Microfluidic measurement system based on single pinhole spatial fluorescence cross-correlation spectroscopy

Abstract

This invention relates to the field of single-molecule fluorescence microfluidics measurement technology, and discloses a microfluidics measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy. The system includes a laser excitation module, a confocal filtering module, a dual-channel spectroscopic detection module, and a signal processing module. The laser source is reflected by a dichroic mirror to a water immersion objective, forming a fly-level focusing excitation volume in the sample. The fluorescence signal is spatially filtered through a single pinhole and then split into two paths by a non-polarizing beam splitter, which are then converted into electrical pulse signals by an avalanche diode (APD) detector. A hardware correlator calculates the cross-correlation function of the two signals in real time, combining the trigonometric relationship between flow velocity and direction to achieve high-precision calculation. This invention simplifies the three-dimensional optical alignment process required by traditional spatial dual-laser beam FCCS, generating two equivalent detection half-volumes within the same focusing excitation volume through a combination of a non-polarizing beam splitter and a lens, enabling non-invasive and precise measurement of the microfluidic dynamics of living biological samples.

Description

Technical Field

[0001] This invention relates to the field of microfluidic measurement technology, and in particular to a microfluidic measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy. Background Technology

[0002] Microfluidic dynamics has been extensively studied over the past few decades due to its unique value in chemical analysis and bioscience research. Microfluidic systems typically possess characteristics such as small volume, high surface-to-volume ratio, and rapid thermal response, enabling precise manipulation of fluids, including processes such as mixing, reaction, and separation. This makes them valuable for applications in biomolecular analysis, drug screening, cell culture, tissue engineering, and developmental biology. To fully explore and optimize the structural properties of microfluidics, effective characterization tools are needed to measure and analyze the microfluidic dynamic behavior within these systems. Currently, particle imaging velocimetry and fluorescence cross-correlation spectroscopy are widely used in microfluidic velocimetry, with fluorescence cross-correlation spectroscopy demonstrating significant advantages in microscale fluid measurements due to its high sensitivity and non-invasive nature.

[0003] However, while traditional fluorescence cross-correlation spectroscopy (FCCS) can measure flow velocity, it lacks the ability to directly acquire microfluidic flow direction information, severely limiting its further application in complex fluid dynamics research. Although traditional spatial two-beam fluorescence cross-correlation spectroscopy can acquire flow direction information, it requires precise separation of two laser beams in three-dimensional space and the generation of two precisely aligned focusing excitation volumes, leading to a complex system optical structure and significantly increased operational difficulty. Furthermore, improvements to reduce operational difficulty include expanding the focusing excitation volume and employing a dual-pinhole mode, achieving the simultaneous inclusion of two detection volumes within the same focusing excitation volume. However, this inevitably sacrifices the original high spatial resolution advantage of fluorescence cross-correlation spectroscopy, limiting its application in high-precision measurement scenarios. Simultaneously, existing technologies still face alignment and calibration difficulties in characterizing the flow direction of microfluidic channels and living biological samples such as zebrafish embryonic blood vessels, and the measurement device may cause additional interference to the original flow characteristics, further restricting its applicability in live biological sample measurements.

[0004] Therefore, there is an urgent need to develop a microfluidic velocity and flow direction measurement technology that can simultaneously meet the requirements of easy operation, high spatial resolution, direct acquisition of flow direction information, and good in vivo applicability, so as to promote the in-depth development of microfluidic dynamics research in biomedical engineering and other related fields. Summary of the Invention

[0005] This invention addresses the shortcomings of existing microfluidic measurement technologies, such as complex system structure, high operational difficulty, limited spatial resolution, difficulty in alignment and calibration, and insufficient applicability to live biological samples. It proposes a microfluidic velocity and direction measurement method and system based on single-pinhole spatial fluorescence cross-correlation spectroscopy (FCCS).

[0006] The microfluidic velocity and direction measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy proposed in this invention includes a laser excitation module, a confocal filtering module, a dual-channel spectroscopic detection module, and a signal processing and analysis module. Specifically: the laser excitation module consists of a laser source, a reflector, a dichroic mirror, and a water immersion objective connected in sequence, used to generate a micrometer-scale focused excitation volume; the confocal filtering module consists of a condenser lens, a reflector, a relay lens, and a single pinhole connected in sequence, used to spatially filter the fluorescence signal to eliminate stray light outside the focal plane; the dual-channel spectroscopic detection module consists of a non-polarizing beam splitter, a first relay lens, a first emission filter, a first APD detector, a second relay lens, a second emission filter, and a second APD detector, used to split the filtered fluorescence signal into two paths and perform photoelectric conversion on each; the signal processing and analysis module consists of a hardware correlator and a computer, used to perform real-time cross-correlation calculations on the two photoelectric signals and calculate the velocity and direction parameters.

