow the stretched cells to escape from the optical tweezers after reaching a dynamic equilibrium state, thereby enabling continuous characterization of multiple incoming cells brought about by the fluid. The blocking frequency was set to about 2Hz (corresponding to a laser exposure time of about 0.25 s) to provide sufficient time for capturing and stretching the cells. Other embodiments based on this embodiment include adjusting and optimizing the blocking frequency to stretch the cells to a state that statistically allows the most sensitive differentiation between different cellular elasticities.

The cells were imaged in the microfluidic channel using a long-range microscope objective with an NA of 0.42 and onto a digital CCD camera, which recorded images of RBCs at a frame rate of about 80fps, with a reduced field of view of about 35 μm × to about 35 μm.

Tubing was used to connect the inlet of the microfluidic channel to the syringe pump. The flow rate in the channel is controlled by controlling the rate of the pump. Another tube is used to connect the outlet of the channel to the container to collect waste.

Cell sample preparation

A rabbit blood sample of about 5m L was obtained from animal and plant Care laboratory (APCF) of hong Kong scientific university, about 10% heparin was added to the blood sample for anticoagulation, diluted in about 0.6 Xphosphate buffered saline (PBS) buffer solutionRelease of blood to osmotically expand RBC, estimated cell concentration of about 10⁸Cells/ml. Prior to the experiment, the red blood cells were incubated in buffer solution for about 10 minutes. For chemical treatment of cells, RBCs were incubated in about 0.6x PBS buffer with a glutaraldehyde concentration of about 0.002 v/v%. Other embodiments based on this protocol include the ability to extract sensitivity to cell stretching by studying the dependence of cell deformation and elasticity on chemical treatment concentration.

Image processing

The editor program processes all frame images in the video recorded by the CCD camera using the image processing toolkit and the computer vision system toolkit in Matlab. Image processing essentially comprises three steps including (i) edge detection, (ii) object segmentation and (iii) feature extraction.

In edge detection, a gradient image is computed from a recorded top view image of the cell and adaptive thresholds are applied to create a binary mask containing the cell edge contours. In object segmentation, cell edge contours are obtained from the binary mask and irrelevant edges are removed. In feature extraction, the centroid position, velocity and size of the cell are measured from the extracted cell profile.

Using the extracted position of the cell in each frame, the cell trajectory can be tracked and the velocity of the cell as it passes through a field of view of about 35 μm × about 35 μm.

Cell capture by optical tweezers in microfluid

Figure 3(a) is a graph showing multiple permeability expanded rabbit RBC extraction traces from about 78s video flowing through a field of view at a flow rate of 2.5 μ l/hr. The x-y position of the optical tweezers is marked with a white circle. The optical tweezers were located at about 17 μm for x and about 18 μm for y, with an estimated optical power of about 45 mW. Considering a beam waist diameter of about 1.1 μm, the estimated optical intensity is about 10⁶W/cm². Other embodiments based on this scheme include the use of lower optical power and reduced flow rates.

A region of interest is defined, with dimensions of about 6 μm (y) × about 25 μm (x), represented by the black box in the figure, where the incoming cells overlap the optical tweezers as they pass through the field of view.A total of N190 cells pass through the region of interest.N'. apprxeq.90 cells are attracted and temporarily trapped on the optical tweezers and then escape along the line in the direction of flow.

For comparison, it was also observed that cells outside the region of interest (without spatial overlap with the optical tweezers) always moved along the flow, unaffected by the optical tweezers.

The velocity of the cells across the field of view is extracted from the video. Fig. 3 (b) shows a scatter plot of cell velocity as a change in x position within the region of interest. The curves show the average of the cell velocities. The gray background bar marks the x position of the optical tweezers. The total number of cells in the region of interest is labeled N and the number of captured cells N 'is labeled (N, N').

The average entry velocity of the cells was observed to be about 120 μm/s. The optical tweezers attract and capture cells at a location where x is about 19 μm, where the average cell velocity drops to almost zero.

