JP6732830B2 - Dual modality image processing system for simultaneous functional and anatomical display mapping - Google Patents

Description

This application claims priority to U.S. Patent Application Nos. 13/667,808, filed November 2, 2012, and 13/667,830. The present application is also filed on June 13, 2012 in U.S. Patent Application, entitled "System and Method for Acquisition of Photoacoustic Data and Its Parametric Map Generation" No. 13/507,217, and application dated December 31, 2011 entitled "System and Method for Adjusting the Light Output of an Optoacoustic Imaging System". U.S. patent application Ser. No. 13/341,950, and a continuation-in-part application of U.S. patent application Ser. No. 13/287,759, filed November 2, 2011, entitled "Handheld Optoacoustic Probe". Is. The entire disclosures of these applications, including appendices, are incorporated herein by reference.

At least some embodiments disclosed herein relate generally to biomedical imaging systems, and more specifically to real-time imaging systems for non-invasive visualization of thin tissue sections through the skin. ..

Medical ultrasonic image processing is an image processing technology established for providing diagnostic information based on anatomical analysis and visualizing the tissue morphology of various organs. Photoacoustic imaging is used in medical applications to perform in vivo and in vitro mapping of animal and human tissues and organs based on changes in tissue optical properties. Photoacoustic tomography can provide anatomical, functional and molecular imaging, but the most important value of photoacoustic imaging is its ability to provide quantitative functional information based on the intrinsic contrast of the molecular components of red blood cells. It is in. The essence of functional imaging is to provide the physician with a map of blood distribution and oxygenation so that he can determine whether a particular tissue is functioning normally. For example, a map of the total hemoglobin amount distribution that simultaneously shows a region of increased concentration and a region of decreased oxygen saturation indicates potential malignancy. The essence of molecular imaging is to provide a map of the distribution and concentration of various target molecules for a particular health condition. For example, the distribution of specific protein receptors in cell membranes provides insights into molecular biology or cells that aid in the design of pharmaceuticals and therapeutic methods to treat human diseases.

In one embodiment, the present invention provides a real-time imaging system that non-invasively visualizes thin tissue sections through the skin and also includes independent and superposed 3 containing biomedical important information. Provide a number of images. Specifically, images of deep biological tissue structure are accurately superimposed with images of tissue function status, such as total hemoglobin concentration and blood oxygen saturation. Therefore, the invention of this embodiment combines ultrasonic imaging and photoacoustic imaging techniques in a new manner. Given the complementary nature of the information provided by these, and the fact that the same set of ultrasound/pressure detectors and the same set of analog and digital electronics can be used to obtain both types of signals from tissue, These techniques can be successfully combined. One or more dual-wavelength short pulse-width lasers, or multiple single-wavelength shorts, to achieve a high level of accuracy of quantitative information and to show it in substantially real time (ie, as it occurs substantially) Design techniques utilizing pulse width lasers, fiber optic light transmission systems, handheld imaging probes, and other electronics and processing software are disclosed.

In one embodiment, an image processing system is disclosed for visualization of a section at least to the depth of tissue in a body part. The system includes a processing subsystem that produces three independent images. The image includes two functional images showing the total hemoglobin concentration distribution and the blood oxygen saturation distribution, as well as one morphological image of histology, which is the same handheld imaging probe. Are superposed in time and space. The system provides a three-dimensional function that provides the ability to construct a three-dimensional image of the body from a two-dimensional section made from deep tissue obtained by scanning a handheld probe along at least the surface of a part of the body. A positioning system may be included.

In one embodiment, the image processing method provides simultaneous functional and anatomical display mapping of at least some body tissue. Ultrasound pulses are transmitted into the tissue and backscattered ultrasound signals reflected from various structural tissue boundaries associated with body morphology are detected. Two optical pulses with different spectral bands of electromagnetic energy are transmitted, resulting in a transient superposition resulting from the selective absorption of different energy moieties from each of the two optical pulses by hemoglobin and oxyhemoglobin of blood-containing tissue. A sound wave signal is detected. The detected ultrasonic signal has the process of removing noise, undoing changes in the signal during signal propagation through the tissue and detection system components, and restoring the temporal shape and ultrasonic spectrum of the original signal. Done. The images are reconstructed and processed and co-displayed with a partially permeable functional image of total hemoglobin concentration and blood oxygen saturation to produce a morphological image of the superimposed tissue structure. The process steps described above are repeated at any video frame rate, such that real-time images can display such changes as substantially functional and morphological changes in tissue occur.

The disclosed embodiments are described by way of example with the help of figures in the accompanying drawings, which are not intended to be limiting. In this figure, similar designations indicate similar elements.

FIG. 6 illustrates an embodiment of a photoacoustic probe that irradiates tissue through the skin with a scattered light beam formed in the tissue by merging two light beams. FIG. 6 is a diagram showing how laser irradiation light and an ultrasonic signal from a photoacoustic probe can be scattered from the skin toward the acoustic lens of the probe. Using the light beams on both sides of the ultrasound transducer array and with detection by a transducer tilted at a large angle to the image plane being generated, transverse ultrasound waves induced by laser pulses on the skin are detected. It is a figure of the photoacoustic signal which shows an influence. FIG. 6 illustrates an embodiment of the appearance of image artifacts associated with edge effects of an optical illumination beam having a sharp change in light fluence. FIG. 6 is a diagram of an embodiment in which light illumination of tissue is performed using a handheld photoacoustic probe that transmits light energy from under the photoacoustic probe or from the sides of the probe at different distances. FIG. 6 shows two embodiments of a handheld photoacoustic ultrasound probe protected from light exposure of an acoustic lens. FIG. 8 shows a photoacoustic image using a probe with an acoustic lens that has no optical reflection, and a probe with a gold optically reflective layer that removes lens-related image artifacts. FIG. 6 illustrates an embodiment of a light beam having a sharp edge that may produce edge effects of sound waves and associated artifacts, and a light beam having a smooth edge that reduces the production of edge-related artifacts. .. FIG. 6 shows an output fiber bundle design that includes multiple sub-bundles shaped to provide even illumination of the image plane and reduce edge-related photoacoustic artifacts. FIG. 6 shows an output fiber bundle design that includes multiple sub-bundles shaped to provide even illumination of the image plane and reduce edge-related photoacoustic artifacts. It is a figure which shows the effect of light irradiation with respect to two probes. In this case, the two fiber bundles on each side of each probe are oriented to directly illuminate the skin under the probe. It is a figure which shows the embodiment of the ultrasonic probe which has a flat, concave, or convex arc shape. It is a figure which shows the handheld photoacoustic probe which has a concave arc shape. It is a figure which shows the detail of the handheld photoacoustic probe which has a concave arc shape. It is a figure which shows the photoacoustic image of three spherical objects, and shows that the resolution in the visual field of an arc space (and especially outside) is excellent even in the case of a large object. FIG. 9 illustrates an alternative embodiment of a photoacoustic/ultrasound handheld probe design. Ultrasound transducer impulse response for relatively narrow ultrasonic sensitivity frequency bands, ultrawideband ultrasonic transducer impulse response, and ultrasonic spectrum of transducer sensitivity as a function of frequency for ultrawideband and narrowband resonant oscillators It is a figure which shows the example of. It is a figure which shows the example of the deconvolution of the impulse response of the transducer from the detected photoacoustic signal. Deconvolution restores the original unaltered N-shaped pressure signal. An example of an N-shaped photoacoustic signal in which the original rectangular pressure profile is recovered by addition of all scales corresponding to the range from low frequency to high frequency through a wavelet filter for 5 scale, 7 scale and 9 scale is shown. It is a figure. FIG. 7 is a diagram of radial backprojection where each transducer element aperture is weighted and normalized by the total aperture of the transducer array. FIG. 6 shows an example of a photoacoustic tomographic image of an imaging section through a tissue with a rectangular grid that allows evaluation of system performance for visualization of small arteries, larger veins and microvessels. FIG. 7 is a diagram showing a photoacoustic tomographic image by a point spread function visualized by a flat linear probe using a back propagation algorithm and an aperture normalized back projection algorithm. It is a figure which shows the example of the photoacoustic image of the phantom containing the hair embedded in different depths. In this case, the first image was generated using the standard palette embodiment and the second image was generated using the depth normalized palette embodiment. It is a figure which shows the example of the photoacoustic image of the spherical simulated tumor phantom obtained by the flat linear probe. The wavelength of maximum hemoglobin absorption in completely hypoxic blood (757 nm) and the minimum ratio wavelength of absorption by hypoxic hemoglobin to absorption by oxyhemoglobin present in normoxic blood (1064 nm) 2 FIG. 6 shows an example of tumor discrimination based on absorption coefficients at species wavelengths: 757 nm and 1064 nm. FIG. 8 shows tumor discrimination based on absorption coefficients at two wavelengths of phantoms simulating benign (box-shaped) and malignant (spherical) tumors. It is a figure which shows the photoacoustic image of two crossing tubes which were filled with blood which has different blood concentration [SO2]. FIG. 6 is a photograph of an experimental device containing an artificial blood vessel that was placed in a milk solution and imaged using an arc-shaped photoacoustic probe. FIG. 4 is a 2D cross-sectional image showing a simultaneous anatomical and functional image of a blood vessel, showing six image panels of different anatomical and functional images. FIG. 21 shows the photoacoustic signal amplitude as a function of blood oxygen saturation under laser irradiation with wavelengths of 1064 nm for 21A and 757 nm for 21B (hematocrit is constant). These plots show that blood oxygen saturation can be monitored by photoacoustic imaging. It is a figure which shows the light absorption spectrum of the main tissue chromophore which absorbs light energy in the near-infrared region of hemoglobin, oxyhemoglobin, and water. FIG. 6 shows functional and anatomical imaging of co-displayed breast tumors in a phantom that accurately reproduces the optical and acoustic properties of the average breast with tumor. FIG. 6 is a diagram showing functional and anatomical image processing of simultaneously displayed breast tumors.

