KR102117132B1 - Dual modality imaging system for coregistered functional and anatomical mapping - Google Patents

Abstract

A real-time imaging system that provides mutually matched ultrasound imaging and optoacoustic imaging through the application of the same hand-held probe to generate and retrieve ultrasound and optoacoustic signals. These signals are digitized, processed, and used to reconstruct superimposed anatomical maps and maps of two functional parameters of blood hemoglobin index and blood oxygenation index. The blood hemoglobin index represents changes in blood hemoglobin concentration in areas of diagnostic interest relative to background blood concentration. The blood oxygenation index represents blood oxygenation changes in areas of diagnostic interest against the background level of blood oxygenation. These mutually matched maps can be used to noninvasively differentiate malignant tumors from benign lumps and cysts.

Description

DUAL MODALITY IMAGING SYSTEM FOR COREGISTERED FUNCTIONAL AND ANATOMICAL MAPPING}

This application claims priority to U.S. Patent Application Nos. (13 / 667,808 and 13 / 667,830) filed on November 2, 2012. This application is also filed on June 13, 2012 entitled “System and Method for Acquiring Optoacoustic Data and Producing Parametric Maps Thereof” Partially continued application of U.S. Patent Application No. 13 / 507,217 and entitled, "System and Method for Adjusting the Light Output of an Optoacoustic Imaging System," 2011 U.S. Patent Application No., filed on December 2, 2011, filed December 31, partly filed in U.S. Patent Application No. 13 / 341,950 and entitled "Handheld Optoacoustic Probe" It is a partial continuation application of 13 / 287,759. The entire disclosures of these applications, including their appendices, are incorporated herein by reference.

At least some embodiments disclosed herein generally relate to systems for biomedical imaging, and more particularly to real-time imaging systems that visualize thin tissue flakes non-invasively through the skin.

Medical ultrasound imaging is a well-established imaging technique for visualization of tissue morphology in various organs that provide diagnostic information based on anatomical analysis. Photoacoustic imaging is used in medical applications for in-flight and in-vivo mapping of animal and human tissues and organs based on changes in tissue optical properties. Optoacoustic tomography can provide anatomical, functional, and molecular imaging, but the most important value of optoacoustic imaging is its ability to provide quantitative functional information based on the endogenous contrast of the molecular configurations of red blood cells. . The essence of functional imaging is to provide the physician with a blood distribution and its oxygenation level, so that the physician can determine whether certain tissue functions are normal. For example, a map of total hemoglobin distribution showing areas with increased concentration and reduced oxygen saturation simultaneously indicates potential malignant tumors. The essence of molecular imaging is to provide maps of distributions and concentrations of various molecules of interest to a particular health condition. For example, the distribution of specific protein receptors in cell membranes provides insights into the molecular biology or cells that help develop drugs and treatment methods to treat human diseases.

In one embodiment, the present invention provides a real-time imaging system that visualizes thin tissue flakes non-invasively through the skin and provides three independent, cross-matched images with biomedically important information. Specifically, images of deep biological tissue structures overlap exactly with images of tissue function status, such as blood levels of total hemoglobin concentration and oxygen saturation. Thus, the invention in this embodiment combines ultrasonic imaging and optoacoustic imaging techniques in a novel way. These techniques are advantageously used to acquire both types of signals from tissues, with the complementary features of the information provided by them, and the same set of ultrasonic / pressure detectors and the same set of analog and digital storage devices. It can be combined considering the fact that it can. One or more double-wavelength short-pulse lasers or a plurality of single-wavelength short-pulse lasers, fiber optics, to achieve a high level of accuracy of quantitative information and provide it in substantially real time (ie, substantially when it occurs) A design using a light delivery system, hand-held imaging probe, other electronic hardware and processing software is disclosed.

In one embodiment, an imaging system is initiated for visualization of flakes to the depth of tissue of at least a portion of the body. The system includes a processing subsystem that produces three independent images, including two functional images showing the distribution of total hemoglobin concentration and the distribution of blood oxygen saturation and one morphological image of tissue structures. In other words, the images are inter-matched in time and space by using the very same hand-held imaging probe. The system is a depth of tissue obtained by scanning a hand-held probe along the surface of at least a portion of the body, but a three-dimensional providing the ability to assemble three-dimensional volume images of the body from two-dimensional slices provided. And a positioning system.

In one embodiment, the imaging method provides cross-matched functional and anatomical mapping of tissue of at least a portion of the body. Ultrasound pulses are delivered to 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, and transient ultrasound signals resulting from the selective absorption of different energy ratios from each of the two optical pulses are detected by hemoglobin and oxygen hemoglobin of blood containing tissues. . The detected ultrasonic signals are processed to remove noise, to return signal changes in the process of signal propagation through the tissue and through the detection system components, and to restore the ultrasonic spectrum and temporal shape of the original signals. Image reconstruction and processing is performed to generate morphological images of overlapping tissue structures that are cross-matched with partially transparent functional images of total hemoglobin concentration and blood oxygen saturation. The above steps of the process are repeated with a video frame rate so that real-time images can substantially display tissue functional and morphological changes when they occur.