[0007] Furthermore, the present invention employs an optical architecture of a single laser source and a single pinhole confocal filter. This architecture eliminates the need for precise alignment of two laser beams in three-dimensional space. High-precision microfluidic measurement can be achieved solely through the combination of a single laser source and a single pinhole, significantly simplifying the system structure and alignment calibration process, and effectively avoiding measurement errors caused by optical path drift.

[0008] Furthermore, this invention utilizes the optical combination of a non-polarizing beam splitter and a lens to "softly divide" two equivalent probe half-volumes (dashed circles) within the same focusing excitation volume (solid circle). The non-polarizing beam splitter divides the fluorescence signal passing through the single pinhole into two independent optical paths with a 50:50 energy ratio. The first relay lens and the second relay lens are located at the two output ends of the non-polarizing beam splitter, respectively, and work together with the non-polarizing beam splitter to optically form two axially closely adjacent equivalent probe half-volumes. The axial directions of these two probe half-volumes form a fixed geometric relationship, and there is a specific geometric ratio between their focal distance and the waist radius of a single focusing volume.

[0009] Furthermore, the laser beam emitted by the laser source propagates horizontally, and after the propagation direction is changed by the reflector, it is incident on the dichroic mirror. The angle between the dichroic mirror and the incident laser is 45 degrees. The dichroic mirror reflects the shorter wavelength of the incident laser while transmitting the longer wavelength of the emitted fluorescence, thus achieving spatial separation of the excitation optical path and the fluorescence collection optical path. The laser beam reflected by the dichroic mirror enters the water immersion objective vertically upward. The water immersion objective focuses the laser beam into the interior of the sample under test with a high numerical aperture, forming a tiny focusing excitation volume on the submicron scale, which excites the fluorescent molecules in the sample to generate a fluorescence signal.

[0010] Furthermore, the fluorescence signal generated by the stimulated sample returns along the original optical path to the immersion objective lens, is transmitted through the dichroic mirror, enters the condenser lens for collimation and focusing, then passes through the reflector to change its propagation direction before entering the relay lens for relay imaging, and finally passes strictly through the central aperture of the single pinhole. The single pinhole, as the core element of spatial filtering, plays a role in strictly limiting the confocal observation volume, eliminating stray fluorescence and background noise from outside the focal plane, and ensuring that only the fluorescence signal from the focal point is received by the subsequent detection system.

[0011] Furthermore, the fluorescence signal after single-pinhole spatial filtering enters the non-polarizing beam splitter and is evenly divided into two paths. The first path of fluorescence is adjusted in spatial position by the first relay lens, and then a specific wavelength of fluorescence signal is selected by the first emission filter. The purified fluorescence enters the first APD detector. The second path of fluorescence is adjusted in spatial position by the second relay lens, and then undergoes the same wavelength selection process by the second emission filter. The purified fluorescence enters the second APD detector.

[0012] Furthermore, the first and second APD detectors convert the collected fluorescent photons into TTL electrical pulse signals. The two pulse signals are transmitted synchronously to the hardware correlator, which calculates the cross-correlation function of the two fluorescent signals in real time and transmits the calculation result to the computer. The dedicated computer software is responsible for system control, data acquisition, cross-correlation curve fitting, flow velocity calculation, flow direction angle calculation, and result display.

[0013] Furthermore, the cross-correlation function of the fluorescence signals from the two detection channels The normalized form can be represented by a specific mathematical expression: exp Where N represents the average number of molecules within the focused excitation volume, and τ represents the time offset. d Let τ represent the diffusion time, K represent the structure coefficient, R represent the ratio between the focal distance between the two probe volumes and the waist radius of a single focusing volume, and τ represent the diffusion time. xfThe value represents the flow time of fluorescent molecules from one detector volume to another in the XY two-dimensional plane space, and α represents the angle between the axial direction of the two detector volumes and the microfluidic flow direction.

[0014] Furthermore, since the axial distance between the two detector half-volumes of the present invention is very close, it is inevitable that fluorescent molecules will be detected repeatedly by the two detectors. In order to obtain a more accurate and reliable cross-correlation function, it is necessary to simultaneously obtain the positive cross-correlation function and the negative cross-correlation function and perform calibration by subtraction, so as to eliminate the molecular repetition counting error caused by the close proximity of the detector volumes and improve the accuracy of flow rate and flow direction calculation.

[0015] Furthermore, during flow velocity measurement, the sample to be tested is rotated so that the axial directions of the two probe volumes gradually coincide with the flow direction of the microfluidic fluid. When the amplitude of the cross-correlation curve reaches its maximum value, it indicates that the axial directions of the two probe volumes completely coincide with the flow direction of the microfluidic fluid. At this time, the flow velocity of the microfluidic fluid can be calculated based on the flow time τF obtained by fitting the corresponding cross-correlation curve and the known geometric distance between the two probe volumes.

[0016] Furthermore, during flow direction measurement, given the known microfluidic velocity, the sample to be tested is rotated so that the microfluidic flow direction forms different angles α with the axial directions of the two detection volumes. Cross-correlation curves are obtained at different angles. By analyzing the variation law of the amplitude of the cross-correlation curves at different angles, and combining the known flow velocity information and the geometric relationship of the cross-correlation function, the flow direction angle α is calculated inversely according to the trigonometric function relationship.

