JP2013027605A - Contrast agent for photoacoustic analysis and photoacoustic imaging method using the same - Google Patents

Abstract

PROBLEM TO BE SOLVED: To flexibly measure high contract by improving selectivity of wavelengths in an photoacoustic analysis.SOLUTION: In the photoacoustic analysis, contrast agent containing fine particles which have an absorption peak within a range from 300 to 700 nm and flat shapes is used. Because of performing photoacoustic imaging by using the contrast agent, wavelength in a waveband where absorption of blood vessels or living tissues is weak is used. Thus, in the photoacoustic analysis, selectivity of the wavelength is improved and measurement of high contract is performed more flexibly.

Description

  The present invention relates to a contrast agent used for photoacoustic analysis performed by detecting a photoacoustic wave generated in a subject by irradiating the subject with light, and a photoacoustic imaging method using the same. is there.

  Conventionally, as a method for acquiring a tomographic image inside a subject, an ultrasonic image is generated by detecting ultrasonic waves reflected in the subject by irradiating the subject with ultrasonic waves. Ultrasonic imaging for obtaining a morphological tomographic image is known. On the other hand, in the examination of a subject, development of an apparatus that displays not only a morphological tomographic image but also a functional tomographic image has been advanced in recent years. One of such devices is a device using a photoacoustic analysis method. This photoacoustic analysis method irradiates a subject with light having a predetermined wavelength (for example, visible light, near infrared light, or mid infrared light), and a specific substance in the subject absorbs the energy of this light. A photoacoustic wave, which is the resulting elastic wave, is detected and the concentration of the specific substance is quantitatively measured. The specific substance in the subject is, for example, glucose or hemoglobin contained in blood. Such a technique for detecting a photoacoustic wave and generating a photoacoustic image based on the detection signal is called photoacoustic imaging (PAI) or photoacoustic tomography (PAT).

  In photoacoustic imaging, a contrast agent for photoacoustic imaging for improving detection sensitivity and contrast may be used (Patent Document 1 and Patent Document 2). The contrast agent administered into the living body is distributed in the living tissue to be observed, absorbs the light energy irradiated to the tissue, and generates a photoacoustic wave. Therefore, contrast agents for photoacoustic imaging can increase the apparent extinction coefficient of the tissue in which they exist, that is, the photoacoustic wave derived from the contrast agent can be added to the photoacoustic wave derived from the endogenous tissue. This improves the detection sensitivity of the target biological tissue.

JP 2008-297289 A JP 2008-528114 A

  However, the conventional contrast agents have a problem that the width of the light absorption peak is narrow and the wavelength selectivity is lacking in relation to a living tissue having an absorption peak such as a blood vessel. In such a case, for example, the light energy that should be absorbed by the contrast agent may be absorbed by a living tissue that is not an observation target and become noise. Further, for example, when changing the wavelength of light, it may be necessary to separately administer a contrast agent corresponding to the new wavelength.

  The present invention has been made in view of the above-mentioned problems, and in photoacoustic analysis, a contrast agent for photoacoustic analysis that can improve wavelength selectivity and perform high-contrast measurement more flexibly, and An object of the present invention is to provide a photoacoustic imaging method used.

  In order to solve the above-mentioned problems, a contrast agent for photoacoustic analysis according to the present invention is characterized by containing fine particles having an absorption peak in a range of 300 to 700 nm and having a flat shape.

  And in the contrast agent which concerns on this invention, it is preferable that the ratio of the length of the flat surface with respect to the thickness of microparticles | fine-particles is 2.0-20.

  In the contrast agent according to the present invention, the thickness of the fine particles is preferably 0.005 to 1 μm, and the length of the flat surface is preferably 0.1 to 2 μm.

  In the contrast agent according to the present invention, the fine particles are selected from the group consisting of Al, Mn, Si, Mg, Cr, Ni, Mo, Cu, Fe, Co, Zn, Sn, Ag, Au, Ti, and Zr. It is preferable that it is comprised from at least 1 type.