The disclosed embodiments are illustrated by way of example and not limitation in the figures of the accompanying drawings, in which like reference signs indicate like elements.
1A illustrates one embodiment of an optoacoustic probe having illumination of tissue through the skin by scattered light beams formed in the tissue by merging two optical beams.
1 (b) illustrates how the acoustic signals from the laser illumination light and the optoacoustic probe can be scattered from the skin toward the probe's acoustic lens.
2A and 2B use optical beams on each side of the ultrasonic transducer array and are guided by laser pulses in the skin using detection by transducers tilted at a large angle to the plane of the images generated therefrom. Photoacoustic signals showing the effects of the lateral ultrasounds are illustrated.
3 illustrates an embodiment of the display of image artifacts associated with the edge effect of an optical illumination beam with abrupt changes in optical fluence.
4A-4C are implemented where the optical illumination of the tissue is achieved using a hand-held optoacoustic probe that transmits light energy on the side of the probe or from below the optoacoustic probe at different distances. Illustrate examples.
5A and 5B illustrate two embodiments of a hand-held optoacoustic ultrasonic probe protected from optical illumination of an acoustic lens.
FIG. 6 illustrates optoacoustic images with a probe with an optically reflective layer of gold, using a probe with an acoustic lens that is not completely optically reflective and eliminating lens related image artifacts.
7A illustrates one embodiment of an optical beam with sharp edges that can produce edge effects of sound waves and related artifacts and an optical beam with smooth edges that produce reduced edge-related artifacts.
7B and 7C illustrate designs of an output fiber bundle with multiple sub-bundles shaped to provide uniform illumination of the image plane and reduce edge-related optoacoustic artifacts.
8 illustrates the effect of optical illumination on two probes, with two fiber bundles on each side of each probe oriented to illuminate the skin directly under the probe.
9A illustrates embodiments of ultrasonic probes whose shapes are flat, concave or convex.
9B shows a hand-held optoacoustic probe of concave shape.
9C illustrates details of a hand-held optoacoustic probe that is concave in shape.
9D shows an optoacoustic image of three spherical objects and shows excellent even for large objects within the field of view of arc space (and especially lateral) resolution.
9E illustrates an alternative embodiment of an optoacoustic / ultrasound hand-held probe design.
10A-10C show the transducer sensitivity as a function of the impulse response of an ultrasonic transducer with a relatively narrow ultrasonic frequency band sensitivity, the impulse response of an ultra-wideband ultrasonic transducer, and the frequencies for ultra-wide and narrow-band resonant transducers. Examples of ultrasonic spectra are shown.
11A and 11B provide exemplary examples of deconvolution of the impulse response of transducers from detected optoacoustic signals, where the deconvolution restores the original, unchanged, N-type pressure signals.
12A-12C are wavelet filtered N-type light restored to their original rectangular pressure profile by summing all ratios corresponding to low to blast furnace frequency ranges for 5 scales, 7 scales, and 9 scales. Provides illustrative examples of acoustic signals.
FIG. 13 provides an exemplary diagram of radial backprojection where each transducer element opening is weighted and normalized with respect to the total opening of the transducer array.
14A and 14B provide illustrative examples of photoacoustic tomography images of imaging flakes through tissue with small arteries, larger veins, and rectangular grids that allow estimation of system performance in the visualization of microvessels.
15A and 15B provide illustrative examples of photoacoustic tomography images of a point spread function as visualized with a flat linear probe using a back propagation algorithm and an aperture normalized back projection algorithm.
16A and 16B show the light of a phantom with hairs embedded in different depths where a first image is created using one embodiment of a standard palette and a second image is created using one embodiment of a depth-normalized palette. Provides illustrative examples of acoustic images.
17A and 17B provide exemplary examples of photoacoustic images of a phantom of a spherical simulated tumor obtained with a flat linear probe.
18 fully matches the local maximum of hemoglobin absorption as in hypoxic blood (757 nm) and the minimum of the ratio of absorption by hypoxic hemoglobin versus absorption by oxygenated hemoglobin as in normally oxygenated blood (1064 nm), A diagram illustrating tumor differentiation based on absorption coefficients at two wavelengths (757 nm and 1064 nm) is shown.
19 illustrates tumor differentiation based on absorption coefficients at two wavelengths in the phantom simulating benign (box) and malignant (old) tumors.
20A shows an optoacoustic image of two crossed tubes filled with blood with different levels of blood [SO2].
20B shows a photograph of an experimental setup containing artificial blood vessels positioned in a milk solution and imaged using an arc-type optoacoustic probe.
20C shows cross-matched 2D cross-section atomic and functional images of vascular tubes showing six image panels with different atomic and functional images.
21A and 21B show photoacoustic signal amplitudes as a function of blood oxygen saturation (with constant hematocrit) under laser illumination at wavelengths of 1064 nm in FIG. 21A and 757 nm in FIG. 21B. These plots illustrate that blood oxygen saturation can be monitored with optoacoustic imaging.
22 illustrates light absorption spectra of main tissue chromophores absorbing light energy in the near-infrared range: hemoglobin, oxygen hemoglobin and water.
23A and 23B illustrate cross-matched functional and anatomical imaging of breast tumors in phantoms that accurately replicate the optical and acoustic properties of the average chest with tumors.
24A and 24B illustrate cross-matched functional and anatomical imaging of breast tumors.