(Detailed Description of the Invention)
The following description and figures are for illustration only and should not be construed as limiting. Many specific details are set forth in order to provide a thorough understanding. However, in certain instances, well-known or usual details have not been included in order to avoid obscuring the description. References to one or any embodiment of the disclosure are not necessarily references to the same embodiment, and such references mean at least one.

Reference herein to “an embodiment” or “the embodiment” refers to at least one of the particular features, structures, or characteristics described in connection with the embodiment. It is meant to be included in the embodiment. The appearances of the phrase "in one embodiment" in various places in this specification are not necessarily all referring to the same embodiment, and may differ from each other in other embodiments. Nor is it an alternative embodiment. Moreover, various features are described that may be exhibited by some embodiments, but not by others. Similarly, various requirements are described that are requirements in some embodiments but not requirements in other embodiments.

(System overview)
In at least some embodiments, the present disclosure relates to a medical diagnostic dual modality ultrasound/photoacoustic system using a handheld probe for scanning along the skin surface of an organ, of two types at tissue depth. : Provides anatomical (morphological) and functional (blood hemoglobin index and blood oxygenation index) two-dimensional maps. In one embodiment, these two maps are spatially co-displayed using the same array of ultrasound transducers, faster than any physiological changes that may occur in the tissue being diagnosed, and The two types of images are acquired in real time for simultaneous display in time. The blood hemoglobin index represents the change in blood hemoglobin concentration of the diagnosis target with respect to the background blood concentration in the diagnosis target region. The blood oxygenation index represents the change in blood oxygenation with respect to the background level of blood oxygenation in the area to be diagnosed. These simultaneous display maps can be used to non-invasively differentiate malignant tumors from benign masses and cysts.

In one embodiment, the dual modality ultrasound/photoacoustic system of the present disclosure utilizes a replaceable handheld probe to transfer optical energy and ultrasonic detection of the resulting transient pressure waves for two-dimensional imaging of the body. I will provide a. Of this probe, the first is flat and is used to perform a translational scan of at least a flat portion of the body under study and the second is to perform a translational scan of at least a cylindrical or curved portion of the body under study. With a curved concave arc shape, both scans contribute to a more complete understanding of the body's normal or pathological function.

In one embodiment, at least a portion of the body under study is a molecule such as a blood component such as hemoglobin or oxyhemoglobin responsible for bodily function, or a receptor in cells responsible for cellular function, water, lipids or other components. including.

In one embodiment, the light energy generated using at least one laser beam is used to irradiate the body with light of at least one wavelength. In one embodiment, the light energy is pulsed with a pulse duration that is shorter than the ultrasonic transit time through a distance in the body that is equal to the desired spatial resolution. In one embodiment, the light energy is in the spectral range of 532 nm to 1064 nm. In one embodiment, the light energy is replaced with other electromagnetic energy with a wavelength of 1 nm to 1 m.

In one embodiment, the electronic signal produced by the ultrasonic transducer is amplified using a high noise, wide band electronic amplifier with a high input impedance. In one embodiment, the analog electronic signals are digitized by a multi-channel analog-to-digital converter and further processed utilizing a field programmable gate array. In one embodiment, the ultrasonic transducer is an ultra-wideband transducer that detects ultrasonic signals with no or minimal reverberation. In one embodiment, the system is integrated with an ultrasound imaging system used to enhance visualization of acoustic boundaries of bodies and parts of bodies having different densities and/or speeds of sound.

In one embodiment, the deconvolution of the hardware transfer function, in combination with digital electronic signal processing, yields optical energy in order to obtain the unique photoacoustic amplitude and profile of such signals and the optical absorption coefficient distribution in the body. Propagation and absorption characterization provides a quantitative measurement of target molecule, cell or tissue concentration.

In one embodiment, the photoacoustic contrast agent is used to visualize a portion of the subject's body or to characterize the distribution of particular molecules, cells or tissues in the body.

In one embodiment, the system includes at least a laser, an optical transmission system, a photoacoustic probe, an electronic system, a computer and an image display.

<Laser>
In one embodiment, the laser is capable of emitting short, nanosecond pulses of near infrared radiation at two (or more) different switching wavelengths, ie, two different spectral bands. In one embodiment, one of the wavelengths is preferentially absorbed by blood hemoglobin and the other wavelength is preferentially absorbed by oxyhemoglobin in the blood. In one embodiment, irradiation of the organ under investigation with a first laser pulse of one wavelength (spectral band), and detection of a first photoacoustic signal profile resulting from the first irradiation, followed by a second. Irradiation with a second laser pulse in the wavelength band and detection of a second photoacoustic signal profile to generate a functional map of the diagnostic target area based on (i) blood hemoglobin index and (ii) blood oxygenation index It is possible to obtain data that can be used to reconstruct two simultaneously displayed tomographic images that can be used for.