[0017] Furthermore, the entire optical system adopts an L-shaped compact confocal layout, and achieves spatial orthogonal separation of the excitation light path and the detection light path through the spatial configuration of the reflector and dichroic mirror, effectively suppressing excitation light crosstalk and improving the signal-to-noise ratio, which is suitable for measuring blood vessel flow in live biological samples such as zebrafish embryos.

[0018] The beneficial effects of the embodiments of the present invention are as follows: By adopting a simplified optical setup with a single laser source and a single pinhole, the three-dimensional spatial optical alignment process required in traditional spatial dual-beam laser fluorescence cross-correlation spectroscopy is eliminated, greatly reducing the operational difficulty and alignment calibration complexity of the system. At the same time, it avoids measurement errors caused by optical path drift, thereby improving the reliability and accuracy of the measurement. By using an optical combination of a non-polarizing beam splitter and a lens, two effective detection half-volumes can be generated within the same focusing excitation volume, eliminating the need for additional scanning mirror groups and additional pinholes. This reduces hardware costs while maintaining the original high spatial resolution advantage of fluorescence cross-correlation spectroscopy, enabling precise measurement of hydrodynamic behavior in microscale structures. By directly measuring the flow velocity and direction of microfluidics, this method solves the problems of insufficient spatial resolution of traditional particle imaging velocimetry at the microscale and the inability of traditional fluorescence correlation spectroscopy to directly obtain flow direction information. It provides a non-invasive and highly sensitive measurement method that is particularly efficient and convenient for measuring blood vessel flow in living tissue samples such as zebrafish embryos. By optimizing the signal processing flow, a hardware correlator is used to perform real-time cross-correlation processing on the pulse signals output by the dual-channel APD detector. Combined with dedicated computer software for data acquisition, curve fitting, and result analysis, the efficiency of data processing and the accuracy of measurement results are further improved. This technical solution has broad prospects for on-chip device applications that rapidly determine whether there is fluid information in a microfluidic channel. Attached Figure Description

[0019] Figure 1 This is a schematic diagram of the optical path of a microfluidic velocity and direction measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy provided in an embodiment of the present invention.

[0020] The attached figures are labeled as follows: 1. Laser source; 201. First reflecting mirror; 202. Dichroic mirror; 3. Water immersion objective lens; 4. Sample to be tested; 203. Condensing lens; 204. Second reflecting mirror; 205. Relay lens; 206. Single pinhole; 207. Non-polarizing beam splitter; 208. First relay lens; 209. First emission filter; 5. First APD detector; 210. Second relay lens; 211. Second emission filter; 6. Second APD detector. Detailed Implementation

[0021] The embodiments described below with reference to the accompanying drawings are exemplary and intended to explain the invention, and should not be construed as limiting the invention.

[0022] like Figure 1 As shown in the figure, the microfluidic velocity and direction measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy provided in this embodiment of the invention includes a laser source 1, a first reflecting mirror 201, a dichroic mirror 202, a water immersion objective lens 3, a sample to be tested 4, a condenser lens 203, a second reflecting mirror 204, a relay lens 205, a single pinhole 206, a non-polarizing beam splitter 207, a first relay lens 208, a first emission filter 209, a first APD detector 5, a second relay lens 210, a second emission filter 211, and a second APD detector 6. The measurement system adopts an L-shaped compact confocal layout design, with each optical element arranged sequentially along a specific optical path to form a complete laser excitation optical path, a fluorescence collection optical path, a dual-channel beam splitting detection optical path, and a signal processing link. The structural connection relationship, working principle, and operation process of each component will be described in detail below with reference to the accompanying drawings.

[0023] The laser beam emitted by the laser source 1 propagates horizontally to the first reflecting mirror 201, which is positioned at a 45-degree angle in the optical path, changing the propagation direction of the laser beam by 90 degrees so that it is perpendicularly incident on the dichroic mirror 202. The dichroic mirror 202 forms a 45-degree angle with the incident laser beam. The surface of the dichroic mirror 202 is coated with a dielectric film of a specific wavelength band. This dielectric film can reflect the shorter wavelength incident laser beam while transmitting the longer wavelength emitted fluorescence, thereby achieving spatial separation and merging of the excitation optical path and the fluorescence collection optical path. The laser beam reflected by the dichroic mirror 202 enters the immersion objective 3 vertically downwards. The immersion objective 3 is mounted on a precision displacement platform and has high numerical aperture characteristics, preferably with a numerical aperture value greater than or equal to 1.0. In one specific embodiment, the immersion objective 3 is an Olympus 60x immersion objective with a numerical aperture of 1.2, which can precisely focus the incident laser beam onto a tiny submicron-sized volume within the sample, forming a focused excitation volume on the micron or even submicron scale. The geometry of this focused excitation volume is determined by the wavelength of the laser beam, the numerical aperture of the immersion objective 3, and the diameter of the pinhole, typically achieving a spatial resolution on the order of approximately 1 micrometer in the lateral direction.

