JP2959813B2 - Particle image analyzer - Google Patents

Description

DETAILED DESCRIPTION OF THE INVENTION [Industrial Application Field] The present invention provides a flat sheath flow of a sample liquid such as blood or urine, and irradiates the sample flat stream with strobe light to obtain a still image. For an apparatus that performs analysis such as classification and counting of particle components in a sample solution by image processing, the details are as follows: the imaging area is constantly monitored, and when a particle comes, a strobe is applied to reduce the particle component content. The present invention relates to a particle image analyzer capable of efficiently imaging a particle component even when the number is small.

[Prior art] Conventionally, when examining a particle component contained in a sample such as blood or urine collected from a living body, the sample is prepared by smearing the sample on a slide glass and observing with a microscope. Classification and counting. However, this method is troublesome and has disadvantages such as lack of accuracy. Therefore, an automatic analyzer was developed in order to save labor and increase the accuracy of the inspection. Specific examples thereof are disclosed in JP-A-57-500995,
And U.S. Patent No. 4,338,024. This device uses a sheath liquid as the outer layer, irradiates the sample liquid that has been made extremely flat with strobe light, captures a still image with a video camera, and performs image processing to classify and count the formed material in the sample. Is performed. A multi-item automatic urine analyzer using this technology has already been commercialized.

FIG. 9 shows a schematic view of such a conventional apparatus. The strobe 202 emits light at fixed time intervals in response to a strobe trigger signal output from the image processor 200 at regular intervals. The strobe light is an optical system 204 composed of a lens and the like.
Thus, the sample liquid flowing through the flat channel 208 of the flow cell 206 is irradiated. The sample flows in the direction of the front and back of the paper in the figure. The sample liquid flows in the flat flow path 208 in the vertical direction in FIG. The light transmitted through the flow cell 206 is transmitted to the CCD light receiving surface 2 of the video camera 212 by the optical system 210.
Imaged on 14. A video signal is output from the video camera 212 in synchronization with the genlock signal from the image processor 200, and image processing is performed. In the figure, reference numeral 216 denotes a strobe power supply.

FIG. 10 is an enlarged view of a sample liquid flow portion viewed from the imaging side. In the figure, 220 is a flat flow of the sample, and 222 is a particle component in the sample.

The sample liquid flows rightward in FIG. In the figure 22
Reference numeral 4 denotes a sample liquid flow portion imaged by a video camera. In the figure, reference numerals 226 and 228 denote a sample liquid flow portion imaged last time and a sample liquid flow portion imaged next time, respectively.

[Problem to be Solved by the Invention] Normally, 1/30 second is defined as one frame screen in imaging by a video camera. If light is emitted every 1/30 second and the measurement time is 45 seconds, an image of 1350 can be captured. However, when the content of the particle component in the sample liquid is small, the particle image is not always shown on all the screens. For example, consider the case where blood is used as a sample and white blood cells are analyzed. In order to analyze leukocytes, a blood sample obtained by hemolyzing and destructing erythrocytes and staining the leukocytes is used.

Now, suppose that a blood sample having a leukocyte content of 5000 cells / μ is subjected to the above pretreatment and finally diluted 10-fold, that is, a blood sample containing 500 leukocytes / μ is passed through a flow cell and analyzed. I do. However, the imaging area is one side
150μm square, flat sheath flow thickness 8μ
m (the volume of the imaging area is 150 μm × 150 μm × 8 μ
m = 1.8 × 10 ⁻⁴ μ ).

FIG. 11 is a perspective view of the imaging area. In the figure
230 are white blood cells. Under these conditions, the imaging screen 1
When the number of leukocytes per sheet is calculated, 500 cells / μ × 1.8 × 10 ^-4
μ = 0.09 pieces. In other words, only one white blood cell is captured in 11 screens. Therefore, as described above, even if a total of 1350 screens are obtained, the screen with the white blood cells is 1/11 of that, that is, about 120 in the simple calculation. To increase the number of screens on which white blood cells are shown, (a) decrease the dilution ratio of the sample (increase the concentration), (b) shorten the imaging cycle, and (c) increase the imaging area (volume). And the like.