<Optical transmission system>
In one embodiment, the optical transmission system includes a fiber optic bundle. In one embodiment, the input of the fiber optic bundle is annular to match the incident laser beam, while the output of the fiber bundle is rectangular to match the size and shape of the ultrasound transducer array. In one embodiment, each fiber has a small diameter (eg, reduced to 50 microns) to provide excellent flexibility to the bundle. In one embodiment, the input tip of the fiber bundle is melted and the bundle is shaped into a hexagon to eliminate air gaps between the fibers in the bundle, resulting in as much as 20% improvement in laser energy transfer. In one embodiment, the output tips of the fiber bundles are completely randomized so that the fibers that appear closer together at the input can be output at the output or at different branches of the bifurcated fiber bundle. They will appear apart from each other.

<Photoacoustic probe>
The probe is designed to provide high contrast and resolution for photoacoustic and ultrasound images. In one embodiment, the probe is a handheld probe with an ultrasonic/photoacoustic transducer array, which can be designed to be 1-dimensional, 1.5-dimensional, or 2-dimensional. In one embodiment, the transducer detects acoustic waves within ultrasonic frequencies in the ultra wideband, which are shaped to fit the spectrum of the photoacoustic signal emitted by the tissue to be diagnosed. Also, in one embodiment, the transducer is designed to emit sound waves as short pulse-width ultrasonic waves with short ringdown times and minimal echoes of gradually decreasing magnitude.

In order to realize such a design, the oscillator material is, for example, a piezoelectric ceramic (for example, PZT, PMN-PT, and PZNT), a piezoelectric single crystal (for example, PZT, PMN-PT, and PZNT), a piezoelectric polymer. (Eg, PVDF and copolymer PVDF copolymer), polymer ceramic composites and polymer crystal piezoelectric composites, and electrostatic ultrasonic microvibrators (CMUTs). In one embodiment, the oscillator element thickness, which defines the center frequency and material of the oscillator backing and front matching layers, is optimized.

In various embodiments, the ultrasonic transducer array may be flat or have a concave arc. The flat design is suitable for scanning the surface of the organ under investigation, which has a radius of curvature much larger than the size of the probe, such as the human body. The concave arc shape design provides maximum aperture for photoacoustic signal detection with minimum physical size. Furthermore, the larger aperture provides improved lateral resolution within the viewing angle formed by the straight lines connecting the focal points of the arcs with each end transducer of the array. Arc-shaped probes are often most efficient for scanning curved body surfaces (radius of average size, neck, arms, and legs, etc.) with radii that closely match the radius of the probe.

FIG. 1A shows a structure in which two light beams (OB) emitted from a fiber bundle (FB) passing through an optical window (OW) after being expanded and passed through a light diffuser (LD) are merged into a tissue. 1 illustrates an embodiment of a photoacoustic probe that provides illumination of tissue (TS) through the skin (SK) by a scattered light (SL) beam formed therein. A sound wave (AW) generated in a blood vessel or a tumor (BV or TM) by scattered light (SL) in a tissue propagates through an acoustic lens (AL) to a transducer (TR), and an electric cable (EC). Are converted into electrical signals that are transmitted to the electronic amplifier through the backing material (BM).

In one embodiment, the fiber optic bundle design is as follows. The input of the fiber bundle is rounded with fused fiber tips to avoid light loss due to air gaps between the fibers. The fiber diameter may be about 200 microns for good flexibility and fiber diameters of 100 microns or even 50 microns may be desirable for certain applications. This fiber bundle is Y-branched into two half-bundles and completely randomized so that virtually any two adjacent fibers seen at the input appear in different half-bundles. At least most of the adjacent fibers need to be randomized here. Each half-bundle is preferably divided into a plurality of sub-bundles, each sub-bundle placed in its groove/in place to form the "paddle" of the fiber bundle. Two paddles are placed on either side of the ultrasonic transducer (TR) array assembly. As discussed below, referring to FIGS. 7B and 7C, the output shape of each fiber bundle paddle is rectangular in width of field of view, typically 40 mm, and may have triangular ends. .. Such a triangular shape makes the end portion of the output beam smooth after passing through the light diffuser (LD) of FIG. 1A. Finally, the light beam from the fiber bundle paddle exits the probe and enters the skin (SK) through the optical window (OW). The optical window comprises a thin antireflective coated glass plate or an antireflective coated polymer plate or plastic plate having an acoustic impedance matching the value of the tissue being imaged.

There are several targets in the present photoacoustic probe design: (i) the light should not substantially propagate through the acoustic lens (AL), or optical block acoustic dampers (OBAD) on either side of the probe. Do not propagate through). (Ii) In the acoustic lens or in the optical block acoustic damper material, substantially no absorption of light should produce sound waves. Sound waves in a wide range of ultrasonic frequencies from 0.1 MHz to 15 MHz must be able to pass through the (AL) without attenuation, and neither ultrasonic wave should be able to pass through the OBAD. (Iii) The light beams (OB) emerging through the optical window (OW) must have smooth edges of the light fluence, which light beams will merge due to light scattering within the skin. It must enter the skin as close as necessary to each other to do, and it must enter the array or transducer underlying tissue and provide maximum fluence in the image plane.

In one embodiment, the optical transmission system directs light under the transducer elements rather than through an array of transducer elements. In one embodiment, the design of the photoacoustic probe is arranged as close to the transducer as possible, taking into account the need to focus the ultrasound beam on the depth of the most probable target. Is based on an ultrasonic transducer array with a fiber optic transmission system on both sides of the ultrasonic array, with the dimensions of the elevation axis of as small as possible. In one embodiment, the fiber optic transmission system allows near-infrared light energy to penetrate into the imaged organ, such as the breast, and also minimizes the light-thermo-mechanical interaction of the light beam with the skin. Designed to let you.

Another alternative design of the light transmission system transmits light to a reflector or prism located underneath the ultrasound transducer, which reflects the light at a right angle and directs it at the skin surface of the organ to be imaged. In such an embodiment, a distance may be placed between the transducer element and the skin/tissue. These alternative embodiments can be combined within the scope of the invention.

(Detailed Description of Aspects of System Elements)
<Light irradiation and probe design>
The acoustic lens is typically placed on a transducer within the photoacoustic probe to focus the ultrasonic beam. The probe can be provided without an acoustic lens, but without the lens, the ultrasound transducer could be exposed to direct light and absorb such light, which could then be pulsed out. Especially if this is the case, this may result in ultrasound signals with very large artifacts. Illumination of the lens onto the ultrasound probe produces very intense transient sound waves that lead to image artifacts. Up to 50% of the near infrared can be diffusely scattered by the skin, depending on the skin color. Acoustic impedance mismatch between the lenses of the transducer elements can cause reverberation with long ringdown times. Therefore, embodiments of the probe design include a strongly scattering white opaque lens. If such a lens is not needed due to the curved shape of each transducer element, it is necessary to employ a white strong scattering front matching layer to protect the transducer element from near infrared light.

FIG. 1B shows how the laser radiation 110 and 120 from the photoacoustic probe can scatter 130 from the skin 140 toward the acoustic lens 150 of the probe.