The following description and drawings are illustrative and not to be construed as limiting. A number of specific details are described to provide a thorough understanding. However, in certain instances, well-known or conventional details are not described to avoid obscuring the description. References to one or one embodiment in the present disclosure are not necessarily references to the same embodiment, and these references mean at least one.

References herein to “one embodiment” or “an embodiment” mean that a particular feature, structure, or characteristic described in connection with the embodiment is included in at least one embodiment of the present disclosure. The appearances of the phrase (“in one embodiment”) in various places in this specification are not necessarily all referring to the same embodiment, or to separate or alternative embodiments, except for other embodiments. In addition, various features are described that may be seen by some embodiments and not by others. Similarly, various requirements are described, which may be requirements for some embodiments rather than other embodiments.

System overview

In at least some embodiments, the present disclosure uses a hand-held probe to scan along the skin surface of an organ and is of two types: anatomical (morphological) and functional (blood hemoglobin index and blood oxygenation index). It relates to a dual-way ultrasound / photoacoustic system for medical diagnosis that provides dimensional maps to the depth of the tissue. In one embodiment, these two maps are spatially matched by using the same array of ultrasonic transducers and images of the two types in real time, faster than any physiological changes can occur in the tissue of diagnosis of interest. By acquiring them, they are mutually matched in time. The blood hemoglobin index represents changes in blood hemoglobin concentration in diagnostic regions of interest relative to background blood concentration. The blood oxygenation index represents blood oxygenation changes in areas of diagnosis of interest against the background level of blood oxygenation. These mutually matched maps can be used to noninvasively differentiate malignant tumors from benign lumps and cysts.

In one embodiment, the dual-mode ultrasound / photoacoustic system of the present disclosure is one of which is flat and is used to perform a mediation scan through at least a flat portion of the body under irradiation, the second of which is at least the body of the body under irradiation. Two-dimensional of the body with the transfer of optical energy and acoustic detection of the resulting transient pressure using interchangeable hand-held probes that are shaped and curved as concave arcs to perform mediation scans through cylindrical or curved parts. It provides imaging, and both scans contribute to a more complete understanding of normal or pathological functions in the body.

In one embodiment, at least a portion of the body under irradiation is in molecules of blood components such as hemoglobin and oxygen hemoglobin responsible for body functions or cells responsible for cell function, water, lipids or other components. Receptors.

In one embodiment, light energy generated using at least one laser beam has at least one wavelength of light and is used for body illumination. In one embodiment, the light energy is pulsed with a pulse duration shorter than the time of ultrasound propagation through the distance from the body, such as the desired spatial resolution. In one embodiment, the light energy is in the spectral range from 532 nm to 1064 nm. In one embodiment, the light energy is replaced by other electromagnetic energy with wavelengths from 1 nm to 1 m.

In one embodiment, the electronic signals generated by the ultrasonic transducers are amplified using low noise wide band electronic amplifiers with high input impedance. In one embodiment, analog electronic signals are digitized by a multi-channel analog-to-digital converter and further processed using a field programmable gate array. In one embodiment, the ultrasonic transducers are ultra-wideband transducers that detect ultrasonic signals with no or minimal reflections. In one embodiment, the system is integrated with an ultrasound imaging system used to enhance the visualization of acoustic boundaries in the body and parts of the body with different density and / or speed of sound.

In one embodiment, quantitative measurements of concentrations of target molecules, cells, or tissues result in a deconvolution of the hardware transfer function to obtain the distribution of light absorption coefficients in the body and the unique optoacoustic amplitude and profile of these signals. It is achieved through the characterization of optical energy propagation and absorption combined with the processing of digital electronic signals.

In one embodiment, optoacoustic contrast agents are used to visualize a portion of the body of interest or to characterize the distribution of specific molecules, cells or tissues in the body.

In one embodiment, the system includes at least a laser, light transmission system, optoacoustic probe, electronic system, computer and image display.

laser

In one embodiment, the laser can emit short, nanosecond pulses of near infrared light in two (or more) different toggling wavelengths, ie two different spectral bands. In one embodiment, one of the wavelengths is preferentially absorbed by hemoglobin in the blood and the other is preferentially absorbed by oxygen hemoglobin in the blood. In one embodiment, illumination of an irradiated organ having a first laser pulse at one wavelength (spectrum band) and detection of a first optoacoustic signal profile resulting from the first illumination, followed by a second laser in the second wavelength band Detection of the pulsed illumination and the second optoacoustic signal profile is based on (i) blood hemoglobin index and (ii) blood oxygenation index, which can be used to generate functional maps of functional maps of the regions of diagnosis of interest. It can provide data that can be used for reconstruction of tomography images.