When x > about 19 μm, the stretched cells are released and accelerated by the flow, and the cell velocity rises from almost zero to about 150 μm/s. It should be noted that the average accelerated velocity of the released cells is higher than the entry velocity of about 120 μm/s. This phenomenon may be due to the fact that: the cell is slightly floated and trapped by the optical tweezers at a plane above its entry plane (closer to the substrate due to gravity). When releasing cells, they have a higher flow rate in laminar distribution (the flow rate is greatest near the center of the channel) than near the substrate.

The plot in fig. 3 (c) shows a scatter plot of cell velocity as a function of cell x position in the region outside the region of interest (as indicated by the white line box in the plot in fig. 3 (a)). A uniform distribution of cell velocities was observed with an average of about 120 μm/s, indicating that the cell velocities outside the region of interest were not affected by the optical tweezers.

Stretching captured cells using microfluidic flow

Fig. 4 (a) - (e) show optical micrographs of representative cells across a region of interest. As shown in fig. 4 (a), the cells exhibited an almost symmetrical shape due to osmotic swelling before the capture. As shown in fig. 4 (b) and (c), during the stretching process, the cell elongates in the flow direction (x direction) due to the fluid drag force. As shown in fig. 4 (d) - (e), after releasing the cells from the optical tweezers, the cells gradually returned to an almost symmetrical shape. In fig. 4 (a) the length of the cells in the flow direction (x-direction) and transverse to the flow direction (y-direction), i.e. dx and dy, is marked.

The dx, dy and dx/dy ratios of the cells in the region of interest are extracted as indicated by the black line boxes in the graph of (a) in fig. 3. FIGS. 4 (f) - (h) show scatter plots of extracted dx, dy and dx/dy ratios as changes in cell x position. In each case, an average value was obtained, as shown by the curve in each figure. The position of the optical tweezers is indicated on the three figures using a grey background bar.

Before the incoming cells reach the x ═ about 17 μm optical tweezers, the average dx and dy values for the cells are both about 6.7 μm (e.g., for x <13 μm), as shown in the (f) and (g) plots of fig. 4. Accordingly, the average dx/dy ratio of the input cells was about 1.0 (for x <13 μm), as shown in (h) of fig. 4. These results indicate that the cell shape prior to capture is almost spherical (osmotic swelling) before the cell interacts with the optical tweezers.

Cells captured at the x-19 μm position are stretched and deformed by microfluidic flow. The average dx increases and the average dy decreases, as shown in (f) and (g) of fig. 4, which respectively show that the cells are elongated in the x-direction and compressed in the y-direction. At x ═ about 20 μm, dx reaches a maximum of about 7.5 μm (Δ dx is about 0.8 μm) and dy reaches a minimum of about 6.1 μm (Δ dy is about-0.6 μm). Accordingly, the dx/dy ratio reaches a maximum of about 1.2 at maximum deformation.

After the cells were released from the optical tweezers, the mean dx returned to about 6.8 μm and the mean dy returned to about 6.6 μm within about 2 μm of displacement. Accordingly, the dx/dy ratio returns to about 1.0, indicating that the cell shape returns to approximately spherical after the optical tweezers and fluid drag process.

The fluid force F exerted on the captured cells is estimated to be at the pN level by the Stokes' law using a first order approximation that F6 π η Rv, where η (10 π η Rv)^-3Nm^-2S) is the dynamic viscosity of the PBS solution, R (about 3.4 μm) is the cell radius, and v (about 150 μm/s) is the relative velocity between the cell and the fluid. Assuming a linear spring model, the spring constant of RBC is estimated to be of the order of μ N/m, considering that the change in cell dx is about 1 μm.

Throughout the video of about 78s, about 90 cells were captured and dragged, with a corresponding average cell spreading flux of about 1.2 cells/sec.