In addition, a (laser) light pulse is produced by a light pulse transversely transverse to the skin surface in a direction perpendicular to the image plane, and a sound wave induced by strong interaction with the skin of the organ being imaged. It may directly affect the ultrasonic transducer. When detected by the transducer array, the spatial distribution of these sound waves is projected on the photoacoustic image at a depth equal to the lateral distance between the transducer array and the light beam on the skin surface, forming artifacts. In addition, the acoustic waves generated in the skin by the reverberation of the acoustic lens and the housing of the probe can further affect the quality of the image processing.

FIG. 2 shows a representative photoacoustic signal showing the effect of lateral ultrasound induced by laser pulses on the skin using light beams on both sides of the ultrasound transducer array. The signal shown is produced by the transducer in a direction approximately perpendicular to the image plane produced therefrom. Such transducers can receive signals at large oblique angles (up to 90°) with respect to the image plane produced therefrom, but this is not desirable. Thus, the transducer array design includes means for rejecting the signal emerging from the image plane. Such means include, but are not limited to, concave arc shapes of transducer elements and acoustic lenses, and transmission of a light beam under the transducer. The 2A detected photoacoustic signal 210 was generated using an effective ultrasonic coupling agent, in this case water. The 2B signal 220 was generated without the acoustic coupling agent, that is, using only the air gap to couple the acoustic signal to the transducer array.

In addition, the finite size of the light beam can affect the acoustic waves generated in response to the light beam's impact on the tissue. Such sound waves may be generated at the sharp ends of the light beam and propagate in the direction of the transducer array, causing artifacts. A system that utilizes a flat linear array of ultrasonic transducers is configured so that the first and last transducers in the array detect these waves first, and the central transducer detects these waves last. As such, this "edge effect" results in V-shaped artifacts on the photoacoustic signal sinogram and low frequency bulk artifacts on the photoacoustic image.

FIG. 3 shows an example of manifestation of V-shaped artifacts 310 on the sinogram 300 of the photoacoustic signal and related artifacts 320 on the photoacoustic image. Since these sound waves are associated with the edge effect of the illumination beam having a sharp change in light fluence, in one embodiment, a V-shaped high intensity signal on the sinogram and a series of associated artifact waves on the photoacoustic image. Can be admitted.

Further, the irradiation shape of the light beam emitted from the photoacoustic probe may affect the image quality. If the light beams of the photoacoustic probe are placed very far apart, this results in dark field illumination (two separate light beams on either side of the probe, so that below the probe in the image plane: There can be a gradual change from no direct light) to brightfield illumination (one beam below the probe that is inserted into the tissue depth along the image plane). This variation creates problems in the image intensity map, defeats the quantitative accuracy of the map, and creates artifacts at the same depth as the initial width between the separate illumination beams on either side of the probe.

FIG. 4 shows an implementation where light illumination of tissue is achieved using handheld photoacoustic probes 410, 420, 430 that transmit light energy either from under the photoacoustic probe or from the side of the probe at different distances. The morphology is shown. In the 4A embodiment, the distribution of light energy in the image plane has a smooth gradient that maximizes at the skin surface when the light beam is transmitted from beneath the ultrasound probe. This optical distribution is useful for obtaining high contrast photoacoustic images. In the 4B embodiment, if the light beam is transmitted near a thin photoacoustic probe, the two beams can be merged by light scattering within the skin, which results in a distribution of light energy in the tissue under the skin. 4A embodiment. In the 4C embodiment, if the light beams are separated by a large distance, they are merged only at a significant depth within the tissue, with dark zones (no light) in the subsurface layer of the tissue, and depth of the tissue. It produces an optical distribution in the image plane where there is a bright zone in it, especially on the photoacoustic image plane of the area of the skin that is brightly illuminated at the same depth as the separation distance of the two beams. In view of the projection on, it is detrimental to the contrast of the photoacoustic image.

Therefore, in the embodiment shown in FIG. 4, the image intensity maps 412, 422 and 432 of the scanned tissue are optimized when the skin illumination is directly below the probe 410. As shown at 420 and 430, as the distance between the center of the transducer and the center of the light beam increases, the image brightness maps 422 and 432 of the scanned tissue become progressively more non-uniform.

Finally, the reflection of laser-induced ultrasound emitted into the tissue after it has been generated in the skin from the boundaries of tissue structures (such as tumors, blood vessels or tissue layers) likewise results in lines, curves, and general It can lead to image artifacts that appear as noise.

In one embodiment, the acoustic lens of the probe is designed so that the lens reflects and scatters light from the illumination element but does not absorb it. Still, it is optically opaque. In various embodiments, such lenses are made with materials that are highly optically scattering, such as silicon rubber filled with titanium dioxide or barium sulfate powder, or with a thin metallic highly reflective layer, such as aluminum or gold or It can be made using a combination of opaque lens material and metal layers. In one embodiment, when a lens of diffuse scattering material and a thin reflective layer (foil) are combined, the lens of diffuse scattering material and the thin reflective layer (foil) are avoided to avoid delamination of the thin metal layer from the front surface of the acoustic lens. ) With a metal reflective layer can be placed between two layers of diffuse scattering material. It is difficult to make a material that has absolutely zero light absorption, and such absorption can generate ultrasonic waves in the thermoelastic material, so the lens material has minimal thermal expansion. It can be made from a thermoplastic material that produces minimal or no ultrasonic waves in response to absorbed light energy.

FIG. 5 shows two embodiments of handheld photoacoustic ultrasound probes 510 and 520, respectively, which are protected from the light exposure of an acoustic lens. In 5A, a fully reflective opaque white lens is utilized, in 5B, a partially reflective white lens is utilized, and the light reflectivity of the lens is enhanced by a gold layer or coating.

FIG. 6 shows a photoacoustic image using a probe 610 with a non-reflective acoustic lens and a probe 620 with a reflective layer of gold. A probe 620 utilizing a gold reflective layer produces an image with reduced artifacts 612 and 614.

In one embodiment, the housing of the probe provides a low echo encapsulation of the probe, which means that the housing of the probe does not absorb (i) laser light (and more specifically near infrared light) (but a small amount of If absorption is unavoidable, it means that it is made of a material that has low thermal expansion that does not emit ultrasonic waves after absorption of laser light) and that (ii) significantly attenuates and suppresses ultrasonic waves and does not reverberate. .. Transducer assemblies within the probe housing are also made from low echo materials. Alternatively, a layer of said low echo material is placed between the transducer assembly and the fiber optic assembly to avoid the generation of any ultrasonic waves during the interaction of the light with the transducer assembly. In various embodiments, such materials are, for example, white porous and anechoic hybrid composites for baffles, foams, polymers, rubbers and plastics (eg, from Micro-Mark, Inc., Berkeley Heights, NJ). CR-600 casting resin available, or AM-37) available from Syntech Materials, Inc. (Springfield, VA). In one embodiment, all such materials are electrically non-conducting insulators, especially to protect the probe from external electromagnetic radiation.

In one embodiment, the light illumination subsystem is configured to transmit a light beam having a smooth intensity edge. In one embodiment, the width of the light beam is equal to the width of the ultrasound transducer array within the photoacoustic probe (eg, about 5 mm). This is achieved by designing the fiber density of the fiber optic bundle to gradually decrease at the ends. This design allows the laser irradiation of the skin of the organ under investigation, so that the beam does not generate ultrasonic waves due to the sharp edges, and such laser-guided ultrasonic waves do not appear in the photoacoustic image sinogram. V-shaped artifact is not generated.