[0024] The sample 4 to be tested is fixed on the sample stage, which can be adjusted in three-dimensional space and rotated relative to the immersion objective 3. When the laser beam is focused into the sample 4 through the immersion objective 3, the fluorescent molecules in the focused excitation volume absorb the laser energy and transition from the ground state to the excited state, subsequently generating a fluorescence signal through de-excitation. The fluorescence signal has a relatively long wavelength and returns to the immersion objective 3 along the original optical path, then passes through the dichroic mirror 202 and enters the condenser lens 203. The condenser lens 203 is located in the transmission optical path of the dichroic mirror 202, and its function is to collimate and converge the fluorescence signal, so that the fluorescence signal from the focused excitation volume can be efficiently collected and guided to the subsequent optical path.

[0025] The fluorescence signal, focused by the condenser lens 203, propagates to the second reflector 204, which is also positioned at a 45-degree angle in the optical path, changing the propagation direction of the fluorescence signal by another 90 degrees, allowing the fluorescence signal to enter the relay lens 205 for relay imaging. The relay lens 205 images the fluorescence signal from the condenser lens 203 onto the plane of the pinhole 206, ensuring that the fluorescence signal passes strictly through the central aperture of the pinhole 206. The pinhole 206, as the core component of the confocal filter, directly affects the spatial resolution and signal-to-noise ratio of the measurement system. In a specific embodiment, the aperture of the pinhole 206 is set to 150 micrometers. This aperture size can achieve a lateral spatial resolution on the order of approximately 1 micrometer under confocal imaging conditions, while ensuring sufficient fluorescence signal throughput. The spatial filtering principle of the single pinhole 206 is as follows: only fluorescence signals from the focal plane (i.e., the plane where the focusing excitation volume is located) can converge into a diffraction-limited focal point at the plane position of the single pinhole 206 and pass smoothly through the pinhole. Stray fluorescence and background noise from outside the focal plane (i.e., non-focal plane) cannot form a clear focal point at the plane of the single pinhole 206 and are therefore blocked and excluded by the single pinhole 206. Through the spatial filtering effect of the single pinhole 206, the measurement system only collects fluorescence signals from the focusing excitation volume, effectively improving the accuracy of imaging and the signal-to-noise ratio.

[0026] The fluorescence signal, spatially filtered by the single pinhole 206, enters the unpolarized beamsplitter 207. The unpolarized beamsplitter 207 is a 50:50 energy splitter, characterized by its insensitivity to the polarization state of the incident fluorescence, capable of splitting the incident fluorescence into two independent optical paths with approximately equal energy ratios. The fluorescence signal from the first optical path branch propagates along the first output end of the unpolarized beamsplitter 207 to the first relay lens 208. The first relay lens 208 further optically adjusts the spatial position of the fluorescence signal, ensuring it enters the first emission filter 209 with an optimized spot size. The passband of the first emission filter 209 precisely matches the emission spectrum of the fluorescent molecules, used to select fluorescence signals of specific wavelengths, while effectively eliminating interference from Rayleigh scattering, Raman scattering, and other background noise from the laser. The purified fluorescence signal enters the first APD detector 5, which adopts avalanche photodiode technology. Its detection surface diameter is preferably 175 micrometers, and its response time is in the order of hundreds of picoseconds. It has the characteristics of low dark count rate and high photon detection efficiency, and can convert each collected fluorescence photon into a TTL electrical pulse signal.

[0027] The fluorescence signal from the second optical path branch propagates along the second output end of the non-polarizing beam splitter 207 to the second relay lens 210. The second relay lens 210 performs the same optical adjustment spatial positioning processing on the fluorescence signal as the first relay lens 208, so that the fluorescence signal enters the second emission filter 211 with appropriate parameters. The second emission filter 211 has the same spectral characteristics as the first emission filter 209. After spectral purification processing of the fluorescence signal, the purified fluorescence signal enters the second APD detector 6. The second APD detector 6 adopts the same technical specifications and performance parameters as the first APD detector 5, ensuring that the two detection channels maintain a high degree of consistency and symmetry in terms of photoelectric conversion efficiency and time response characteristics, providing two equivalent and balanced signal channels for subsequent cross-correlation analysis.

[0028] The first relay lens 208 and the second relay lens 210 are located at the two output ends of the unpolarized beamsplitter 207, respectively. Together with the unpolarized beamsplitter 207, they optically form two axially adjacent equivalent probe half-volumes (dashed circles) within the same focusing excitation volume (solid circle). The axial directions of these two equivalent probe half-volumes form a fixed geometric relationship, with a specific geometric ratio between their focal distance and the waist radius of a single focusing excitation volume. This "soft segmentation" design is one of the core innovations of this measurement system. It eliminates the need for additional pinholes and scanning mirrors, achieving the separation and acquisition of two probe half-volumes within the same focusing excitation volume solely through the optical combination of the unpolarized beamsplitter 207, the first relay lens 208, and the second relay lens 210. This significantly simplifies the system structure while maintaining high spatial resolution.