However, regarding (a), there is a risk of poor hemolysis of red blood cells,
Problems such as an increase in required blood volume occur. Also,
As for (b), there is a special video camera capable of capturing 100 or more screens per second, but it is very expensive. Further, it is necessary to speed up image processing. Further, since the strobe irradiation cycle is shortened, the light amount becomes unstable and the life is shortened. Regarding (c), when the imaging area is widened (when the magnification is reduced), the size of the cell image becomes relatively small. In addition, when the thickness of the flat sheath flow is increased, the number of out-of-focus cell images increases, and the resolution of any cell image is reduced.

As described above, a cell image cannot be obtained efficiently by a method of simply irradiating strobe light every 1/30 second to capture an image.

In order to efficiently obtain a cell image, imaging may be performed by irradiating strobe light when the cell arrives in the imaging area without imaging when the cell is not in the imaging area. For this purpose, a method of providing a detection region for cell detection on the upstream side of the sample liquid flow can be considered. For example, an electric resistance type detection unit including a microhole and an electrode pair and an optical type detection unit including a light emitting element and a light receiving element are conceivable. By doing so, the presence of the cell is detected by the detection unit. However, there is no guarantee that cells that have passed through the detection unit will always pass through the imaging area. In the vicinity of the imaging area, the sample solution flows widely and spreads, and cells may pass out of the imaging area. Further, the time from when the cell passes through the detection unit to when it reaches the imaging area is likely to vary depending on various conditions. For this reason, cells cannot always be imaged at the correct timing.

In view of the above, an object of the present invention is to provide a particle image analyzer that can always efficiently capture an image of a particle component even in a sample having a small content of the particle component.

[Means for Solving the Problems] The particle image analyzer of the present invention flows a sample liquid containing a particle component such as a cell in a flat flow path of a flow cell with a sheath liquid as an outer layer and an extremely flat flow, A first light source for capturing a still image, a first light source for capturing a still image, a first light source for capturing a still image of a sample flow, and performing analysis such as classification and counting of particle components by image processing; In addition to the imaging means, a second light source, a second imaging means, and a control circuit for detecting particles passing through the first imaging area are provided.

The second imaging region formed in the flat flow portion of the sample by the second imaging means is formed in the first imaging region formed in the flat flow portion of the sample by the first imaging means.

The second imaging means is a line sensor in which picture elements are arranged one-dimensionally, and the second imaging area (imaging line) crosses the first imaging area (imaging area) with respect to the flow of the sample. It is formed as follows.

The second light source constantly emits light, and the light passes through the second imaging region and forms an image on the second imaging unit.

The control circuit receives a signal from the second imaging unit, detects the arrival of a particle component, and issues a trigger signal for a strobe as the first illumination unit.

The first light source emits light for a short time in response to the trigger signal, and the light passes through the first imaging region and forms an image on the first imaging unit.

The light of the first light source includes visible light. The light of the second light source does not include visible light. The second light source light is infrared light.

In front of the first and second image pickup means, there is provided a light selection means for selecting light so as to apply only visible light to the first image pickup means and to supply infrared light to the second image pickup means.

In the configuration described in the above, it is preferable to provide a cylindrical lens between the second light source and the flow cell.

In the configuration described in or above, the sample to be measured is blood subjected to hemolysis and staining, and the target particles are white blood cells.

In the configuration described in or above, the sample to be measured may be urine subjected to a staining treatment, and the target particles may be formed components in the urine.

[Operation] The second light source constantly emits light and irradiates the second imaging region. This transmitted light reaches the second imaging unit by the light selection unit. The signal from the second imaging means is sent to the control circuit, and it is determined in real time whether or not particles have arrived at the second imaging area. When the arrival of the particle component is detected, a trigger signal is issued from the control circuit, and the first light source emits light for a short time by the trigger signal. Light from the first light source illuminates the first imaging area for a short time. The second imaging region is formed in the first imaging region. In addition, since the infrared light from the second light source does not reach the first imaging means by the light selection means, only the light from the first light source is selected and reaches the first imaging means. Then, a still image of the particles is captured.