FIG. 7A illustrates a light beam having a sharp edge 710 that can produce ultrasonic edge effects and associated artifacts, and a light beam 720 having a smooth edge with reduced edge associated artifact generation. An embodiment is shown. FIGS. 7B and 7C show that a fiber optic bundle design with a triangular structure at the ends of multiple sub-bundles and fiber bundle assemblies can be used to create a light beam with fluence smooth ends.

In one embodiment, the fiber bundles are placed at a distance from the skin that is sufficient to spread the light beam to the desired width. If the probe dimensions are compact, the fibers used in the fiber bundle can be selected to have a higher numerical aperture (eg, greater than 0.22). In one embodiment, the beam is transmitted through an optical window to provide better coupling of the light beam to the skin. In such embodiments, the optical window contacts the skin and flattens the surface for better light penetration while removing excess coupling gel from the imaged skin surface. In one embodiment, the fiber bundle and the optical window are incorporated into the probe housing, which protects the air gap between the fiber bundle and the optical window.

In one embodiment, the optical window is designed such that interaction of both the light beam and the laser guided ultrasound with such window is minimized. In one embodiment, the windows are very thin and made from an optically clear material with an anti-reflective (AR) optical coating. In one embodiment, such material has anechoic properties. These anechoic ultrasonic properties and the denting of the irradiated skin during photoacoustic scanning results in the suppression of ultrasonic waves that propagate laterally from the laser-irradiated skin surface to the transducer array, resulting in Artifacts are reduced.

In one embodiment, the probe is designed and made as thin as technically possible so that the light beam is very close to the transducer elements on both sides of the ultrasound probe. In one embodiment, the thickness of the probe is very small (eg, 5 mm) and the light beam transmitted to the skin at this distance, d, from the probe center is within the skin thickness (about z=5 mm). To allow the tissue of the organ under investigation to receive one beam under the transducer element. This is called bright field illumination. In one embodiment, the photoacoustic probe is designed such that the optical light beam is transmitted to the skin directly below the transducer element.

FIG. 8 shows the effect of light irradiation of the two probes 810 and 820. The two fiber bundles on both sides of each probe are arranged in a direction that irradiates the skin directly below the probe 812 and the skin on both sides of the probe 822. If the skin is illuminated just below the probe 812, the tumor 814 is clearly visible and no scatter is seen on the image background 816. If the skin is illuminated on both sides of the probe 822, the tumor 824 is not discernible and many artifacts are visible on the image background 826.

In one embodiment, the beam width is designed to deliver increased light to the tissue section to be imaged. Since the non-uniform beam creates sources due to non-uniformity and further creates artifacts in the photoacoustic image, in one embodiment the beam is uniform and has a constant fluence across the beam. The fluence level is defined by the ANSI Laser Safety Standard for Laser Skin Irradiation. The beam width is limited by the light-scattering ability in the tissue for transmitting photons of light to the center of the section (section to be imaged) located under the transducer element. In one embodiment, the length of the light beam is equal to the length of the transducer array. Also, since the sharp edges produce strong edge artifacts on the photoacoustic image, in one embodiment the light beam has smooth edges, ie, a fluence that gradually decreases at the edges.

In one embodiment, the design features of the light illumination system and photoacoustic probe of the present disclosure are summarized in the table below.

In various embodiments, the shape of the integrated ultrasound array for photoacoustic/ultrasound imaging may be flat or convex arc shaped. In one embodiment, the probe shape for photoacoustic image processing is a concave arc shape. Such a concave shape provides a large aperture of minimal physical size of the imaged object, a wider field of view, as well as an improved lateral resolution of the imaged object and a better reconstruction of the shape.

9A-9C show embodiments of a flat or concave arc shape 910 photoacoustic/ultrasound handheld probe (FIG. 9A) and a linear shape 920 handheld transrectal probe (FIG. 9B). FIG. 9C shows a photoacoustic/ultrasound handheld probe design with an ultrasonic transducer assembly, two layers on either side of low echo light reflection and ultrasonic attenuation material, and a surface with two optical windows for transmitting the light beam. Shows the details of.

FIG. 9C shows details of the concave arc handheld photoacoustic probe. An electrical cable 930 is provided for bidirectional communication with the probe and an optical fiber bundle 940 is provided for transmitting light to the probe. The broadband ultrasonic transducer array 950 sends and receives acoustic energy. The transducer array 950 is covered with an opaque white cylindrical lens (not shown for clarity) that scatters and reflects near infrared light over a wide range. Optical window 960 delivers the light beam output. In FIG. 9C of the embodiment, the ultrasound transducers in the probe may be designed to be insensitive to lateral sound waves (ultrasound) and, in particular, to be laterally reverberant. This can be achieved by choosing the piezoelectric composite material, the shape of the piezoelectric ceramic elements in the matrix, and the anechoic properties of the matrix. Also, in one embodiment, the ultrasonic transducer is designed to be highly sensitive within the ultra-wide band of ultrasonic frequencies. This further minimizes the echoes that cause artifacts on the photoacoustic/ultrasound image.

FIG. 9D is a photoacoustic image showing the advantage of the concave arc shape handheld probe from the viewpoint of the resolution of the photoacoustic image. As shown in this embodiment, when the object is in the field of view of the probe aperture, a large spherical shape and sharp edges are clearly shown. Outside the probe aperture, the reproducibility of resolution and shape accuracy is diminished, but it remains better than that of a flat linear probe of similar width.

FIG. 9E shows an alternative embodiment of a photoacoustic/ultrasound handheld probe design that allows two-dimensional imaging in a plane that is parallel to the skin surface at various selected depths. The same applies to the case of a three-dimensional image.

In one embodiment, a handheld probe that scans along the skin surface and produces a real-time two-dimensional image of the tissue under study in the body also has components that serve for accurate global 3D positioning of the probe. This design causes the image processing system to remember the location of all tissue sections and reconstruct a three-dimensional image at the end of the scanning procedure.

<Electronic data acquisition system>
In one embodiment, the present disclosure relates to a photoacoustic imaging system having an electronic data acquisition system that operates in both photoacoustic and ultrasonic modes and can rapidly switch between these modes. In one embodiment, this is accomplished using firmware that controls the functionality of the main microprocessor of the electronic data acquisition system, the field programmable gate array (FPGA). In one embodiment, a re-programmable FPGA allows real-time switching between photoacoustic and ultrasonic operating modes, thus enabling simultaneous display of ultrasonic and photoacoustic images, which can be displayed simultaneously. It can be used for diagnostic imaging based on physical and anatomical maps. In one embodiment, FPGA functionality includes controlling, acquiring, and storing photoacoustic and/or ultrasound data, signal processing and transmission data for real-time image reconstruction and processing. Be done. Also, in one embodiment, FPGAs may be employed for ultrasound beamforming and image reconstruction.