[0029] After the first APD detector 5 and the second APD detector 6 convert the collected fluorescent photons into TTL electrical pulse signals, the two pulse signals are synchronously transmitted to the hardware correlator via a coaxial cable or dedicated signal line. The hardware correlator uses high-speed digital signal processing technology to calculate the cross-correlation function of the two fluorescence signals in real time. The processing principle of the hardware correlator is as follows: using the signal from one detection channel as the reference channel and the signal from the other detection channel as the delay channel, the signal from the delay channel is multiplied and averaged relative to the reference channel at different time delays to obtain the cross-correlation value of the two fluorescence signals at different time offsets. The hardware correlator transmits these cross-correlation calculation results to a computer in real time through a data transmission interface. The computer is equipped with dedicated measurement and analysis software, which is responsible for system control, data acquisition, real-time display of cross-correlation curves, curve fitting, flow velocity calculation, flow direction angle calculation, and graphical display of measurement results.

[0030] During the measurement process, the fluorescent molecules in sample 4 undergo two forms of motion simultaneously within the focusing excitation volume. One is the Brownian diffusion motion of the fluorescent molecules themselves, that is, the fluorescent molecules undergo random diffusion within the focusing excitation volume due to thermal motion. The diffusion rate is related to the viscosity of the solution, temperature, and the size of the fluorescent molecules themselves. The time constant corresponding to the diffusion motion is called the diffusion time τ. d Another type is the convective motion of fluorescent molecules carried by microfluidic flow. When microfluidic flow exists in the sample 4, fluorescent molecules move directionally along the flow direction driven by the fluid. The flow rate is related to factors such as fluid viscosity, channel geometry, and driving pressure. The time constant corresponding to this flow motion is called the flow time τ. xf When fluorescent molecules move within the focusing excitation volume, their emitted fluorescence signals are detected simultaneously by two APD detectors. Due to the axial offset between the two detection half-volumes in space, a time delay related to the flow direction and velocity is generated between the output signals of the two detection channels when the fluorescent molecules pass through the two detection half-volumes in sequence.

[0031] The cross-correlation function of the fluorescence signals from two detection channels can be expressed as the normalized cross-correlation result of the fluorescence intensity signals at a time offset τ. When the two fluorescence signals originate from the same color excitation fluorescence within two closely spaced focusing excitation volumes, the specific mathematical form of their cross-correlation function includes the following parameters: N represents the average number of molecules within the focusing excitation volume, τ represents the time offset between the two fluorescence signals, and τ... d Let τ represent the diffusion time, K represent the structure coefficient, R represent the ratio between the focal distance between the two probe volumes and the waist radius of a single focusing volume, and τ represent the diffusion time. xf The value represents the flow time of fluorescent molecules from one detector volume to another in the XY two-dimensional plane space, and α represents the angle between the axial direction of the two detector volumes and the microfluidic flow direction.

[0032] Because the axial distance between the two detector half-volumes in this embodiment is very close, it is inevitable that fluorescent molecules will be detected simultaneously or repeatedly by both detectors. To obtain a more accurate and reliable cross-correlation function and eliminate the molecular repetition counting error caused by the close proximity of the detector volumes, the system adopts the following calibration method: calibration is performed by simultaneously acquiring the positive cross-correlation function and the negative cross-correlation function, and then subtracting these two cross-correlation functions. The positive cross-correlation function is the cross-correlation function calculated using the signal from the first APD detector 5 as the reference channel and the signal from the second APD detector 6 as the delay channel; the negative cross-correlation function is the opposite. Differential processing can effectively eliminate the cross-correlation amplitude shift caused by repeated detection of fluorescent molecules, thereby obtaining a more accurate estimate of the flow time.

[0033] In microfluidic velocity measurement, the sample 4 to be tested is mounted on a rotatable precision sample stage, the rotation angle of which can be precisely controlled by computer software. The operator first adjusts the sample stage to an initial angle, ensuring that the flow direction of the microfluidic fluid in the sample 4 forms a specific relative angle with the axial directions of the two probe half-volumes. The measurement system is then activated, and the hardware correlator calculates and outputs the cross-correlation curve in real time. The amplitude and shape of the cross-correlation curve can be observed using the computer software as a function of the sample stage angle. Subsequently, the sample stage is rotated gradually in small angle steps (e.g., 1 to 5 degrees), and the amplitude of the corresponding cross-correlation curve is recorded after each rotation. When the amplitude of the cross-correlation curve reaches its maximum value, it indicates that the axial directions of the two probe half-volumes are completely aligned with the flow direction of the microfluidic fluid, meaning the flow direction of the microfluidic fluid is consistent with the arrangement direction of the two probe half-volumes. At this optimal angle, the flow velocity of the microfluidic fluid can be calculated using the formula: velocity V equals geometric distance D divided by flow time τF, obtained from the fitted cross-correlation curve.