Light transmission system

In one embodiment, the light delivery system includes bundles of optical fibers. In one embodiment, the input of the optical fiber bundle is circular to match the incident laser beam, while the output of the fiber bundle is rectangular to match the size and shape of the ultrasonic transducer array. In one embodiment, each fiber has a small diameter (eg, down to 50 microns) to provide good flexibility of the bundles. In one embodiment, the input tip of the fiber bundle is fused to shape the bundle into a hexagon and to remove spaces between fibers in the bundle, thereby providing good transmission up to 20% of the laser energy. . In one embodiment, the output tip of the fiber bundle is completely randomized so that fibers that appear close to each other at the input appear to be far from each other even at different outputs of the fiber bundle, either at the output or in two branches.

Optoacoustic Probe

The probe is designed to provide high contrast and resolution for both optoacoustic and ultrasound images. In one embodiment, the probe is a hand-held probe with an array of ultrasonic / optoacoustic transducers, which can be designed to be single-dimensional, 1.5-dimensional or 2-dimensional. In one embodiment, the transducers detect sound waves within the ultra-wide ultrasound frequencies, and the ultra-wide band is shaped to match the spectrum of optoacoustic signals emitted by the tissue of diagnosis of interest. In one embodiment, the transducers are also designed to emit sound waves as short pulses of ultrasonic waves with shorter ring-down times and minimal reflections of gradually decreasing magnitude.

To achieve this design, the transducer material is, for example, piezoelectric ceramics (such as PZT, PMN-PT, and PZNT), piezoelectric single crystals (such as PZT, PMN-PT, and PZNT), piezoelectric polymers Field (such as PVDF, and copolymer PVDF copolymer), synthetic polymer-ceramic and polymer-crystalline piezoelectric materials and capacitive microfabricated ultrasonic transducers (CMUT). In one embodiment, the thickness of the transducer elements providing the center frequency and materials of the support layer and front surface matching layer of the transducers is optimized.

In various embodiments, the shape of the ultrasonic transducer array may be a flat or concave arc. The flat design is suitable for scanning the surface of the organ under investigation with a radius of curvature much larger than the size of the probe, such as the human body. The concave arc-shaped design provides maximum aperture for optoacoustic signal detection with minimal physical dimensions. The large arc results in improved lateral resolution within the angle of view formed by lines connecting the focal point of the arc with each edge transducer in the array. Arc-shaped probes are often most effective for scanning curved body surfaces (such as average-sized chest, neck, arms, and legs) with a radius that roughly matches that of the probe.

FIG. 1 (a) then enlarges the optical diffusers LD passing through the optical windows OW and merges the two optical beams OB from the fiber bundles FB passing through them into the tissue. Illustrates one embodiment of an optoacoustic probe that provides illumination of the tissue TS through the skin SK by the scattered light SL beam formed. The sound waves AW generated in the blood vessels or tumors (BV or TM) by the light scattered from the tissue propagate through the acoustic lens AL to the transducers TR and the electrical cables EC ) To electrical signals transmitted through the support material BM to the electronic amplifiers.

In one embodiment, the design of the optical fiber bundle is as follows. The input of the fiber bundle is circular with fused fiber tips to avoid loss of light through the spaces between the fibers. The fiber diameter can be approximately 200 microns for good flexibility, and a fiber diameter of 100 microns or 50 microns may be desirable in certain applications. This fiber bundle is Y-separated into two half-bundles and is completely randomized, allowing substantially any two neighboring fibers from the input to appear in different half-bundles. At least a number of neighboring fibers must be randomized in this regard. Each half-bundle is preferably separated into a number of sub-bundles, each sub-bundle being positioned in its slot / niche to form a fiber bundle ("paddles"). Two paddles are located on each side of the ultrasonic transducer (TR) array assembly. As discussed below with reference to Figures 7B and 7C, the output shape of each fiber bundle paddle can be rectangular for the width of the field of view, typically 40 mm, and has triangular ends. This triangular shape allows the output beam to have smooth edges after passing through the light diffuser (LD) (Fig. 1 (a)). Finally, the optical beam from the fiber bundle paddle is an optical window (OW) comprising thin anti-reflection-coated glass plates, anti-reflection-coated polymer or plastic plates with acoustic impedance matching those of the tissues to be imaged. ) Through the probe to the skin (SK).