In one embodiment, the electronic data acquisition system design utilizes one or more multi-core graphical processor units (GPUs) for image reconstruction and processing. In ultrasound mode, in one embodiment, the FPGA controls ultrasound transmission and performs ultrasound and photoacoustic data acquisition on a multi-channel board. An external memory buffer may be used to enhance the operability of the memory of the FPGA. In one embodiment, the FPGA is capable of rapidly rewriting ultrasound data acquisition at a repetition rate of about 2-20 kHz signal/frame to photoacoustic data acquisition at a repetition rate of about 10-20 Hz signal/frame. And configured to enable real-time switching of the gate structure and internal memory structure and size between ultrasonic generation and detection, laser synchronization, and system control. In one embodiment, multiple FPGAs may be used to enhance system performance. In one embodiment, in ultrasonic and photoacoustic modes, the FPGA clock can be modified by appropriate time division multiplexing (TDM). In one embodiment, the design of the multi-channel electronic data acquisition system may be module-based, where the modules are typically 16-128 channels, but in some applications 256 channels or more. Sometimes appropriate. In one embodiment, the multi-channel electronic data acquisition system design is a 64-channel configuration.

Separate optoacoustic electronic systems may be combined with separate ultrasonic electronic systems via a single probe to achieve dual modality operation of the optoacoustic/ultrasound system. In one embodiment, the probe has a cable with a Y-branch that connects the probe to the optoacoustic and ultrasound electronic system. In one embodiment, the programmable electronic switch sends the detection signal from the probe (transducer array) to the photoacoustic electronics (to function in the photoacoustic mode) or ultrasound electronics, and further from the ultrasound electronics to the probe. (To work in ultrasonic mode). In one embodiment, synchronization trigger signals are sequentially transmitted to the photoacoustic and ultrasound systems, which in turn acquire photoacoustic and ultrasound images.

<Image processing, reconstruction and display>
Signal Processing In various embodiments, the purpose of the diagnostic imaging procedure is to display each pixel at a brightness that accurately reproduces the signal originally generated at each voxel of the tissue displayed on the image. On the other hand, the intrinsic pressure signal generated by the light pulse in the tissue may be significantly altered during the process of propagation through the tissue, and in particular the detection and recording by the ultrasound transducer and electronics subsystem. ..

In one embodiment, the detected signal is processed to reverse the change and restore the original signal. In one embodiment, such reversal can be accomplished by deconvolution of the system impulse response (IR). In one embodiment, the impulse response can be measured by recording and digitizing a delta function ultrasonic signal generated by a short (nanosecond) laser pulse in a strongly absorptive optical medium of high thermoelastic expansion coefficient.

One component of the impulse response is the acoustoelectric impulse response, which provides photoacoustic or ultrasonic signal distortion due to the properties of ultrasonic transducers, cables and analog electronics. The second component of the impulse response is the spatial impulse response which imparts signal distortion associated with the finite dimensions of the ultrasonic transducer. In various embodiments, large transducers can integrate ultrasonic incidence at all angles, while point source-like transducers can provide perfect or near-perfect delta function space impulse response.

In one embodiment, any distortion that may occur during the acoustoelectric impulse response is reversible by deconvolution of the impulse response from the detected signal. However, possible distortions in the spatial impulse response can be avoided by designing transducers with small dimensions in the image plane. In one embodiment, the dimensions of the transducer are significantly smaller than the shortest wavelength of ultrasonic waves that can be detected or generated by the transducer.

FIG. 10 is a function of the impulse response of the ultrasonic transducer 1010 having a relatively narrow band sensitivity, the impulse response of the ultra wideband ultrasonic transducer 1020, and the frequency of the transducer sensitivity with respect to the ultrawideband and narrowband resonant oscillators 1030. An example of the ultrasonic spectrum is shown below.

In one embodiment, the first step in photoacoustic signal processing in an image processing system that produces a two-dimensional photoacoustic image is the deconvolution of the acoustoelectric impulse response.

FIG. 11 shows an example of deconvolution of the impulse response of the transducer from the detected photoacoustic signal 1110. The deconvolution restores the original, same N-shaped pressure signal 1120 as before.

In one embodiment, the second step in processing the photoacoustic signal is signal filtering to remove noise using a signal filter. In one embodiment, the signal filter is based on a wavelet transform that works in the frequency and time domains simultaneously. In one embodiment, such wavelet filters filter certain frequency components that belong to noise and appear at a given time, while preserving useful signal-like frequency components that appear at different times. Is possible. In one embodiment, the frequency spectrum of the wavelet filter reproduces the frequency band of a typical N-shaped photoacoustic signal, while at the same time providing a smooth window edge that does not cause signal distortion during convolution.

In one embodiment, such wavelet filters are useful in photoacoustic imaging in terms of their ability to restore the original pressure profile created in tissue prior to pressure propagation. In the course of propagation through the tissue, the original positive pressure signal is converted into a bipolar (compression/tension) profile. Therefore, reconstruction of an image of absorbed light energy (photoacoustic image) requires a transformation that starts with a bipolar signal and gives a positive photoacoustic image intensity. In one embodiment, a multi-scale wavelet filter, for example, a filter that simultaneously integrates the signal over time and provides the summation of several frequency bands present in the signal, produces a bipolar pressure signal in the form of thermal energy or first. Can be converted into a unipolar signal representing the positive pressure generated.

FIG. 12 shows, for the 5 scale 1210, 7 scale 1220 and 9 scale 1230, an N-shaped light restored to the original rectangular pressure profile by wavelet filter addition of all scales corresponding to the range from low to high frequencies. An example of an acoustic signal is shown.

In various embodiments, wavelet filtering enables enhancement of objects on the image within a particular size range. The image processing operator (ultrasound technologist or diagnostic radiologist) typically wants better visualization of tumors of particular dimensions and other objects of particular dimensions, eg blood vessels. In one embodiment, the wavelet filter allows the operator to apply a wavelet filter scale of his/her own choice, as well as enhancing objects of a particular size and suppressing objects of other insignificant size. In one embodiment, the boundaries are well visualized for objects of any size, and the scale of the high frequency wavelet is beneficial to the image quality and is included in the scale selection. In one embodiment, a ramp filter can be applied to the signal for mathematically accurate tomographic reconstruction, which can linearly enhance the contribution of higher frequencies.

Image Reconstruction In various embodiments, image reconstruction typically uses radial backprojection of the processed and filtered signal to the image plane. However, due to the limited field of view available from small handheld probes, only incomplete data sets can be obtained. As a result, the 2D photoacoustic image may contain artifacts that distort the shape and brightness of the object displayed on the image. In one embodiment, an aperture integrated normalized radial back projection is used to correct some reconstruction artifacts observed in photoacoustic tomography with limited aperture.

FIG. 13 is an explanatory diagram of radiation backprojection in the case where each transducer element aperture is weighted and normalized with respect to all the apertures of the transducer array.

In one _{embodiment, T k ~T k + 4,} 1311~1315 are transducers 1310 in the array. When visualized by transducer #k, B _i,j is the brightness (intensity) of the pixel at coordinate (i,j) and ω _i,j,k 1320,1330 is emitted by pixel (i,j). It is the angular portion of the photoacoustic wavefront that is generated. When visualized by a full transducer array, Ω _i,j =Σω _i,j,k (sum of all ω _i,j,k ) is the photoacoustic wavefront portion emitted by the optical pixel (i,j). is there. In addition, S _i,j,k are the photoacoustic signal samples measured by the k ^th transducer and used to reconstruct the brightness at pixel (i,j). Various backpropagation algorithms can be used to normalize the photoacoustic image.