For comparison, the dx, dy and dx/dy ratios of the cells outside the region of interest (as indicated by the white line boxes in the (a) diagram in fig. 3) are extracted. FIGS. 4 (i) - (k) show the dx, dy and dx/dy ratios (N about 170) of the extracted cells as changes in the x-position of the cells. The plots (i) - (j) in FIG. 4 show that dx and dy are consistent along the x position, with an average of about 6.7 μm. The plot in fig. 4 (k) shows that the dx/dy ratio is consistent along the x position, with an average of about 1.0, indicating that there is no significant deformation of cells that do not interact with the optical tweezers.

Distinguishing between healthy (untreated) and unhealthy (chemically treated) cells

To further demonstrate the applicability of optical tweezers and fluid-dragged cell stretching, osmotically swollen red blood cells were measured with and without chemical treatments that changed cell stiffness. For chemical treatment of cells, erythrocytes were treated with glutaraldehyde at a concentration of about 0.002% v/v. Glutaraldehyde is known to cross-link cellular proteins, thereby increasing cell rigidity.

Erythrocytes treated with and without glutaraldehyde were characterized at different flow rates of about 3.0. mu.l/hr to about 1.0. mu.l/hr. For each test condition, the dx/dy ratio of the cells in the region of interest is extracted. Fig. 5 (a) - (j) illustrate a comparison of dx/dy ratios of untreated and chemically treated RBCs at different flow rates. The mean value of the dx/dy ratio was calculated from the plurality of cells as shown by the black curve. Standard error is indicated using a grey shading around the black curve. For each plot, the total number of cells N and the number of captured cells N 'are labeled as (N, N').

The average dx/dy ratio into the cells was about 1.0 under all conditions tested, indicating that the flow rate or chemical treatment had no significant effect on the cell shape before reaching the optical tweezers. It was found that under most of the conditions tested (except for the (j) graph in FIG. 5), the average dx/dy ratio increased when the cells were stretched. Under all conditions tested, the average dx/dy ratio returned to about 1.0 after the cells were released from the optical tweezers, indicating that the cell shape returned to a nearly spherical shape.

Fig. 6 shows the maximum dx/dy ratio of extracted untreated and chemically treated RBCs. For untreated and chemically treated cells, the maximum dx/dy ratio increases with flow rate. This phenomenon can be attributed to the fact that: the trapped cells experience large stretching forces when flowing at high velocities. At all flow rates, the chemically treated cells showed less deformation than the untreated cells, consistent with cell hardening due to the chemical treatment. Other embodiments based on embodiments include testing the cell deformation for dependence on chemical treatment concentration, thereby establishing the sensitivity of the stretching procedure to chemical treatment.

Other embodiments of the present application are described below.

Fig. 7 schematically illustrates another embodiment of the present application using two optical tweezers for two parallel optical tweezers and a fluid drag process. The two optical tweezers are arranged at a sufficiently large separation in a direction transverse to the flow direction so that they can perform optical tweezers action on two cells or particles independently and simultaneously, enabling two independent optical tweezers and fluid drag processes in the same microfluidic flow. This embodiment can achieve a 2-fold characterization throughput compared to the above embodiment.

Fig. 8 schematically illustrates a method of producing two optical tweezers in a microfluidic channel, wherein two separate laser sources are used, which do not necessarily have the same output power or wavelength. After simultaneously focusing the two laser beams using the same high NA microscope objective, two optical tweezers can be generated within the microfluidic channel.

Fig. 9 schematically illustrates another method of producing two optical tweezers in a microfluidic channel, wherein a single laser source is employed. An input laser beam is split into two parallel propagating displaced laser beams using an interferometer setup comprising two beam splitters and two mirrors. Two optical tweezers can also be generated within the microfluidic channel after simultaneously focusing two laser beams using the same high NA microscope objective.