  In the contrast agent according to the present invention, the fine particles are preferably composed of Ag.

  Further, the contrast agent according to the present invention may further contain an antibody modified on the surface of the fine particles.

  Alternatively, the contrast agent according to the present invention may further contain a dye modified on the surface of the fine particles.

The photoacoustic imaging method according to the present invention includes:
Operate the light irradiation means to irradiate the measurement light having a wavelength included in the absorption peak of the contrast agent into the subject to which the contrast agent described above is administered,
Detect photoacoustic waves generated in the subject by irradiation of measurement light and convert the photoacoustic waves into electrical signals,
Generate a photoacoustic image based on the electrical signal,
A photoacoustic image is displayed.

In the photoacoustic imaging method according to the present invention, the wavelength of the wavelength that is different from the wavelength of the first measurement light and that is included in the absorption peak of blood is in the same region as the region irradiated with the first measurement light. Activating the light irradiation means to irradiate the second measurement light;
Detecting a photoacoustic wave generated in the subject by irradiation of the second measurement light, and converting the photoacoustic wave into a second electric signal;
Generating a second photoacoustic image based on the second electrical signal;
The first photoacoustic image and the second photoacoustic image can be superimposed and displayed.

Further, in the case where the first photoacoustic image and the second photoacoustic image are displayed in a superimposed manner, the photoacoustic imaging method according to the present invention is an electroacoustic conversion unit so as to irradiate ultrasonic waves into the subject. And
Detecting ultrasonic waves reflected in the subject and converting the ultrasonic waves into a third electrical signal;
Generating an ultrasound image based on the third electrical signal;
The first photoacoustic image, the second photoacoustic image, and the ultrasonic image can be displayed in a superimposed manner.

Alternatively, the photoacoustic imaging method according to the present invention operates the electroacoustic conversion means so as to irradiate the subject with ultrasonic waves,
Detecting ultrasonic waves reflected in the subject and converting the ultrasonic waves into a third electrical signal;
Generating an ultrasound image based on the third electrical signal;
The first photoacoustic image and the ultrasonic image can also be displayed in a superimposed manner.

  Since the contrast agent for photoacoustic analysis according to the present invention includes fine particles having an absorption peak in the range of 300 to 700 nm and having a flat shape, the contrast agent has a broader absorption peak than in the case of a spherical shape. Become. Therefore, the absorption peak of the contrast agent can be provided in a wavelength band where the absorption of blood vessels and biological tissues is weak. As a result, in photoacoustic analysis, it is possible to improve wavelength selectivity and perform high-contrast measurement more flexibly.

  In addition, since the photoacoustic imaging method according to the present invention performs photoacoustic imaging using the contrast agent of the present invention, it is possible to use a wavelength in a wavelength band in which absorption of blood vessels and biological tissues is weak. As a result, in photoacoustic analysis, it is possible to improve wavelength selectivity and perform high-contrast measurement more flexibly.

It is a schematic perspective view which shows the contrast agent for photoacoustic analysis of this embodiment. It is a graph which shows each absorption spectrum of the contrast agent of this embodiment, and the conventional contrast agent. It is the schematic which shows the structure of a photoacoustic imaging device.

  Hereinafter, although an embodiment of the present invention is described using a drawing, the present invention is not limited to this. In order to facilitate visual recognition, the scale of each component in the drawings is appropriately changed from the actual one.

"Contrast agent for photoacoustic analysis"
FIG. 1 is a schematic perspective view showing a contrast agent for photoacoustic analysis of the present embodiment. The contrast agent C of the present embodiment is a fine particle P having an absorption peak in the range of 300 to 700 nm and a flat shape.