There are a number of objectives for current optoacoustic probe design: (i) virtually no light must propagate through the acoustic lens (AL) or through the light blocking acoustic damper (OBAD) on the sides of the probe, and (ii) ) Practically no sound waves should be generated in the acoustic lens or light blocking acoustic damper materials through the imprint of light; Sound waves at a wide range of ultrasonic frequencies from 0.1 MHz to 15 MHz must be able to pass (AL) without any attenuation, and any sound waves must be able to pass through the OBAD; (iii) The optical beams OB exiting through the optical windows OW must have smooth edges of the light fluence, these optical beams being maximally in the image plane and for merging due to light scattering within the skin. Under the arrays or transducers that provide strength, you should enter the skin as close to each other as necessary to enter the underlying tissue.

In one embodiment, the light delivery system delivers light underneath the transducer elements, not through an array of transducer elements. In one embodiment, the design of the optoacoustic probe is an optical fiber on each side of the ultrasound array, positioned as close to the transducers as possible, given the need to focus ultrasound beams to the depth of the most likely target. It has transmission systems and is based on an array of ultrasonic transducers with dimensions on the height axis of the transducer as small as possible. In one embodiment, the optical fiber delivery system is designed to allow penetration of light energy of near infrared light into an imaged organ, such as the chest, and minimal photo-thermal-mechanical interaction of the light beam with the skin.

Another alternative design of the light transmission system delivers light to a mirror or prism (s) located below the ultrasonic transducers to reflect light at right angles to the skin surface of the organ being imaged. In these embodiments, a standoff can be placed between the transducer elements and the skin / tissue. These alternative embodiments can be combined within the scope of the present invention.

Detailed description of aspects of system components

Optical lighting and Probe  design

The acoustic lens is positioned on the transducers in the optoacoustic probe to focus the ultrasound beams. The probe may not be provided without an acoustic lens, but without the lens, the ultrasonic transducers are directly exposed to light and can absorb this light, which in particular can cause very large artifact ultrasound signals to be pulsed. have. Optical illumination of the lens on the ultrasonic probe results in very strong transients that cause image artifacts. Depending on the skin color, up to 50% of near infrared light can be scattered widely by the skin. Mismatch of the acoustic impedance between the lenses of the transducer elements can cause reflections with a long ring down time. Therefore, one embodiment of the probe design includes a white strongly scattering opaque lens. If such a lens is not required due to the curved shape of each transducer element, a white strongly scattering front matching layer should be used to protect the transducer elements from near-infrared light.

FIG. 1B illustrates how the laser illumination light 110, 120 from the photoacoustic probe can be scattered 130 from the skin 410 to the probe's acoustic lens 150.

Moreover, (laser) optical pulses have a direct effect on the ultrasonic transducers of sound waves introduced by the strong interaction of optical pulses with the skin of the imaged organ traversing laterally along the skin surface in a direction orthogonal to the image plate. Can give. When detected by an array of transducers, the spatial distributions of these sound waves are projected into an optoacoustic image at a depth equal to the lateral distance between the array of transducers and the optical beams on the skin surface, creating artifacts. Moreover, the acoustic waves generated through the acoustic lens reflection and the housing of the probe may further affect the quality of the imaging.

2A and 2B illustrate representative optoacoustic signals showing the effect of lateral ultrasound waves introduced by laser pulses on the skin using optical beams on each side of the ultrasound transducer array. The illustrated signals are generated by the transducers in a direction almost perpendicular to the plane of the images generated therefrom. These transducers are capable of receiving signals with a large square (up to 90 degrees) relative to the plane of the images generated therefrom, which is undesirable. Therefore, the design of the transducer array includes means for rejecting signals coming from outside the image plane. Such means include, but are not limited to, concave arc-shaped transducer elements and acoustic lenses and transmission of optical beams under the transducers. The detected photoacoustic signals 210 in FIG. 2A were generated using an effective acoustic binder, in this case water. The signals 220 in FIG. 2B were generated in the absence of such an acoustic binder, ie using only the airspace to couple the acoustic signals to the transducer array.

Moreover, the limited dimensions of the optical beam can affect the sound waves generated in response to the impact of the optical beam on the tissue. These sound waves are generated at the sharp edges of the optical beam, propagate towards the array of transducers, and can cause artifacts. Systems using a flat linear array of ultrasonic transducers are configured so that the first and last transducers in the array are configured to detect these waves, so the first and central transducers detect these waves last, and this “edge effect”. Causes v-type artifacts on the sinogram of optoacoustic signals and low frequency bulk artifacts on optoacoustic images.

3 illustrates an example of the display of v-type artifacts 310 on the sinogram 300 of acoustic signals and associated artifacts 320 on acoustic images. In one embodiment, since these sound waves are associated with the edge effect of the optical illumination beams with abrupt changes in the light fluence, one has the V-shaped bright signals on the sinogram and the related series of artifact waves on the optoacoustic image. can see.

Moreover, the illumination geometry of the optical beams projected by the optoacoustic probe can affect image quality. The optical beams of the optoacoustic probe are located too far apart, which is the bright field of illumination (two separate optical beams on each side of the probe that cause the absence of light directly in the probe term in the image plane). (One beam below the probe entering the depth of the tissue along the image plane) can cause a gradual transition. This transition creates a problem with the image brightness map that makes the map quantitatively inaccurate and causes artifacts to the same depth as the initial width between separate optical illumination beams on each side of the probe.