FIGS. 10 (a) and (b) schematically illustrate another embodiment of the present application that uses one-dimensional and two-dimensional arrays of optical tweezers for multiple parallel optical tweezers and fluid drag processes the array of optical tweezers comprises a plurality of optical tweezers that may be arranged as a one-dimensional 1 × N grating, a two-dimensional N × M rectangular grating, a two-dimensional hexagonal grating, etc. when multiple cells or particles enter the array region, the arrayed optical tweezers in a microfluidic may simultaneously and independently perform the optical tweezers and fluid drag processes, provided that the spacing between the optical tweezers should be sufficiently large.

Fig. 11 schematically illustrates a method of producing a one-dimensional array of optical tweezers in a microfluidic channel, wherein a single laser source is employed. The input laser beam is split into two spatially overlapping, parallel propagating laser beams using an interferometer setup comprising two beam splitters and two mirrors. After simultaneously focusing the two laser beams using the same high NA microscope objective and interfering inside the microfluidic channel, a one-dimensional array of optical tweezers can be generated inside the microfluidic channel. By controlling the phase difference and spatial overlap between the two beams, the optical interference pattern can be fine tuned.

FIG. 12 schematically illustrates a method of producing one-and two-dimensional arrays of optical tweezers in a microfluidic channel using a single laser source, shaping an input Gaussian beam into an output multi-beam array based on multi-mode interference inside the waveguide using passive multi-mode interference (MMI) waveguides, generating multiple optical lattices using MMI waveguides by using different cross-sectional shapes and waveguide sizes and lengths, the inset schematically illustrates the 3 × 3 optical lattice generated from the waveguide output facet as an example, input and output coupling using a microscope objective with MMI waveguides.

FIG. 13 schematically illustrates another method of producing one-and two-dimensional arrays of optical tweezers in a microfluidic channel, wherein a single laser source is employed, using a computer-controlled spatial light modulator (S L M) to shape the expanded single beam into an optical tweezers lattice (referred to as holographic optical tweezers).

Finally, it should be understood that while the various aspects of the present specification describe specific embodiments, those skilled in the art will readily appreciate that the disclosed embodiments are merely illustrative of the principles of the subject matter disclosed herein. Accordingly, it is to be understood that the disclosed subject matter is not limited to the specific combinations, methods, and/or formulations, etc., described herein, unless otherwise specified. Moreover, those of ordinary skill in the art will recognize that certain changes, modifications, permutations, variations, additions, subtractions and sub-combinations may be made in accordance with the teachings herein without departing from the spirit of the present specification. It is therefore intended that the following appended claims be interpreted as including all such alterations, modifications, permutations, variations, additions, subtractions and sub-combinations as fall within the true spirit and scope thereof.

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Claims (2)

1. A method for determining a mechanical property of a biological cell or biological cell-like particle, comprising the steps of:

passing a fluid comprising the biological cells or biological cell-like particles through a microfluidic channel in a first direction;

capturing at least a portion of the biological cells or biological cell-like particles with optical tweezers in a second direction transverse to the first direction during passage of the fluid through the microfluidic channel;

collecting an image of the biological cells or biological cell-like particles through the microfluidic channel; and

determining a mechanical property of the biological cell or biological cell-like particle based on the amount of deformation of the biological cell or biological cell-like particle captured by the optical tweezers in the image.

2. A system for determining a mechanical property of a biological cell or biological cell-like particle, comprising:

a microfluidic device comprising a microfluidic channel for passing a fluid comprising the biological cells or biological cell-like particles in a first direction;

optical tweezers providing means for capturing at least a part of said biological cells or biological cell-like particles with optical tweezers in a second direction transverse to said first direction during passage of said fluid through said microfluidic channel;

an image acquisition device for acquiring images of the biological cells or biological cell-like particles passing through the microfluidic channel; and

and the analysis device is used for measuring the mechanical property of the biological cells or the biological cell-like particles based on the deformation quantity of the biological cells or the biological cell-like particles captured by the optical tweezers in the image.

Images