In one embodiment, the backpropagation algorithm can be expressed as:

However, in at least some embodiments, aperture-normalized backprojection yields superior images. In one embodiment, the aperture-normalized backprojection can be expressed as:

FIG. 14 is examples 1410 and 1420 of photoacoustic tomographic images of an imaging section of a tumor angiogenesis model. In the first image 1410, the image is normalized using a backpropagation algorithm, such as the first algorithm immediately above. The resulting image has strong, bright arc-shaped artifacts 1412 around blood vessels 1414 near the array surface. In the second image 1420, the image is normalized using an aperture-normalized backprojection algorithm, such as the second algorithm immediately above. As can be seen, the aperture-normalized backprojection algorithm corrects image brightness and reduces arc-shaped artifacts.

FIG. 15 shows a flat straight line using a backpropagation algorithm like the first algorithm just above (1510) and an aperture-normalized backprojection algorithm like the second algorithm just above (1520). It is the example 1510 and 1520 of the photoacoustic tomography image by a point spread function when visualized with a shape probe. As can be seen, the aperture-normalized backprojection algorithm corrects the image brightness and reduces artifacts.

Image Processing and Display In one embodiment, the photoacoustic image palettes are made identical to reduce the effects of light distribution in the tissue. Such equalization changes the dynamic range of the photoacoustic image in order to better visualize both shallow and deep objects.

FIG. 16 is an example 1610 and 1620 of photoacoustic images of a phantom with hair embedded at different depths. The first image 1610 was generated using the standard palette embodiment and the second image 1620 was generated using the depth normalized palette embodiment. As can be seen, the use of depth-normalized pallets enhances the visibility of deeper objects in the illustrated embodiment.

In one embodiment, single photoacoustic image acquisition (different channels) principal component analysis (PCA) is used to remove cross-correlated signal noise. Correlated image clutter can be removed by principal component analysis of the photoacoustic signal dataset. Further, the correlation image clutter can be removed by the principal component analysis of the photoacoustic frame.

FIG. 17 is an example 1710 and 1720 of a photoacoustic image of a phantom of a spherical artificial tumor obtained with a flat linear probe. The first image 1710 is a raw image that has not been subjected to principal component analysis processing. The second image 1720 was subjected to principal component analysis processing by deconvolution of the first principal component. As can be seen, the use of principal component analysis processing enhances the image quality, especially by reducing artifacts.

In one embodiment, the signal and image processing features in the design of the present disclosure can be summarized as in Table 2 below.

Diagnostic Image Reprocessing The principles of functional imaging can be based on tumor pathophysiology. For example, in malignancy, total hemoglobin concentration is increased and oxygen saturation levels in blood hemoglobin are reduced. In one embodiment, the photoacoustic image can be reprocessed and converted, inter alia, into images of (i) total hemoglobin [tHb] and (ii) oxygen saturation of hemoglobin [SO2]. FIG. 18 shows an example of two breast tumors.

FIG. 18 shows two wavelengths, 755 nm (1810) and 1064 nm (1820) (these are wavelengths at which the ratio of absorption by hemoglobin (hypoxic blood) to absorption by oxidative hemoglobin is maximal: 757 nm, and minimal. (Corresponding to a wavelength of 1064 nm), an example of tumor identification based on the absorption coefficient is shown. As can be seen, the malignant tumor (1830) has a higher absorption coefficient at 757 nm than the benign tumor (1840), while the benign tumor (1840) is 1064 nm higher than the malignant tumor (1830). Has an absorption coefficient.

FIG. 19 shows an example of tumor identification by photoacoustic image processing based on absorption coefficients at two wavelengths (1910 and 1920) in a phantom. At 757 nm (1920), the malignant tumor model (1922) is clearly discernible, whereas the malignant tumor model (1922) is not discernible at 1064 nm (1910).

FIG. 20A shows photoacoustic images of two cross-tubes filled with blood having different blood concentrations [SO2] (98% left tube, 31% right tube). The tubes were placed in 1% fat milk with optical properties similar to those found in human breasts. The laser irradiation wavelength used for this image is 1064 nm. FIG. 20B is a photograph of an experimental setup containing an artificial blood vessel placed in a milk solution and imaged with an arc-shaped photoacoustic probe. FIG. 20C is a 2D cross-sectional anatomical and functional image panel of 6 vessels displayed simultaneously: (1-upper left) anatomical ultrasound image of the body with blood vessels; (2-upper right). ) Photoacoustic image acquired at a wavelength of 757 nm; (3-bottom right) Photoacoustic image acquired at a wavelength of 1064 nm; (4-bottom left) Functional image of total hemoglobin amount [tHb]; (5-bottom) (Center) Functional image of blood oxygen saturation [SO2]; (6-Upper center) Functional image of blood oxygen saturation that appears only in the maximum concentration region of total hemoglobin. Raw photoacoustic images in the upper right and lower right panels of FIG. 20C show different brightness of blood vessels containing blood at different levels of total hemoglobin concentration [tHb] and blood oxygen saturation [SO2], but accurate quantification. Measurements can be made as a function of depth under conditions of normalized fluence of light irradiation of body tissue. Using these photoacoustic images, functional images of total hemoglobin content [tHb] and blood oxygenation [SO2] were reconstructed. All functional images shown in FIG. 20C were co-displayed and overlaid with anatomical images of histology for better feature correlation.

FIG. 21 shows the amplitude of the photoacoustic signal as a function of blood oxygen saturation when laser irradiated at a wavelength of 1064 nm for 21A and a wavelength of 757 nm for 21B (hematocrit value is constant). These plots show that blood oxygen saturation can be monitored with photoacoustic imaging. Specifically, this embodiment provides measurements of photoacoustic signal amplitude in blood with varying levels of oxygen saturation (30%-98%) and hematocrit of hemoglobin in erythrocytes (38 g/dL) [tHb]. Shows quantitative data based on. As predicted by the reported blood absorption spectra, the amplitude of the photoacoustic signal at 1064 nm irradiation increases with increasing levels of oxygen saturation, while at 757 nm irradiation wavelength, blood oxygenation increases. The photoacoustic signal amplitude decreases with increasing.

FIG. 22 shows light absorption spectra of major tissue chromophores that absorb light energy in the near infrared range: hemoglobin, oxyhemoglobin and water. The preferred laser wavelengths for functional image processing are 757 nm and 1064 nm, corresponding to maximum and minimum [HHb]/[O2Hb] ratios. On the other hand, a wavelength of 800 nm is optimal for the purpose of calibrating the entire measurement of the total hemoglobin amount [tHb].

FIG. 23 shows functional and anatomical imaging of co-displayed breast tumors in a phantom that accurately reproduces the optical and ultrasound properties of an average breast with tumors. 23A is the same anatomy in which a 2D image of a morphological model of a malignant tumor based on ultrasound (left), a functional image of total hemoglobin concentration (center) and a functional image of blood oxygenation (right) are displayed simultaneously. Shows a biological image. 23B is a 2D image of a benign tumor model, in which a morphological image based on ultrasound (left), a functional image of total hemoglobin concentration (center) and a functional image of blood oxygenation (right) are displayed simultaneously. The same anatomical image.

Figure 24 shows functional and anatomical image processing of breast tumors displayed simultaneously. 24A shows a 2D image of invasive ductal carcinoma of the malignancy with rough boundaries, heterogeneous morphology, high concentrations of total hemoglobin and hypoxia saturation (hypoxia). The morphology of this malignant tumor is shown on the same anatomical image where the ultrasound image in the left image and the blood oxygenation functional image in the middle image and the total hemoglobin concentration functional image in the right image are displayed simultaneously. Based on. 24B shows a 2D image of a breast with a fibroadenoma, a benign tumor with a relatively rounded border, normal oxyhemoglobin concentration and a relatively low total hemoglobin content. The breast morphology is based on ultrasound of the left image and the same anatomical image is displayed simultaneously with a functional image of central blood oxygenation and a right functional image of total hemoglobin concentration.