4A-4C are hand-held optoacoustic props 410, 420, 430 that transmit light energy on the side of the probe or from below the optoacoustic probe at different distances where the optical illumination of the tissue is different. ) To illustrate one embodiment achieved. In the embodiment of Fig. 4A, when the optical beams are transmitted under the ultrasonic probe, the distribution of light energy in the image plane has a maximum at the skin surface and has a smooth slope. This optical distribution is beneficial for high contrast of optoacoustic images. In the embodiment of Fig. 4B, when the optical beam is delivered close to the thin optoacoustic probe, the two beams can merge due to optical scattering in the skin, so that the light energy in the tissue under the skin The distribution can be made similar to the embodiment of Fig. 4A. In the embodiment of Fig. 4 (C), when the optical beams are separated by a large distance, they only merge to a considerable depth in the tissue, and the dark region (no light) and tissue of the tissue's substrate layer Contrast of acoustic images, taking into account the projection of the brightly illuminated areas of the skin to the optoacoustic image plane at a depth equal to the separation distance of the two beams, with a light zone at depth and creating an optical distribution in the image plane. Is harmful to.

Thus, in the embodiments illustrated in FIGS. 4A-4C, the image brightness maps 412, 422, and 432 of the tissue to be scanned where the illumination of the skin is directly below the probe 410. Is optimized. As shown at 420 and 430, as the distance between the center of the transducers and the center of the optical beams increases, the image brightness maps 422 and 432 of the tissue being scanned continue to become more non-uniform.

Finally, the reflection from the boundaries of the tissue structures (tumor, blood vessels or tissue layers) of laser-guided ultrasound waves produced by the skin and then started into the tissue is also an image expressed by lines, curves, and total noise. It can lead to artifacts.

In one embodiment, the probe's acoustic lens is designed such that the lens reflects and scatters light from the illumination components, but does not absorb, which is optionally opaque. In various embodiments, such lenses use thin, selectively scattering materials, such as titanium dioxide or barium sulfate powder-filled silicone rubber, or thin metallic, such as a combination of aluminum or gold or white opaque lens materials and metal layers. It can be made using a highly reflective layer. In one embodiment, in the case of a combination of a widely scattering material of a lens and a thin reflective layer (foil), to avoid peeling of the thin metallic layer from the front surface of the acoustic lens, the metallic reflective layer is two layers of a widely scattering material Can be located between. It is difficult to make a material with absolutely zero light absorption, and since this absorption can generate ultrasonic waves from thermoelastic materials, the lens material can be made from thermoplastics with minimal thermal expansion, which is absorbed. Generates minimal ultrasound or no ultrasound in response to light energy.

5A and 5B illustrate two embodiments of hand-held optoacoustic ultrasonic probes 510 and 520, each protected from optical illumination of an acoustic lens. In FIG. 5A, an entirely opaque white lens is used, and in FIG. 5B, a partially reflective white lens is used, and the light reflecting ability of the lens is enhanced by a gold layer or coating.

6 illustrates optoacoustic images using a probe with a non-reflective acoustic lens 610 and a probe with a reflective layer of gold 620. The probe using the reflective layer of gold 620 produces an image with reduced artifacts 612 and 614.

In one embodiment, the probe housing acts as a low-reflection encapsulation of the probe, which means that if the probe housing does not (i) absorb laser light (more specifically near-infrared light), but small absorption cannot be avoided, low heat This means that materials with expansion do not emit ultrasonic waves after absorption of laser light, and (ii) are made from materials that strongly attenuate, attenuate, and do not echo ultrasonic waves. The transducer assembly within the probe housing is also made of low-reflective material. Alternatively, the layer of low-reflective material is positioned between the transducer assembly and the fiber optic assembly to avoid generating any ultrasound upon interaction of light in the transducer assembly. In various embodiments, these materials are, for example, baffles, foams, polymers, rubbers, and plastics (CR-600 resins available from Micro-Mark, Berkeley Heights, NJ, or Virginia, White color porosity and anechoic heterogeneous composites (such as AM-37 available from Syntech Materials of Springfield). In one embodiment, any of these materials are electrically sub-conducting insulators to protect the probe from inter alia external electromagnetic emissions.

In one embodiment, the optical illumination subsystem is configured to deliver optical beams with smooth intensity edges. In one embodiment, the width of the optical beams is the same as that of the array of ultrasonic transducers in the acoustic probe (eg, about 5 mm). This is achieved by designing a bundle of optical fibers to have a density of fibers that slowly decreases at the edges. This design allows the beam to deliver laser illumination to the skin of the organ under investigation so that it does not produce sharp edge-related sound waves, which do not generate V-shaped artifacts in the sinogram of optoacoustic images. .