The light from the second light source is narrowed down to an oblong shape by the cylindrical lens, and efficiently irradiates the second imaging region.

Embodiment An embodiment of the present invention will be described with reference to the drawings.

FIG. 1 is a block diagram showing a main part of an embodiment of the present invention. In the figure, reference numeral 22 denotes a flow cell for flowing a blood sample which has been subjected to hemolysis and staining treatment very flatly with the sheath liquid as an outer layer. The sample liquid flows from the front side to the back side perpendicular to the paper surface. The flow path 24 of the flow cell 22 has a dimension of the order of 100 to 200 μm in a direction parallel to the irradiation optical axis (left-right direction in FIG. 1), and in a direction perpendicular to the irradiation optical axis (vertical direction in FIG. 1). It is of the order of a few mm. Therefore, the sample liquid flowing through the flat flow path 24 is also extremely thin (for example, 5 to 10 μm) in the direction of the irradiation optical axis, and extremely wide (for example, several hundred μm) in the direction orthogonal to the irradiation optical axis. Become.

In the figure, reference numeral 10 denotes a first for capturing a still image of a cell.
Reference numeral 36 denotes a first imaging unit. Specifically, the first light source 10 is a strobe, and the first imaging means 36 is a color video camera. In the figure, reference numerals 40 and 48 denote a second light source and a second imaging means for constantly monitoring whether or not cells have come to the imaging area of the still image. More specifically, the second light source 40 is a semiconductor laser that emits near-infrared light having a wavelength of 780 to 830 nm, and the second imaging means 48 is a CCD line sensor in which picture elements are arranged in a line.

The light emitted from the semiconductor laser 40 is a collimator lens
As shown in FIG. 2, the laser beam 60 is narrowed in the direction of flow of the sample liquid 62 and is directed in a direction orthogonal to the parallel direction by collimating the light into parallel light by the lens 42 and strongly converging the light in the direction of flow of the sample liquid by the cylindrical lens 44. Irradiates the first imaging region 64 widely. FIG. 2 is an enlarged view of the sample liquid flow portion viewed from the imaging side in FIG.

FIG. 3 is an enlarged view of the first imaging region.

In FIG. 3, reference numeral 64 denotes a first image pickup area by the first image pickup means 36, and reference numeral 66 denotes a second image pickup area by the second image pickup means 48 (indicated by oblique lines to the left). First imaging area
Numeral 64 denotes a two-dimensional square, each row being a square having a side of 150 μm. The second imaging region 66 is formed one-dimensionally so as to cross the first imaging region 64 orthogonally to the flow direction of the particles. Its width is 1 μm and its length is 150 μm. Hereinafter, the first imaging area is referred to as an “imaging area”, and the second imaging area is referred to as an “imaging line”. The irradiation area 60 by the second light source is shown by oblique lines descending to the right and covers the imaging line 66 widely.
The irradiation area of the first light source is not shown, but naturally covers the imaging area 64 widely.

As described above, by narrowing the light from the second light source into an elongated shape, even a semiconductor laser with a small output can efficiently concentrate the light on the imaging line, and the S / N ratio in the light receiving unit can be improved. In addition, since a change in the amount of light in the imaging line direction can be reduced, shading correction may not be required when the signal from the line sensor is binarized later.

In FIG. 1, reference numerals 16 and 30 denote visible light reflecting and infrared light transmitting dichroic mirrors, and reference numeral 20 denotes a condenser lens. Characteristics of this dichroic mirror (incident angle 45
An example of the degree is shown in FIG.