[0034] When measuring the flow direction of a microfluidic fluid, it is assumed that the flow velocity of the microfluidic fluid has been obtained through the aforementioned methods or under known conditions. The operator rotates the sample stage to form several different angles α between the flow direction of the microfluidic fluid and the axial directions of the two probe half-volumes, for example, 0 degrees, 15 degrees, 30 degrees, 45 degrees, 60 degrees, 75 degrees, and 90 degrees. For each angle setting, the corresponding cross-correlation curve and its amplitude are obtained. Since there is a specific trigonometric function relationship between the amplitude of the cross-correlation curve and the angle α, the cross-correlation amplitude reaches its maximum value when the angle α is 0 degrees, i.e., the flow direction coincides with the axial direction of the probe half-volume; as the angle α gradually increases, the cross-correlation amplitude gradually decreases. According to the mathematical expression of the cross-correlation function, the relationship between the cross-correlation amplitude G(α) and the angle α can be expressed as a function proportional to cos²α. Therefore, given the flow velocity V and the relationship curve of the cross-correlation amplitude as a function of angle, the specific value of the flow direction angle α can be calculated using trigonometric functions.

[0035] In specific experimental operations, laser source 1 can be a solid-state laser with a wavelength of 488 nm or 561 nm. Laser source 1 should have good beam quality (M² value close to one) and power stability to meet the requirements of long-term precision measurement. A linear polarizer can be set at the laser output end to adjust the laser intensity to adapt to different types of test samples 4. For the measurement of aqueous samples, the laser power is usually set in the range of 0.1 mW to 1 mW; for the measurement of live cells or live tissue samples, the laser power should be appropriately reduced to avoid photobleaching and phototoxicity effects, and is usually set in the range of 0.01 mW to 0.1 mW.

[0036] The selection of the immersion objective 3 requires comprehensive consideration of factors such as numerical aperture, working distance, and chromatic aberration correction. In a preferred embodiment, the immersion objective 3 is an Olympus 60x immersion objective with a numerical aperture of 1.2 and a working distance of approximately 0.28 mm. The front lens portion of the objective is filled with deionized water to match the refractive index of the sample medium. This immersion objective 3 can focus a 488 nm wavelength laser beam to a lateral beam waist radius of approximately 0.2 μm and an axial beam waist length of 0.8 μm, forming a focusing excitation volume of approximately 0.3 liters.

[0037] The aperture size of the single pinhole 206 has a significant impact on the performance of the measurement system. A smaller pinhole aperture provides higher spatial resolution and a better off-focal rejection rate, but at the same time reduces the fluorescence signal throughput and increases measurement time and signal acquisition difficulty. A larger pinhole aperture can increase signal throughput and shorten measurement time, but at the cost of some spatial resolution. In a preferred embodiment of the present invention, the aperture of the single pinhole 206 is set to 150 micrometers. This size achieves a good balance between spatial resolution and signal throughput, and when used with a water immersion objective with a numerical aperture of 1.2, a lateral spatial resolution on the order of approximately 1 micrometer can be achieved.

[0038] Both the first APD detector 5 and the second APD detector 6 employ avalanche photodiode single-photon detectors, such as the Excelitas SPCM-AQR-14 model. This detector has a detection surface diameter of 175 micrometers, a dark count rate of less than 50 counts per second, a photon detection efficiency of approximately 25% to 65% in the 400 to 900 nanometer wavelength range, and a response time of approximately 500 picoseconds. The detector's output signal is a standard TTL level pulse, with the pulse width proportional to the number of incident photons, allowing direct interface connection with a hardware correlator.

[0039] The hardware correlator can be the Flex02-01D cross-correlator, which has two independent high-speed input channels, each with a sampling rate of up to one billion times per second. It can calculate the cross-correlation function of two signals in real time and transmit the data to the computer via USB interface. The dead time of the cross-correlator is controlled within the nanosecond range, effectively avoiding signal loss and cross-correlation calculation errors caused by dead time.

[0040] The dedicated measurement and analysis software installed on the computer can be developed using platforms such as IGOR PRO, Matlab, or LabVIEW. The software functional modules include system initialization and parameter configuration, data acquisition and real-time display, cross-correlation curve fitting, flow velocity calculation, flow direction angle calculation, and data storage and report generation. During system initialization, the software first establishes communication with the hardware correlator, detects the working status of the APD detector, and loads the optical path alignment parameters. During data acquisition, the software reads cross-correlation data points from the hardware correlator at set time intervals (e.g., every ten milliseconds), plots the cross-correlation curve in real time, and updates the display interface. The curve fitting module uses a nonlinear least squares algorithm to fit the measured cross-correlation curve, obtaining fitted values ​​for key parameters such as diffusion time and flow time. The flow velocity calculation module calculates the flow velocity value of the microfluidic based on the flow time and the distance between the two probe half-volume focal points. The flow direction angle calculation module calculates the flow direction angle based on the cross-correlation amplitude data at different angles combined with the known flow velocity value.