  The fine particles P have a flat shape. In the present specification, the “flat shape” means a shape in which the ratio of the length D of the flat surface to the thickness d of the fine particles P is 2 to 20. In addition, the thickness d of the fine particles P and the length D of the flat surface mean average values of the whole particles respectively calculated by measurement with an electron microscope. The ratio of the length D of the flat surface to the thickness d of the fine particles P is preferably 2.3 to 17, and more preferably 2.6 to 15. This is because if the ratio is too small, the effect compared to the case of fine particles having a spherical shape may not be obtained, and if the ratio is too large, production may be difficult. The thickness d of the fine particles is preferably 0.005 to 1 μm, and the length D of the flat surface is preferably 0.1 to 2 μm. The reason for the lower limit is that in order for the fine particle P to interact with light, at least the fine particle needs to have a size of about 1/8 of the wavelength. Therefore, at least 0.1 μm is required in one direction. The reason for the upper limit is that if the particles are too large, they do not enter the details of the living body.

  In the present embodiment, the fine particles P are made of a material having an absorption peak in the range of 300 to 700 nm. In the present specification, the “absorption peak” means a portion having an absorbance higher than the minimum value on the short wavelength side in an absorption spectrum curve of a certain material. Such materials include metal materials such as Al, Mn, Si, Mg, Cr, Ni, Mo, Cu, Fe, Co, Zn, Sn, Ag, Au, Ti and Zr, or oxides or semimetals thereof. Materials. Moreover, these may be used as arbitrary 2 or more types of compounds. The fine particles P are preferably composed of Ag. This is because Ag fine particles are easy to create an arbitrary flat shape.

  Since the contrast agent C of the present invention is a fine particle P having an absorption peak in the range of 300 to 700 nm and having a flat shape as described above, the contrast agent C absorbs as compared with a contrast agent composed of fine particles having a spherical shape. The spectrum becomes broad. Specifically, it is as follows. FIG. 2 is a graph showing respective absorption spectra of Ag fine particles having a flat shape and Ag fine particles having a spherical shape. Note that these two absorption spectra are normalized with respect to the maximum value. From FIG. 2, the absorption peak range L1 of the Ag fine particles having a spherical shape is approximately 260 to 500 nm (width is 240 nm), whereas the absorption peak range L2 of the Ag fine particles having a flat shape is approximately 250 to 630 nm ( It can be seen that the width is 380 nm. Furthermore, the wavelength at which the absorption peak becomes maximum can be shifted to the long wavelength side. In addition, the tendency of the absorption peak of the fine particles having the flat shape to be broader than the absorption peak of the fine particles having the spherical shape does not depend on the material. Therefore, according to the contrast agent C of the present invention, the absorption peak of the contrast agent C can be provided in a wavelength band where the absorption of blood vessels and biological tissues is weak. As a result, in photoacoustic analysis, it is possible to improve wavelength selectivity and perform high-contrast measurement more flexibly.

<Design changes>
In the above embodiment, the case where the contrast medium C is composed of the fine particles P having an absorption peak in the range of 300 to 700 nm and having a flat shape has been described, but the contrast medium C of the present invention is not limited to this.

  For example, the contrast agent C can further include an antibody modified on the surface of the fine particle P. That is, in this case, the contrast agent C is composed of fine particles P and antibodies modified on the surface of the fine particles P. In such a case, it becomes possible to specifically administer the contrast medium C to a living tissue in which an antigen against the antibody exists. Therefore, the presence or absence of an antigen can be measured by photoacoustic analysis via the contrast agent C.

  Further, for example, the contrast agent C can further include a dye modified on the surface of the fine particle P. That is, in this case, the contrast agent C is composed of fine particles P and a dye modified on the surface of the fine particles P. This is effective as a method for causing the contrast medium C to have an absorption peak when the fine particles P themselves do not have an absorption peak in the range of 300 to 700 nm. The method for modifying the pigment on the fine particles P is not particularly limited and can be carried out by a known method. For example, as such a method, as described above, the antibody is modified on the surface of the fine particle P and the dye is chemically bonded via the antigen, or a linker such as a silane coupling agent modified on the surface of the fine particle P. And a method of chemically bonding the dye via

"Photoacoustic imaging method"
Next, an embodiment of the photoacoustic imaging method will be described with reference to the drawings. FIG. 3 is a block diagram showing a basic configuration of the photoacoustic imaging apparatus 10 used in the photoacoustic imaging method of the present embodiment. The photoacoustic imaging apparatus 10 includes an ultrasonic probe 11, an ultrasonic unit 12, and a laser light source unit 13. The photoacoustic imaging apparatus 10 is configured to be able to generate both an ultrasonic image and a photoacoustic image.