Now, the infrared light transmitted through the flow cell 22 is
6. An image is formed on the line image sensor 48 via the dichroic mirror 30 and the projection lens 46. An image of the imaging line 66 shown in FIG. 3 is formed on the line image sensor 48. A voltage corresponding to the amount of light accumulated for each picture element is sequentially output from the line image sensor 48 and input to the control circuit 50. After the output from the line image sensor 48 is amplified by the control circuit 50, the output is compared with a threshold level and binarized to determine whether or not cells have arrived. Then, when a cell is detected, a strobe trigger signal is output to cause the strobe 10 to emit light.

Since the transmitted light is imaged on the line image sensor 48, the light amount is reduced in the portion where the cell is shown, and the binarized signal of the picture element corresponding to the portion becomes LOW.

FIG. 7 shows this state. In FIG. 7, reference numeral 70 denotes a white blood cell. In order to monitor all cells passing through the imaging area with the line image sensor 48, the distance that the target cell moves within the scanning cycle time of the line image sensor 48 must be smaller than the size of the cell. No. That is, it is necessary to keep the moving speed of the cells to some extent or less.

The imaging range per pixel in the line direction of the line image sensor 48 is 1 μm, and the size of white blood cells is 15 μm.
μm. Simply, the binary signal sequence is LOW 15 times in a row.
Becomes However, in practice, the size and staining of white blood cells are different for each type, or there is a red blood cell membrane (ghost) reduced by hemolysis. Therefore, it is necessary to determine white blood cells in consideration of these factors. A specific example of the cell determination will be described later. The determination in the control circuit 50 is performed in real time, and if it is determined that white blood cells have come, a trigger signal for irradiating strobe light is issued. 54 in FIG.
Is a strobe power supply.

The strobe light from the strobe light source 10 is a collimator lens
12, dichroic mirror 16 via collector lens 14
The light is condensed by the diaphragm 18 and the condenser lens 20 and is irradiated to the imaging area. The transmitted light is the objective lens
The light is reflected by a dichroic mirror 30 via 26, is condensed by a projection lens 34 via an infrared cut filter 32, and is imaged on a light receiving surface 38 in the video camera. Since the imaging line is formed in a line so as to cross the imaging area, when the cells are detected and the white blood cells are imaged by the video camera, the white blood cell 70 in the imaging screen 68 is captured as shown in FIG. Is limited to the hatched area. For this reason, it is not necessary to process the entire area of the screen at the time of image processing, so that software or hardware for processing can be simplified, and an effect such as an improvement in processing speed also occurs. The location is random.) The video signals from the video camera 36 are mutually synchronized by a genlock signal from the image processor 52, and the captured image signals are sent to the image processor 52 to perform various image processing.

Next, determine whether the cells have arrived at the imaging area,
The control circuit 50 that performs control for strobe irradiation will be described.

 FIG. 6 is a circuit diagram of the control circuit 50.

The light imaged on each picture element of the line image sensor (CCD image sensor) 48 is photoelectrically converted, and the electric charge is accumulated for the scanning cycle time of the line image sensor 48. The electric charge accumulated for each picture element is transferred in synchronization with the transfer clock, and a voltage corresponding to each electric charge is output from the line image sensor 48. This output voltage is
, And is compared with a certain appropriate threshold level by a comparator 74, and HIGH (hereinafter simply referred to as “H”)
Say ) Or LOW (hereinafter simply referred to as “L”).
Valued. The appropriate threshold level is a level determined so that the cell image can be distinguished from the background.

FIG. 7 shows a state in which white blood cells 70 flow from above in the figure and reach the imaging line 66 portion. The binarized signal is L due to the white blood cell signal crossing the imaging line.

 Returning to the description of FIG.

The binarized signal, which has become H for the picture element of the part where the cell image is not formed and becomes L for the picture element of the part where the cell image is formed, is sent to the cell detection circuit 82, When a cell is detected, a cell detection signal is output. The cell detection signal is sent to the strobe control circuit 84, and at that time, a strobe trigger signal for causing the strobe to emit light is output during the period in which imaging is possible.