Conclusion Although some embodiments may be implemented on fully functioning computers and computer systems, various embodiments may be distributed in various forms as computer products and may be used in the actual distribution. It is applicable regardless of type of machine or computer-readable medium.

At least some disclosed aspects can be embodied in software, at least in part. That is, the techniques described herein are specialized applications for processors such as memory, eg, ROM, volatile RAM, non-volatile memory, microprocessors that execute a series of instructions contained in cache or remote storage. Alternatively, it can be implemented using a general purpose computer system or other data processing system.

Routines performed to implement an embodiment include an operating system, firmware, ROM, middleware, service provisioning platform, SDK (software development kit) component, web service, or other specific application, component, program, object, They may be executed as part of a procedure of instructions called a module or "computer program". The activation interface for these routines may be exposed to the software development community as an API (application programming interface). A computer program typically includes repeated one or more instruction sets in various memory and storage devices in a computer that are read by one or more processors in the computer. When executed and executed, it causes a computer to perform operations necessary for execution of elements including various aspects.

Machine-readable media may be used to store software and data, which, when executed by a data processing system, cause the system to perform various methods. Executable software and data may be stored in various locations such as ROM, volatile RAM, non-volatile memory and/or cache, for example. Portions of this software and/or data may be stored on any one of these storage devices. In addition, data and instructions are available from centralized servers or peer-to-peer networks. Different portions of data and instructions are available at different times and in different communication sessions, or in the same communication session, from different centralized servers and/or peer-to-peer networks. The data and instructions are available in their entirety prior to execution of the application. Alternatively, some data and instructions can be dynamically obtained at the timing required for execution. Therefore, it is not necessary to have the entire data and instructions on a machine-readable medium at a particular time.

Examples of computer readable media include, but are not limited to, writable and read-only type media, such as volatile and non-volatile storage, read-only memory (ROM), random access memory (RAM), flash memory, among others. Devices, floppies and other removable disks, magnetic disk storage media, optical storage media (eg, compact disk read only memory (CD ROM), digital versatile disk (DVD), etc.).

Generally, machine-readable media is a form of information that is accessible by a machine (eg, a computer, network device, personal digital assistant, manufacturing tool, any device that comprises a series of one or more processors, etc.). All mechanisms for providing (eg, storing) are included.

In various embodiments, hardwired circuitry may be combined with software instructions to perform the techniques. Thus, the technology is not limited to any combination of hardware circuitry and software, or to any particular source that issues instructions for execution by the data processing system.

Although some figures exemplify some operations in a particular order, operations that are not order dependent may be reordered, other operations may be combined, or suddenly occur. Some reordering or other groupings are specifically mentioned, while others will not be presented as an exhaustive list of alternatives will be apparent to those skilled in the art. Further, it should be appreciated that the stages can be implemented in hardware, firmware, software, or any combination of these.

In the preceding specification, the present disclosure has been described in the context of particular exemplary embodiments. It will be apparent that various modifications can be made without departing from the broader spirit and scope referred to in the following claims. Therefore, the specification and figures are to be regarded in an illustrative rather than a restrictive sense.

Claims (18)

An image processing system for image processing tissue,
A handheld image processing probe comprising an ultrasonic transducer array and a light emitting unit, wherein the light emitting unit outputs a first optical fiber bundle for generating a first light beam and a first light scatterer, and a second light beam. A second optical fiber bundle for generating and a second light scatterer, wherein the first and second optical fiber bundles are arranged on both sides of the ultrasonic transducer array;
A processing system configured to receive data from the handheld image processing probe and process at least three independent images based on at least a portion of the data, the processing comprising: A deconvolution of the impulse response associated with the child array, said impulse response being based on a delta function ultrasound signal generated by a laser pulse in an absorbing optical medium, said three independent images ,
i) a first image reflecting the total hemoglobin concentration distribution;
ii) a second image reflecting the blood oxygen saturation distribution; and iii) a morphological image of the histology, together a processing system configured to include:
An image processing system including. The image processing system of claim 1 , further comprising a signal filter that filters the delta function ultrasound signal to remove noise from the delta function ultrasound signal and is in communication with the processing system. The image processing system according to claim 2 , wherein the signal filter is a wavelet filter based on a wavelet transform that functions simultaneously in the frequency and time domains . The image processing system according to claim 3 , wherein the wavelet filter holds a frequency component of the delta function ultrasonic signal while filtering a specific frequency component that is noise . The image processing system according to claim 3, wherein the frequency spectrum of the wavelet filter reproduces a frequency band of an N-shaped photoacoustic signal . The image processing system of claim 5, wherein the wavelet filter provides a smooth window edge that does not cause signal distortion during convolution . The image processing system according to claim 5 , wherein the wavelet filter converts a bipolar pressure signal into a unipolar signal . The image processing system of claim 1, wherein the impulse response deconvolution restores an N-shaped pressure signal . The image processing system of claim 1 , wherein the handheld image processing probe has a concave arc shape . A handheld image processing probe comprising an ultrasonic transducer array and a light emitting part, wherein the light emitting part comprises
A first optical fiber bundle for generating a first light beam and a first light scatterer; and a second optical fiber bundle and a second light scatterer for generating a second light beam, wherein the first and second optical fiber bundles are provided. Is a handheld image processing probe disposed on both sides of the ultrasonic transducer array,
A first optical window through which the first light beam passes;
A second optical window through which the second light beam passes,
Is equipped with
The first and second fiber bundles have a structure in which the first light beam passes through the first optical window when the first and second optical windows are located near the tissue, Configured to cover the tissue with a second light beam that has passed through the window,
A processing system is configured to receive data from the handheld imaging probe and process independent images based on at least a portion of the data, the processing being associated with the ultrasound transducer array. Comprising deconvolution of the impulse response, said impulse response being based on an ultrasonic signal generated by a short laser pulse in an absorbing optical medium,
Handheld image processing probe. 11. The handheld imaging probe of claim 10, further comprising a signal filter that filters the delta function ultrasound signal to remove noise from the delta function ultrasound signal and is in communication with the processing system . The handheld imaging probe according to claim 10 , wherein the signal filter is a wavelet filter based on a wavelet transform that operates simultaneously in the frequency and time domains . The handheld image processing probe according to claim 12, wherein the wavelet filter retains a frequency component of the delta function ultrasonic signal while filtering a specific frequency component that is noise. The handheld image processing probe according to claim 12, wherein the frequency spectrum of the wavelet filter reproduces a frequency band of an N-shaped photoacoustic signal. 15. The handheld imaging probe of claim 14, wherein the wavelet filter provides a smooth window edge that does not produce signal distortion during convolution. 15. The handheld imaging probe of claim 14, wherein the wavelet filter converts a bipolar pressure signal into a unipolar signal. 11. The handheld imaging probe of claim 10, wherein the impulse response deconvolution restores an N-shaped pressure signal. The handheld imaging probe according to claim 10, wherein the handheld imaging probe has a concave arc shape.