  In the photoacoustic imaging method of the present embodiment, the photoacoustic imaging apparatus 10 is used to irradiate measurement light having a wavelength included in the absorption peak of the contrast agent into the subject to which the contrast agent of the present invention has been administered. Operate the light irradiation means, detect the photoacoustic wave generated in the subject by irradiation of the measurement light, convert this photoacoustic wave into an electric signal, generate a photoacoustic image based on the electric signal, An acoustic image is displayed.

  The photoacoustic imaging apparatus 10 includes an ultrasonic probe 11, an ultrasonic unit 12, and a laser light source unit 13. The photoacoustic imaging apparatus 10 is configured to be able to generate both an ultrasonic image and a photoacoustic image.

  The laser light source unit 13 emits laser light to be irradiated on the subject as measurement light. This laser light source unit 13 corresponds to the light irradiation means in the present invention. The laser light source unit 13 includes, for example, a first laser beam having a wavelength included in the absorption peak of blood and a second laser having a wavelength different from the wavelength of the first laser beam and included in the absorption peak of the contrast agent. One or more light sources that generate at least light. As the light source, a light emitting element such as a semiconductor laser (LD), a solid-state laser, or a gas laser that generates a specific wavelength component or monochromatic light including the component can be used. For example, in this embodiment, the laser light source unit 13 includes a flash lamp 35 that is an excitation light source and a Q switch laser 36 that controls laser oscillation. When the control means 34 outputs a flash lamp trigger signal, the laser light source unit 13 turns on the flash lamp 35 and excites the Q switch laser 36. Then, the laser light source unit 13 switches and outputs the first laser light and the second laser light in accordance with instructions from the control means 34.

  In the laser light source unit 13, the wavelength of the first laser light is appropriately determined according to the light absorption characteristics of the substance in the subject to be measured. The hemoglobin in the living body has an optical absorption coefficient that varies depending on its state (oxygenated hemoglobin, reduced hemoglobin, methemoglobin, carbon dioxide hemoglobin, etc.). For example, when the measurement target is hemoglobin in the living body (that is, when imaging blood vessels inside the living body), the living body has good light permeability, and various hemoglobins have light absorption peaks of about 600 to 1000 nm. It is preferable. Further, from the viewpoint of reaching the deep part of the subject, the wavelength of the first laser is preferably 600 to 1000 nm.

  On the other hand, in the laser light source unit 13, the wavelength of the second laser light is appropriately determined depending on the absorption characteristics of the contrast agent.

The laser light source unit 13 preferably outputs pulsed light having a pulse width of 1 to 100 nsec as the first and second laser light. The output of the laser beam is 10 μJ / cm ² to several tens of mJ / cm ² from the viewpoints of propagation loss of laser beam and photoacoustic wave, efficiency of photoacoustic conversion, detection sensitivity of the current detector, and the like. Is preferred. Further, the repetition of the pulsed light output is preferably 10 Hz or more from the viewpoint of the image construction speed. Further, the laser beam may be a pulse train in which a plurality of the above pulsed beams are arranged. The laser light output from the laser light source unit 13 is guided to the vicinity of the ultrasonic probe 11 using light guide means such as an optical fiber, a light guide plate, a lens, and a mirror. The subject is irradiated from the vicinity.