The cell detection circuit 82 includes a shift register 76, a cell determination unit 80,
It comprises a counter 78. Binarization signal is sequentially input to the shift register 76, bit by bit in synchronization with a transfer clock to obtain a shifted parallel outputs b ₅ ~b ₀ (bits b ₅ is the most recent of the binary signal, the following b ₄ , b ₃ ,
It becomes old in the order of b ₂ , b ₁ , b ₀ . ). Or Figure 6 of 78 counts up the transfer clock by the bit b ₅ as an example be a counter, is performed of the control hold not count up. If the bit b ₅ is H is held, when the L is a count-up. Since the binary signal is synchronized with the transfer clock, the counter 78 counts the number of picture elements where the binary signal is L,
The counted value is output as 4-bit data Q ₃ , Q ₂ , Q ₁ , Q ₀ and input to the cell determination unit 80. When the count value becomes equal to or more than a predetermined value, a cell detection signal is output from the cell determination unit 80. For example, to generate the cell detection signal HIT when the count value is 9 or more, the following logic is used in the cell determination unit 80.
Cell detection signal _{HIT = (Q 3 × Q 0} ) + (Q 3 × Q 1) + (Q 3 ×
Q ₂ ), where Q ₃ is the most significant bit.

The counter 78 is cleared by the shifted parallel data.
The b ₅ ~b ₀ is performed by clearing signal generated based on. The counter 78 performs counting until it is cleared.
When the output of the picture element corresponding to the background portion is output, the clear signal becomes L and the counter 78 is cleared. When the output of the picture element corresponding to the cell portion is output, the clear signal becomes H and the counter 78 performs counting. The clear signal is generated by the following logic as an example.

In this logic, (b ₅ × b ₄ ) is for making the clear signal active (L) when two consecutive picture elements output the signal of the background portion, Is to set the clear signal to L when four consecutive picture elements output H, L, H, L signals.

Simply, when a white blood cell comes on an imaging line of a line sensor, a cell signal is output from a picture element corresponding to that part. That is, cell signals are output from a plurality of continuous picture elements. Therefore, the number of times that the binarized signal continuously becomes L may be counted. However, in reality, the output of the picture element corresponding to the cell may not always be L depending on the staining state of the nucleus, granules, cytoplasm, and the like (see FIG. 7). Therefore, it is necessary to optimize the above-mentioned threshold level and cell determination logic depending on the type, size, staining state and the like of the target cell. In the present embodiment, an example in which a counter is used in the cell determination unit has been described. However, it is also possible to increase the number of parallel output bits of the shift register without using the counter and determine the parallel data as input. In addition to these, various embodiments can be considered according to the specifications.

The selection signal S input to the cell determination unit 80 shown in the lower right of FIG. 6 selects an optimum determination condition when several types of target cells are considered or when fine adjustment of the determination condition is necessary. Is a signal for The content of the selection signal S is instructed from the image processor 52. Further, the logic of the cell determination unit 80 can be easily realized with one commercially available programmable array logic IC (PAL). The cell detection circuit 82 is a single IC, including the shift register and counter.
You can put it inside.

Next, based on the cell detection signal from the cell determination unit 80, a trigger signal for emitting a strobe light is generated. The strobe is fired by this trigger signal. When a video camera of a type in which one frame screen is composed of an odd field screen and an even field screen is used, it is necessary to use a frame accumulation type video camera. When a field accumulation type video camera is used, one field screen must be one screen, so that the vertical resolution is halved and the resolution is reduced.

When using a frame accumulation type video camera,
The strobe must be fired during even field periods. FIG. 8 is a chart for explaining the timing of strobe irradiation.

Charges accumulated in the odd field accumulation period K ₂₁ is outputted in an odd field period T _21, the even field accumulation period charges accumulated in K ₂₂ is the even field period T
Output at ₂₂ . Therefore, the period K ₂₁ and duration K ₂₂ irradiates flash light during the period T ₁₂ to wrap, odd field screen of the still image is output to the odd field period T _21, the even field screen even field period T
₂₂ to form a one-frame screen. Similarly,
Is irradiated with strobe light during T _22, the still picture is output to the odd field period T ₃₁ and the even field period T _32.