[0041] For different types of samples to be tested, the measurement parameters need to be optimized accordingly. For the measurement of fluorescent molecules in transparent solutions, such as a solution of the fluorescent dye Rhodamine 6G dissolved in water, the measurement time is usually set to 30 to 300 seconds to obtain a sufficient number of photons for statistically reliable cross-correlation curve fitting. For the measurement of fluorescently labeled molecules in living cells, due to the complex intracellular environment, the small number of fluorescent molecules, and the high background noise, the measurement time usually needs to be extended to 300 to 600 seconds, and a higher laser power is required to obtain sufficient fluorescence signal intensity. For in vivo measurements of vascular flow in zebrafish embryos, due to the strong absorption and scattering of laser energy by living tissue, and to avoid photodamage to the biological sample, the laser power should be controlled in the range of 0.01 mW to 0.05 mW, the measurement time should be set to 60 to 120 seconds, and the reliability of the data should be improved by averaging multiple measurements.

[0042] This measurement system has wide-ranging practical applications in biomedical research. In zebrafish embryonic development studies, zebrafish embryos, due to their near-transparent optical appearance, are important model organisms for studying vascular development and hemodynamics. Traditional microparticle imaging velocimetry techniques require the injection of tracer particles into the embryonic circulatory system, which may interfere with embryonic development. The measurement system provided in this invention does not require the introduction of additional tracer particles into the organism; it utilizes only naturally occurring endogenous fluorescent substances in the organism's circulatory system (such as the autofluorescence of hemoglobin or fluorescent dyes introduced through simple staining) to achieve high-precision, non-invasive measurement of blood flow velocity and direction within blood vessels. In practice, anesthetized zebrafish embryos are fixed in a specially designed low-fluorescence background sample chamber. A water immersion objective is used to precisely locate the ventricle or tail vessel region. Once the measurement system is activated, real-time information on blood flow velocity and direction within the vessels can be acquired. This measurement method is non-invasive and allows for repeated measurements of the same embryo at different developmental stages, facilitating the study of the dynamic changes in hemodynamics during vascular development.

[0043] In the field of microfluidic chip laboratory applications, microfluidic chips typically contain multiple microchannels and intersections. The flow behavior of microfluidics within these structures has a decisive impact on chemical reaction efficiency, mixing uniformity, and product separation. The measurement system provided in this invention can be used to rapidly characterize the velocity distribution and flow direction characteristics within the microchannels of a microfluidic chip. By fixing the microfluidic chip under test on a microscope stage and selecting a specific microchannel cross-section or a dual-channel intersection region for measurement, the microfluidic velocity and flow direction information at that location can be obtained. Compared with traditional pressure or temperature sensor measurement methods, the measurement system provided by this invention has advantages such as high spatial resolution, fast time response, and no interference with flow, and can be used in microfluidic chip design verification, process optimization, and quality control.

[0044] In the fields of chemical analysis and drug screening, microfluidic systems are widely used in high-throughput chemical reactions and drug screening experiments. The mixing efficiency and reaction time of reagents within microchannels directly affect the yield and selectivity of chemical reactions. The measurement system provided in this invention allows for real-time monitoring of flow velocity and direction distribution at different locations within the microchannel, thereby optimizing the microchannel geometry and flow parameters to improve mixing efficiency and reaction rate. In drug screening experiments, microfluidic platforms can be used to culture cells and apply drug stimulation. The fluid shear forces in the cell microenvironment have a significant impact on the cell's physiological state and drug response. The measurement system provided in this invention can accurately measure the fluid environment parameters surrounding cells, providing accurate mechanical stimulation data for drug screening experiments.

[0045] In the field of tissue engineering, in vitro cultured tissue-engineered constructs require simulation of the in vivo hydrodynamic environment to promote tissue maturation and functionalization. The efficiency of perfusion culture systems in delivering nutrients and oxygen to tissue-engineered constructs is closely related to fluid parameters. The measurement system provided in this invention allows for precise measurement of microfluidic velocity and flow direction distribution within tissue-engineered constructs, evaluation of the performance of the perfusion culture system, and optimization of fluid perfusion parameters, thereby improving the culture quality and maturity of tissue-engineered constructs.

[0046] The entire optical system adopts an L-shaped compact confocal layout design, achieving spatial orthogonal separation between the excitation and detection optical paths through the spatial arrangement of the first reflecting mirror 201, the second reflecting mirror 204, and the dichroic mirror 202. This orthogonal separation design effectively suppresses crosstalk from the excitation laser to the detection optical path, significantly improving the system's signal-to-noise ratio and ensuring high-sensitivity measurements of live biological samples. All optical components are mounted on an optical platform equipped with an active vibration isolation system to eliminate interference from environmental vibrations. The system's operating temperature range is set to 20 to 25 degrees Celsius to reduce optical path drift and measurement errors caused by changes in ambient temperature.