  The ultrasonic probe 11 irradiates an object with ultrasonic waves and detects an acoustic wave propagating through the object. That is, the ultrasonic probe 11 performs irradiation (transmission) of ultrasonic waves to the subject and detection (reception) of the reflected waves of the ultrasonic waves that are reflected back from the subject. Further, the ultrasonic probe 11 also detects photoacoustic waves generated in the subject as the observation object in the subject absorbs the laser light. In this specification, “acoustic wave” means an ultrasonic wave and a photoacoustic wave. Here, “ultrasonic wave” means an elastic wave and its reflected wave generated in the subject due to vibration of the electroacoustic transducer, and “photoacoustic wave” means a subject due to a photoacoustic effect caused by irradiation of measurement light. It means the elastic wave generated inside. For this purpose, the ultrasonic probe 11 has a transducer array including, for example, a plurality of ultrasonic transducers arranged one-dimensionally or two-dimensionally. The ultrasonic vibrator 11 is a piezoelectric element made of a polymer film such as piezoelectric ceramics or polyvinylidene fluoride (PVDF). The ultrasonic transducer 11 has a function of converting a received signal into an electric signal when an acoustic wave is received. This electrical signal is output to the receiving circuit 21 described later. The ultrasonic probe 11 is selected according to the diagnostic region from among sector scanning, linear scanning, and convex scanning.

  The ultrasonic probe 11 may include an acoustic matching layer on the surface of the transducer array in order to efficiently detect acoustic waves. In general, the acoustic impedance of the piezoelectric element material and the living body are greatly different. Therefore, when the piezoelectric element material and the living body are in direct contact with each other, the reflection at the interface is increased and the acoustic wave cannot be detected efficiently. For this reason, an acoustic wave can be efficiently detected by arranging an acoustic matching layer having an intermediate acoustic impedance between the piezoelectric element material and the living body. Examples of the material constituting the acoustic matching layer include epoxy resin and quartz glass.

  The ultrasonic unit 12 includes a reception circuit 21, an AD conversion unit 22, a reception memory 23, a data separation unit 24, a photoacoustic image reconstruction unit 25a, and a detection / logarithmic conversion unit that receives signals from the photoacoustic image reconstruction unit 25a. 26a, a photoacoustic image construction unit 27a for constructing a photoacoustic image, an ultrasonic image reconstruction unit 25b, a detection / logarithm conversion unit 26b for receiving a signal from the ultrasonic image reconstruction unit 25b, and an ultrasound for constructing an ultrasonic image It has a sonic image construction means 27b, an image composition means 28, a transmission control circuit 33, and a control means 34. The control means 34 controls each part in the ultrasonic unit 12.

  The receiving circuit 21 receives the electrical signal of the acoustic wave output from the ultrasonic probe 11. The AD conversion means 22 is a sampling means, which samples the electric signal received by the receiving circuit 21 in synchronization with an AD clock signal with a clock frequency of 40 MHz, for example, and converts it into a digital signal. The AD conversion means 22 samples the electric signal at a predetermined sampling period in synchronization with, for example, an AD clock signal input from the outside.

  The AD conversion means 22 stores the sampled digital signal (sampling data) in the reception memory 23. The sampling data stored in the reception memory 23 is data related to photoacoustic waves (photoacoustic data), data related to ultrasonic waves (ultrasound data), or a mixed data thereof.

  The data separation means 24 separates the sampling data stored in the reception memory 23 into photoacoustic data and ultrasonic data. A method for separating the sampling data is not particularly limited. For example, when the ultrasonic irradiation and the laser light irradiation are performed while being shifted in time, the sampling data can be separated into photoacoustic data and ultrasonic data by dividing the sampling data at a certain time. . In addition, for example, sampling data can be separated into photoacoustic data and ultrasonic data by utilizing the difference in frequency and delay amount related to the photoacoustic data and ultrasonic data. The data separation unit 24 inputs the separated photoacoustic data to the photoacoustic image reconstruction unit 25a, and outputs the ultrasonic data to the ultrasonic image reconstruction unit 25b.

  For example, the photoacoustic image reconstruction unit 25a obtains the photoacoustic data obtained from the output signals of 64 ultrasonic transducers of the ultrasonic probe 11 with a delay time corresponding to the position of the ultrasonic transducer. Addition to generate data for one line (delay addition method). In addition, this photoacoustic image reconstruction means 25a may be reconstructed by the CBP method (Circular Back Projection) instead of the delay addition method. Alternatively, the photoacoustic image reconstruction unit 25a may perform reconstruction using a Hough transform method or a Fourier transform method. The photoacoustic image reconstruction means 25a outputs the photoacoustic data added and matched as described above to the detection / logarithm conversion means 26a.