7A illustrates an embodiment of an optical beam 710 with sharp edges that can produce edge effects of sound waves and related artifacts and an optical beam 720 with smooth edges that produce reduced edge-related artifacts. For example. 7B and 7C illustrate that an optical beam with the strength of smooth edges can be generated using a fiber optic bundle design with multiple sub-bundles in a triangular shape at each end of the fiber bundle assembly.

In one embodiment, the fiber bundle is positioned at a distance from the skin sufficient for the optical beam to enlarge to the desired width. If the dimensions of the probe are small, the fibers used in the fiber bundle can be selected to have a higher number of openings (eg> 0.22). In one embodiment, the beam is delivered through an optical window to achieve better coupling of the optical beam to the skin. In this embodiment, the optical window touches the skin, flattening its surface for better light transmission, while simultaneously removing any excess of binding gel from the imaged skin surface. In one embodiment, the fiber bundle and optical window are integrated into the probe housing, so that the voids between the fiber bundle and window are protected.

In one embodiment, the optical window is designed to allow minimal interactions of both the optical beam and laser-guided sound waves with this window. In one embodiment, the window is made of an optically transparent material with a very thin and antireflective (AR) optical coating. In one embodiment, this material has anechoic acoustic properties. These anechoic acoustic properties and the fact that the illuminated skin is depressed during photoacoustic scanning results in attenuation of ultrasound waves laterally propagating from the laser-illuminated skin surface to the transducer array, thereby reducing associated artifacts.

In one embodiment, the probe is designed such that the optical beams are very close to the transducer elements on each side of the ultrasonic probe, which is made as technically as thin as possible. In one embodiment, the thickness of the probe is very small (eg, 5 mm) so that the light beams delivered to the skin at this distance d from the center of the probe are one beam within the thickness of the skin (about z = 5 mm). The organization of the organ being investigated will receive a beam under the transducer elements. This is called bright field illumination. In one embodiment, the optoacoustic probe is designed such that the optical light beam is delivered to the skin just below the transducer elements.

FIG. 8 shows two probes 810 and 820 with two fiber bundles on each side of each of the beams oriented to illuminate the skin directly under the probe 812 and on one side of the probe 822. Illustrates the effect of optical illumination on. When the skin is illuminated just below the probe 812, the tumor 814 is clearly identifiable and there is no clutter on the image background 816. If the skin is illuminated on either side of the probe 822, the tumor is not identifiable (824) and there are multiple artifacts on the image background (826).

In one embodiment, the optical beam width is designed to deliver increased light to the flakes of the tissue being imaged. In one embodiment, the beam is homogeneous, so it has a constant intensity through the beam, since the heterogeneous beam creates heterogeneous acoustic sources, which in turn creates artifacts in optoacoustic images. The intensity level is defined by ANSI laser safety standards for laser illumination of the skin. The beam width is limited by the ability of optical scattering in the tissue to deliver photons of light to a central lamella (imaged lamella) located under the transducer elements. In one embodiment, the length of the optical beam is equal to the length of the transducer array. In one embodiment, because the sharp edges produce strong edge artifacts on optoacoustic images, the optical beam also has smooth edges, that is, a gradually decreasing intensity at the edges.

In one embodiment, the design features of the optical illumination system and acoustic probe of the present disclosure can be summarized in the following table.

System features Advantages Arc hand-held probe Higher hole-lower distortions Light transmission to the imaging plane Improves optoacoustic image contrast and reduces artifacts by increasing the ratio of useful information (from the imaging plane) to noise (outside the imaging plane) Optical shielding of the probe Reduces artifacts from direct and scattered light hitting the acoustic lens, probe housing, etc. Acoustic shielding of the probe Acoustic shielding of the probe's housing reduces artifacts (clutters) from sound waves propagating through the probe's housing Using ultra-wideband transducers for both ultrasound and optoacoustic imaging Allows to have the same array operating in ultrasound and optoacoustic imaging

In various embodiments, the shape of the array of ultrasonic transducers for combined optoacoustic / ultrasound imaging may be flat or convex arc-shaped. In one embodiment, the probe shape for optoacoustic imaging is concave arc-shaped. This concave shape provides a large aperture with minimal physical dimensions, a wider field of view of the imaged object, which in turn provides better reconstruction of the shape of the imaged object and improved lateral resolution.

9A-9C are implementations of optoacoustic / ultrasound hand-held probes with flat or concave arc shapes 910 (FIG. 9A) and hand-held stiff probes with linear shape 920 (FIG. 9B). Illustrate examples. 9C is an optoacoustic / ultrasound hand with a side showing the assembly of the ultrasonic transducers, two layers of low-reflection light reflection and ultrasonic damping material on each side, and two optical windows for the transmission of the optical beam. -Illustrate the details of the Held Probe design.