However, when irradiated with strobe the odd field period T ₁₁ and T _21, the odd field screen frame screen becomes the screen obtained is irradiated on the odd field period T _11, the even field screen frame screen, odd field period T _The screen is obtained by irradiating ₂₁ and a correct frame screen cannot be constructed.

Moreover, when irradiated with strobe even field period T ₁₂ and the odd field period T _21, the even field screen frame screen becomes double exposed screen.

As described above, when one frame screen is composed of the odd field screen and the even field screen, there is a period in which the strobe can be irradiated and a period in which the strobe cannot be irradiated. There is a need.

For this reason, it is necessary to inhibit strobe irradiation during the odd field period. Further, it is necessary to prohibit irradiation more than once during one even field period.

FIG. 6 shows a strobe control circuit 84 for generating a strobe trigger signal when a cell detection signal is detected in a state where correct imaging is possible. In the figure, reference numeral 86 denotes a D-type flip-flop. The Q output of the flip-flop 86 becomes H due to the rise of the vertical field synchronization signal generated at the start of the even field. A strobe trigger signal is obtained by ANDing the Q output signal and the cell detection signal (H). Since this strobe trigger signal is fed back to the clear terminal of the flip-flop 86, when the strobe trigger signal is output, the strobe trigger signal is cleared and the Q output becomes L. Therefore, even if a cell detection signal is obtained again during the same even field period, no strobe trigger signal is output. The flip-flop 86 is also cleared by the vertical synchronizing signal at the start of the odd field period, so that the Q output of the flip-flop 86 becomes L and no strobe trigger signal is output.

The Q output of the flip-flop 86 becomes H by the next even field vertical synchronization signal.

 Next, usefulness when the present invention is implemented will be described.

The probability density function of the space between particles flowing in the sheath flow is generally expressed by the following equation.

^{_{f β (t) = βe -βt}} ......... (1) However, t is the particle interval. β is a constant determined by various conditions.

When normalized by the average value t _av of the particle intervals in the equation (1), the probability density function f (t) is represented by f (t) = e− ^t (2), which simplifies the handling. Hereinafter, t indicates a ratio to t _av .

If the probability that the particle interval is equal to or greater than t _i is F (t _i ), Is represented by

From the above equation, the probability that particles (cells) pass through the imaging area within one field period is determined.

Here, assuming that the value obtained by dividing the average interval of particles by the field period is t _c , (a) t _c = 3, ie, when the average interval of particles is 1/3 of 1/60 second of one field period, approximation The imaging probability becomes 98.7% by the calculation according to the formula, and (b) t _c = 2, that is, when the average interval of the particles is の of the period of one field, the imaging probability becomes 95%, and (c) t _c = 1 That is, when the average particle interval is equal to one field period, the imaging probability is 77%.

Then raise the examples to obtain the t _c, try to calculate the imaging probabilities of white blood cells. However, the conditions are as follows.

Sample: Blood measurement sample containing 5000 white blood cells / μm: The above blood sample was subjected to hemolysis and staining treatment.
Double-diluted sample (containing 500 leukocytes / μm) Imaging area: 150 μm × 150 μm × 8 μm = 1.8 × 10 ⁻⁴
μ Line sensor scanning cycle: 33 μsec Flat sheath flow rate: 10 μm / 33 μsec = 5 mm / field, where 1 field is 1/60 second The scanning cycle of the line sensor is determined by the response performance of the line sensor and the required light accumulation time. Try to be as short as possible. On the other hand, the flow rate of the flat sheath flow, that is, the moving speed of the particles, needs to be set so as not to move more than the size of the particles during the scanning cycle of the line sensor. This is because if the movement amount is too large, the line sensor cannot clearly detect cells passing through the imaging line.