  The detection / logarithm conversion means 26a generates an envelope of the photoacoustic data output from the photoacoustic image reconstruction means 25a, and then logarithmically converts the envelope to widen the dynamic range. Then, the detection / logarithm conversion means 26a outputs the photoacoustic data subjected to signal processing as described above to the photoacoustic image construction means 27a.

  The photoacoustic image construction unit 27a constructs a tomographic image (photoacoustic image) based on the photoacoustic data of each line subjected to logarithmic transformation. For example, the photoacoustic image construction unit 27a constructs a photoacoustic image by converting the position of the time axis of the photoacoustic data into the position of the displacement axis representing the depth in the tomographic image.

  On the other hand, the ultrasonic image reconstruction unit 25b delays the ultrasonic data obtained from the output signals of the 64 ultrasonic transducers of the ultrasonic probe 11 according to the positions of the ultrasonic transducers, for example. Add by time to generate data for one line (delay addition method). The ultrasound image reconstruction means 25b may perform reconstruction by the CBP method (Circular Back Projection) instead of the delay addition method. Alternatively, the ultrasonic image reconstruction unit 25b may perform reconstruction using a Hough transform method or a Fourier transform method. The ultrasonic image reconstruction means 25b outputs the ultrasonic data added and matched as described above to the detection / logarithm conversion means 26b.

  The detection / logarithm conversion means 26b generates an envelope of the ultrasonic data output from the ultrasonic image reconstruction means 25b, and then logarithmically converts the envelope to widen the dynamic range. Then, the detection / logarithm conversion unit 26b outputs the ultrasonic data signal-processed as described above to the ultrasonic image construction unit 27b.

  The ultrasonic image constructing unit 27b constructs a tomographic image (ultrasonic image) based on the ultrasonic data of each line subjected to logarithmic transformation. For example, the ultrasonic image constructing unit 27b constructs an ultrasonic image by converting the position of the time axis of the ultrasonic data into the position of the displacement axis representing the depth in the tomographic image.

  The control means 34 outputs a flash lamp trigger signal and a Q switch trigger signal to the laser light source unit 13 to emit laser light from the laser light source unit 13. Further, the control unit 34 outputs an ultrasonic transmission trigger signal to the transmission control circuit 33 and causes the probe 11 to output ultrasonic waves. Further, the control unit 34 outputs an AD trigger signal to the AD conversion unit 22 in synchronization with the irradiation of laser light or ultrasonic transmission, and starts sampling in the AD conversion unit 22.

  The control means 34 outputs a flash lamp trigger signal that instructs the laser light source unit 13 to output laser light. Thereby, in the laser light source unit 13, the flash lamp 35 is turned on in response to the flash lamp trigger signal, and laser excitation is started. Thereafter, the control means 34 outputs a Q switch trigger signal at a predetermined timing. Thereby, in the laser light source unit 13, the Q switch of the Q switch laser 36 is turned on in response to the Q switch trigger signal, the laser light is output, and the subject is irradiated with the laser light. The time required from when the flash lamp 35 is turned on until the Q-switched laser 36 is sufficiently excited can be estimated from the characteristics of the Q-switched laser 36 and the like. Instead of controlling the Q switch from the control means 34, the Q switch laser 36 may be turned on in the laser light source unit 13 after the Q switch laser 36 is sufficiently excited. In this case, a signal indicating that the Q switch is turned on may be notified to the ultrasonic unit 12 side.

  The control unit 34 outputs an ultrasonic trigger signal for instructing ultrasonic transmission to the transmission control circuit 33. When receiving the ultrasonic trigger signal, the transmission control circuit 33 transmits ultrasonic waves from the ultrasonic probe 11. The control means 34 outputs a flash lamp trigger signal first, and then outputs an ultrasonic trigger signal. That is, the control means 34 outputs an ultrasonic trigger signal following the output of the flash lamp trigger signal. After the flash lamp trigger signal is output and the subject is irradiated with laser light and the photoacoustic wave is detected, the ultrasonic trigger signal is output and the ultrasonic wave is transmitted to the subject and its reflected wave. Is detected.