[0047] In summary, the microfluidic velocity and direction measurement method and system based on single-pinhole spatial fluorescence cross-correlation spectroscopy provided in this invention significantly reduces the operational difficulty and alignment complexity by employing a simplified optical setup with a single laser source and a single pinhole, eliminating the need for the three-dimensional spatial optical alignment process required in traditional spatial dual-beam fluorescence cross-correlation spectroscopy. By using an optical combination of a non-polarizing beam splitter and a lens, two effective detection half-volumes can be generated within the same focusing excitation volume, eliminating the need for additional scanning mirrors and pinholes, thus reducing hardware costs while maintaining the high spatial resolution advantage of fluorescence cross-correlation spectroscopy. Real-time cross-correlation processing of the pulse signals output from the dual-channel APD detector is performed using a hardware correlator, and data acquisition, curve fitting, and result analysis are performed using dedicated computer software, improving data processing efficiency and measurement accuracy. This measurement method is non-invasive for microfluidic systems, particularly suitable for living biological samples, and has broad application prospects in chemical analysis, bioscience research, drug screening, tissue engineering, and developmental biology.

[0048] Although embodiments of the present invention have been shown and described above, it is understood that the above embodiments are exemplary and should not be construed as limiting the present invention. Those skilled in the art can make changes, modifications, substitutions and variations to the above embodiments within the scope of the present invention.

Claims

1. A microfluidic measurement system based on single-pinhole spatial fluorescence cross-correlation spectroscopy, characterized in that: The system includes a laser excitation module, a confocal filtering module, a dual-channel beam splitting detection module, and a signal processing and analysis module. The laser excitation module includes a laser source (1), a first reflector (201), a dichroic mirror (202), and a water immersion objective (3). The laser beam emitted by the laser source (1) changes its propagation direction after passing through the first reflector (201) and is incident on the dichroic mirror (202). After being reflected by the dichroic mirror (202), it enters the water immersion objective (3) and is focused on the sample to be tested (4). The confocal filtering module includes a condenser lens (203), a second reflector (204), a relay lens (205), and a single pinhole (206). The fluorescence signal generated by the sample to be tested (4) is collected by the water immersion objective (3) and passes through the dichroic mirror (202) before passing through the condenser lens (203), the second reflector (204), and the relay lens (205) in sequence. Finally, it is spatially filtered through the single pinhole (206). The dual-channel beam splitting detection module includes a non-polarizing beam splitter (207), a first relay lens (208), a first emission filter (209), a first APD detector (5), a second relay lens (210), a second emission filter (211), and a second APD detector (6). The fluorescence signal after spatial filtering by the single pinhole (206) enters the non-polarizing beam splitter (207) and is divided into two paths. The first path of fluorescence passes through the first relay lens (208) and the first emission filter (209) in sequence and then enters the first APD detector (5). The second path of fluorescence passes through the second relay lens (210) and the second emission filter (211) in sequence and then enters the second APD detector (6). The signal processing and analysis module includes a hardware correlator and a computer. The electrical pulse signals output by the first APD detector (5) and the second APD detector (6) are synchronously transmitted to the hardware correlator for real-time cross-correlation calculation, and the calculation results are transmitted to the computer for processing and analysis.

2. The measurement system as described in claim 1, characterized in that, The non-polarizing beam splitter (207) divides the fluorescence signal passing through the single pinhole (206) into two independent optical path branches. The first relay lens (208) and the second relay lens (210) are located at the two output ends of the non-polarizing beam splitter (207), respectively. The first relay lens (208), the second relay lens (210), and the non-polarizing beam splitter (207) work together to form two axially closely adjacent equivalent detection half-volumes (dashed circles) within the same focusing excitation volume (solid circle).

3. The measurement system as described in claim 2, characterized in that, The axial directions of the two equivalent probe half-volumes form a fixed geometric relationship, and there is a specific geometric ratio between their focal distance and the waist radius of a single focusing excitation volume.

4. The measurement system as described in claim 1, characterized in that, The laser source (1) is a solid-state laser with a wavelength of 488 nm, 561 nm, or other commonly used fluorescent laser wavelengths.

5. The measurement system as described in claim 1, characterized in that, The hardware correlator calculates the positive and negative cross-correlation functions of the two fluorescence signals in real time, and performs calibration by subtracting the positive and negative cross-correlation functions to eliminate the molecular repetition counting error caused by the extremely close axial distance between the two detector half-volumes.

6. The measurement system as described in claim 1, characterized in that, The cross-correlation function of the electrical pulse signals output by the first APD detector (5) and the second APD detector (6) for: exp Where N represents the average number of molecules within the focused excitation volume, and τ represents the time offset. d Let τ represent the diffusion time, K represent the structure coefficient, R represent the ratio between the focal distance between the two probe volumes and the waist radius of a single focusing volume, and τ represent the diffusion time. xf The value represents the flow time of fluorescent molecules from one detector volume to another in the XY two-dimensional plane space, and α represents the angle between the axial direction of the two detector volumes and the microfluidic flow direction.

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