Now, from the above conditions, during the even field period (1/60
The amount of the sample liquid passing through the imaging area at (sec) is 150 μm × 8 μm × 5 mm = 6 × 10 ⁻³ μm. As a result, the number of white blood cells passing through the imaging area during the even field period is on average 500 cells / μ × 6 × 10 ⁻³ μ = 3.

Therefore, the average value of the interval between adjacent particles is 3 of the field period, and t _c = 3. That is, the imaging probability is 98.7%. This value is a great improvement as compared with the conventional imaging probability of 9% in the conventional method of imaging at a fixed cycle. Assuming that the measurement time is 45 seconds, about 1350 white blood cell images are captured.

A further advantage of the present invention is that a high imaging probability can be maintained even when the imaging area (volume) is reduced. For example, in the above example, the imaging area is 100 μm × 100 μm × 4 μm
= 4 × 10 ⁻⁵ μ and 1 / 4.5, the volume of the sample solution passing through the imaging area during the even field period (1/60 sec) is 100 μm × 4 μm × 5 mm = 2 × At 10 ⁻³ μ, the average number of white blood cells passing through the imaging area during the even field period is 500 / μ × 2 × 10 ⁻³ μ = 1 / field, and t _c = 1. For this reason, the imaging probability is about 77%. In the conventional method, the imaging probability drops greatly to 2% of 1 / 4.5 of the original.

By reducing the imaging area, a relatively large or in-focus cell image can be obtained by reducing the thickness of the sample liquid flow, so that highly reliable analysis can be performed.

By the way, the imaging of the cell image according to the present invention is not performed at a fixed cycle, but the cells are intentionally imaged. Therefore, even if the cell images in the imaging screen are counted, the number of white blood cells per unit volume is correctly obtained. It is not possible. However, the number of white blood cells per unit volume can be obtained by counting the determination result of cells by the line sensor, that is, by counting the number of cell detection signals output from the cell detection circuit 82 in FIG. It is possible to determine whether or not the imaged cells are white blood cells, determine the percentage of white blood cells in the imaged cell image, and calculate a more accurate white blood cell count by multiplying the above cell passage number. If the counting time is 45 seconds in the above example, it is possible to count about 8000 white blood cells, and it may also be able to serve as a blood cell counter. In the conventional method, the number of white blood cells that can be imaged is about 100, and cannot be used as a blood cell counter.

In the above embodiment, the case where blood is used as a sample has been described. However, the analysis target of the apparatus of the present invention is not limited to blood, of course, and the sample may be urine, and particle components in urine (such as blood cells). The present apparatus can also be used when analyzing cells, columns, and the like.

Urine contains particulate components of significantly different sizes. Therefore, in order to efficiently and accurately analyze the urine sample, it is necessary to change the magnification during the measurement and take an image. For details, refer to Japanese Patent Application No. 1-2431, filed by the present applicant.
See Issue 07.

Here, let us consider a case where the present invention is applied only to the high magnification mode (the optical system becomes complicated if it is applied to both the low magnification mode and the high magnification mode).

First of all, the probability of imaging cells is significantly improved. In particular, the effect is large in the high magnification mode. An improvement in the cell imaging probability can be regarded as substantially increasing the number of urine samples to be analyzed, and the analysis accuracy is improved.

In addition, the measurement time in the high-magnification mode should be slightly shortened, and the measurement time in the low-magnification mode should be lengthened to effectively use the limited measurement time, as the amount of urine analyzed in the high-magnification mode increases. Can also.

[Effects of the Invention] The particle image analyzer of the present invention is provided with a second light source, an imaging means, and a control circuit, and is configured to irradiate a strobe when a particle reaches an imaging area to capture a still image. . Therefore, even if the sample has a small particle content, the imaging probability can be significantly improved. For this reason, the analysis accuracy and processing capability of the device are improved.