  The control means 34 further outputs a sampling trigger signal that instructs the AD conversion means 22 to start sampling. This sampling trigger signal is output after the flash lamp trigger signal is output and before the ultrasonic trigger signal is output, more preferably at the timing when the subject is actually irradiated with the laser light. Therefore, the sampling trigger signal is output in synchronization with the timing at which the control means 34 outputs the Q switch trigger signal, for example. When receiving the sampling trigger signal, the AD conversion means 22 starts sampling the electric signal detected by the ultrasonic probe 11.

  The control unit 34 controls the ultrasonic probe 11, the ultrasonic unit 12, the laser light source unit 13, and the image display unit 14 so that the photoacoustic image is displayed on the image display unit 14 in various forms.

  For example, the control means 34 has the second measurement light having a wavelength different from the wavelength of the first measurement light and included in the blood absorption peak in the same region as the region irradiated with the first measurement light. The ultrasonic wave is controlled so as to irradiate the photoacoustic wave generated in the subject by the irradiation of the second measurement light and convert the photoacoustic wave into a second electric signal. The probe 11 is controlled, the ultrasonic unit 12 is controlled to generate a second photoacoustic image based on the second electrical signal, and the first photoacoustic image, the second photoacoustic image, The image display means 14 is controlled so as to be superimposed and displayed. Thereby, the photoacoustic image about the some structure | tissue in the same imaging area can be simultaneously displayed using the difference in the absorption wavelength of a biological tissue. When irradiating the second measurement light, a contrast agent in which the size of the fine particles (the thickness of the fine particles and the length of the flat surface) is adjusted so as to absorb the second measurement light may be administered.

  Furthermore, in the case where the first photoacoustic image and the second photoacoustic image are superimposed and displayed, the control unit 34 applies ultrasonic waves to the same region as the region irradiated with the first and second measurement lights. The ultrasonic probe 11 is controlled so as to irradiate, the ultrasonic wave reflected in the subject is detected, and the ultrasonic probe 11 is controlled so as to convert the ultrasonic wave into a third electric signal. The ultrasonic unit 12 is controlled to generate an ultrasonic image based on the third electrical signal, and the first photoacoustic image, the second photoacoustic image, and the ultrasonic image are superimposed and displayed. The image display means 14 can also be controlled. Thereby, the photoacoustic image which shows the functional information of a biological tissue, and the ultrasonic image which shows the morphological information can be superimposed and displayed, and the information of a biological tissue can be provided more easily.

  Alternatively, the control unit 34 controls the ultrasonic probe 11 to irradiate ultrasonic waves to the same area as the area irradiated with the first measurement light, and detects ultrasonic waves reflected in the subject. The ultrasonic probe 11 is controlled so as to convert this ultrasonic wave into a third electric signal, and the ultrasonic unit 12 is controlled so as to generate an ultrasonic image based on the third electric signal, The image display means 14 can also be controlled so that the first photoacoustic image and the ultrasonic image are superimposed and displayed.

  In this specification, “irradiating the second measurement light or ultrasonic wave to the same region as the region irradiated with the first measurement light” means a photoacoustic image obtained by irradiating the first measurement light. The second measurement light or the ultrasonic wave is overlapped at least partially with the imaging range of the photoacoustic image or the ultrasonic image obtained by irradiating the second measurement light or the ultrasonic wave. It means to irradiate.

  The image synthesizing unit 28 synthesizes the photoacoustic image and the ultrasonic image constructed in the image constructing units 27a and 27b, respectively. More specifically, the image synthesizing unit 28 synthesizes the photoacoustic image of the blood vessel based on the irradiation with the first laser light and the photoacoustic image of the contrast agent based on the irradiation with the second laser light, The photoacoustic image of the blood vessel, the photoacoustic image of the contrast agent, and the ultrasonic image are superimposed and synthesized, or the photoacoustic image and the ultrasonic image of the contrast agent are superimposed and synthesized. In the case where the synthesized image is not displayed, the photoacoustic image and the ultrasonic image may be output from the image synthesizing unit 28 without being synthesized. Further, the image synthesizing unit 28 performs necessary processing on the synthesized image and generates a final image (display image) to be displayed on the image display unit 14.