In addition, the second imaging region for detecting particles is linear, and is formed in the two-dimensional first imaging region so as to cross the region. Therefore, the position where the cell is shown in the imaging screen is limited. Therefore, it is not necessary to process the entire screen at the time of image processing, and it is sufficient to process a limited area, so that the processing can be simplified, speeded, and the like.

Further, even if the volume of the first imaging region is reduced, the imaging probability does not decrease so much. By taking advantage of this, it is possible to reduce the imaging area, that is, to increase the magnification to make the cell image relatively large, or to reduce the thickness of the sample liquid flow to obtain a more focused image. And more reliable image processing and analysis.

Further, the second imaging region has a line shape, and is configured to irradiate the imaging region with light from the second light source in an oblong shape. Therefore, light can be concentrated on the second imaging region, and a signal having a good S / N ratio can be detected. In addition, since a change in light intensity in the line direction can be reduced, signal processing and the like can be easily performed.

[Brief description of the drawings]

FIG. 1 is a block diagram of a main part of a first embodiment of the present invention. FIG. 2 is an enlarged view of a sample liquid flow portion viewed from the imaging side in FIG. FIG. 3 is an enlarged view of a first imaging region. FIG. 4 is a diagram illustrating an example of characteristics of a dichroic mirror. FIG. 5 is a diagram showing imaging by a video camera. FIG. 6 is a block diagram showing details of a control unit 50 according to the first embodiment of the present invention. FIG. 7 is a diagram showing the relationship between the connection by the image sensor and the output waveform of each block in the control unit 50. FIG. 8 is an explanatory diagram of strobe irradiation timing. FIG. 9 is a block diagram of a main part of a conventional device. FIG. 10 is an enlarged view of a sample liquid flow portion viewed from the imaging side of the apparatus of FIG. FIG. 11 is a perspective view of an imaging area. 10 first light source, 38 first imaging means 40 second light source 48 second imaging means 50 control circuit

Claims (7)

(57) [Claims] 1. A flow cell in which a flat flow path for forming a flat flow of a sample liquid containing a particle component to be detected is formed, and a flow path for irradiating the sample liquid flowing in the flow cell with instantaneous light. A first light source; a first imaging unit configured to capture a still image of the particles to be detected in the sample liquid irradiated by the first light source; and a first imaging unit configured to capture image data from the first imaging unit. Processing means for performing a desired analysis process based on: a second light source that constantly irradiates the sample liquid in the flow cell; and a processing unit that irradiates the sample liquid in the flow cell irradiated by the second light source. A second imaging unit for imaging the sample liquid; and detecting that the detected particle component is present in an imaging region of the first imaging unit based on image data from the second imaging unit. , Based on this detection And control means for operating the first light source during a predetermined imaging period of the first imaging means, further comprising:
A second imaging means having a linear imaging area on the flow of the sample liquid, wherein an imaging area of the second imaging means is an imaging area of the first imaging means. A particle image analyzer, wherein the particle image analyzer is formed inside. 2. An image pickup area of said second image pickup means is said first image pickup area.
The particle image analyzer according to claim 1, wherein the particle image analyzer is formed so as to cross the direction in which the sample liquid flows in an imaging region of the imaging unit. 3. An irradiation light of the first light source is a visible light,
3. The particle image analyzer according to claim 1, wherein the light of the second light source is non-visible light. 4. A method according to claim 1, wherein only visible light is applied to said first imaging means.
4. The apparatus according to claim 1, further comprising a light selecting unit that selects the first and second light sources so as to provide only the invisible light to the second imaging unit. 3. The particle image analyzer according to item 1. 5. The particle image analyzer according to claim 1, wherein a cylindrical lens is provided between the second light source and the flow cell. 6. The particle image according to claim 1, wherein said sample solution is blood subjected to hemolytic staining, and said particles to be detected are white blood cells. Analysis equipment. 7. The sample liquid is urine subjected to a staining treatment,
The particle image analyzer according to any one of claims 1 to 6, wherein the particles to be detected are formed components in urine.