  The image display unit 14 displays the display image generated by the image synthesis unit 28.

  As described above, since the photoacoustic imaging method according to the present invention performs photoacoustic imaging using the contrast agent of the present invention, it is possible to use a wavelength in a wavelength band in which absorption of blood vessels and biological tissues is weak. As a result, in photoacoustic analysis, it is possible to improve wavelength selectivity and perform high-contrast measurement more flexibly.

DESCRIPTION OF SYMBOLS 10 Photoacoustic imaging device 11 Ultrasonic probe 12 Ultrasonic unit 13 Laser light source unit 14 Image display means 21 Reception circuit 22 AD conversion means 23 Reception memory 24 Data separation means 25a Photoacoustic image reconstruction means 25b Ultrasonic image reconstruction Means 27a Photoacoustic image construction means 27b Ultrasound image construction means 28 Image composition means 33 Transmission control circuit 34 Control means C Contrast agent D Flat plane length d Fine particle thickness P Fine particles

Claims (11)

A contrast agent for photoacoustic analysis,
A contrast agent comprising fine particles having an absorption peak in a range of 300 to 700 nm and a flat shape   2. The contrast agent according to claim 1, wherein the ratio of the length of the flat surface to the thickness of the fine particles is 2.0 to 20. 5.   The contrast agent according to claim 2, wherein the thickness is 0.005 to 1 μm, and the length is 0.1 to 2 μm.   The fine particles are composed of at least one selected from the group consisting of Al, Mn, Si, Mg, Cr, Ni, Mo, Cu, Fe, Co, Zn, Sn, Ag, Au, Ti, and Zr. The contrast agent according to claim 1, wherein:   The contrast agent according to claim 4, wherein the fine particles are composed of Ag.   6. The contrast agent according to claim 1, further comprising an antibody modified on the surface of the fine particles.   The contrast agent according to claim 1, further comprising a modified dye on the surface of the fine particles. A light irradiation means is operated so as to irradiate measurement light having a wavelength included in an absorption peak of the contrast agent into a subject to which the contrast agent according to any one of claims 1 to 7 is administered,
Detecting a photoacoustic wave generated in the subject by irradiation of the measurement light and converting the photoacoustic wave into an electrical signal;
Generating a photoacoustic image based on the electrical signal;
A photoacoustic imaging method, wherein the photoacoustic image is displayed. The same region as the region irradiated with the first measurement light is irradiated with the second measurement light having a wavelength different from the wavelength of the first measurement light and included in the absorption peak of blood. Activate the light irradiation means,
Detecting a photoacoustic wave generated in the subject by irradiation of the second measurement light, and converting the photoacoustic wave into a second electric signal;
Generating a second photoacoustic image based on the second electrical signal;
The photoacoustic imaging method according to claim 8, wherein the first photoacoustic image and the second photoacoustic image are superimposed and displayed. Actuating the electroacoustic conversion means to irradiate ultrasonic waves into the subject;
Detecting the ultrasonic wave reflected in the subject and converting the ultrasonic wave into a third electrical signal;
Generating an ultrasound image based on the third electrical signal;
The photoacoustic imaging method according to claim 9, wherein the first photoacoustic image, the second photoacoustic image, and the ultrasonic image are superimposed and displayed. Actuating the electroacoustic conversion means to irradiate ultrasonic waves into the subject;
Detecting the ultrasonic wave reflected in the subject and converting the ultrasonic wave into a third electrical signal;
Generating an ultrasound image based on the third electrical signal;
The photoacoustic imaging method according to claim 8, wherein the first photoacoustic image and the ultrasonic image are superimposed and displayed.

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