9C illustrates details of a hand-held optoacoustic probe with a concave arc shape. Electrical cables 930 are provided for two-way communication to and from the probe, and fiber optic bundles 940 are provided to transmit light to the probe. The array of light-band ultrasonic transducers 950 transmits and receives acoustic energy. Transducer array 950 is covered by an opaque white cylindrical lens (not shown for clarity) that broadly scatters and reflects near-infrared light. Optical window 960 provides optical beam outputs. In the embodiment of Fig. 9c, the ultrasonic transducers in the probe can be designed to be insensitive to lateral acoustic (ultrasound) waves and free of reflections, especially in the lateral direction. This can be achieved by the choice of piezoelectric composite material, the shape of the piezoelectric ceramic elements in the matrix and the anechoic properties of the matrix. In one embodiment, ultrasonic transducers are also designed to have high sensitivity within ultra-wideband ultrasonic frequencies. This in turn results in minimal reflections that cause artifacts for optoacoustic / ultrasonic images.

9D shows an optoacoustic image illustrating the advantages of a concave-arc hand-held probe for resolution in optoacoustic images. As provided in this embodiment, the large spherical shape and sharp edges are well depicted in cases where the object is within the field of view of the probe opening. The resolution and accuracy of shape reproduction is reduced out of the probe opening, but remains better than for flat linear probes of similar width.

9E illustrates an alternative embodiment of an optoacoustic / ultrasound hand-held probe design that also enables three-dimensional images, and three-dimensional images in a plane parallel to the skin surface at various selected depths.

In one embodiment, a hand-held probe scanned along the skin surface that produces real-time two-dimensional images of tissues in the body under investigation also has a component that acts for accurate global 3D positioning of the probe. . This design allows the imaging system to remember the locations of all tissue slices and reconstruct the three-dimensional images at the end of the scanning procedure.

Electronic data acquisition system

In one embodiment, the present disclosure relates to an optoacoustic imaging system having an electronic data acquisition system that operates in both optoacoustic and ultrasonic modes and is capable of quickly switching between these modes. In one embodiment, this is achieved with a field programmable gate array (FPGA) on the electronic data acquisition system, firmware controlling the functions of the main microprocessor. In one embodiment, the reprogrammable FPGA can toggle between optoacoustic and ultrasound operational modes in real time, thus enabling cross-matching of ultrasound and optoacoustic images, which is based on functional and anatomical maps. It can be used for diagnostic imaging. In one embodiment, FPGA functions include controlling, acquiring, and storing photoacoustic and / or ultrasound data, signal processing, and passing data for real-time image reconstruction and processing. In one embodiment, the FPGA can also be used for ultrasound beam formation and image reconstruction.

In one embodiment, the electronic data acquisition system design uses one or more multi-core graphical processor units (GPUs) for image reconstruction and processing. In ultrasonic mode, in one embodiment, the FPGA controls ultrasonic transmission, which performs both ultrasonic and optoacoustic data acquisitions on a multi-channel board. To enhance the operation of the FPGA's memory, an external memory buffer can be used. In one embodiment, the FPGA allows for rapid reprogramming of ultrasound data acquisition with a signal / frame repetition rate of 2-20 kHz for photoacoustic data acquisition with a signal / frame repetition rate of about 10-20 Hz, ultrasound The structure of the gates and the internal memory structure and size are also configured to allow real-time switching between emission and detection, laser synchronization, and system controls. In one embodiment, multiple FPGAs can be used to enhance system performance. In one embodiment, in ultrasonic and optoacoustic modes, the FPGA clock can be changed by appropriate time-division multiplexing (TDM). In one embodiment, the design of the multi-channel electronic data acquisition system may be based on modules, and the module is typically 16 to 128 channels, but more than 256 channels may be suitable in some applications. In one embodiment, the design of a multi-channel electronic data acquisition system has 64 channels.

In order to achieve the dual way operation of the optoacoustic / ultrasonic system, a separate optoacoustic electronic system can also be combined with a separate ultrasonic electronic system via a single probe. In one embodiment, the probe has a cable with Y-separation to connect the probe to optoacoustic and ultrasonic electronic systems. In one embodiment, the programmable electronic switch is capable of converting a signal detected from a probe (transducer array) into optoacoustic electronic devices (to operate in optoacoustic mode) or to ultrasonic electronic devices and from ultrasonic electronic devices. Allows transmission to the probe (to operate in ultrasonic mode). In one embodiment, the synchronization trigger signal is sequentially transmitted to optoacoustic and ultrasound systems, such that optoacoustic and ultrasound images are alternately acquired.

Processing, reconstruction and display of images

Signal processing

In various embodiments, the purpose of the diagnostic imaging procedure is to display each pixel with brightness that accurately replicates the signals originally generated in each voxel of tissue displayed on the image. On the other hand, the intrinsic pressure signals generated by optical pulses within the tissues will be significantly altered in the process of propagation through the tissue, and especially in the process of detection and recording by ultrasonic transducers and electronic subsystems. Can be.

In one embodiment, the detected signals are processed to reverse changes and restore the original signals. In one embodiment, this reversal may be achieved through deconvolution of the system's impulse response (IR). In one embodiment, the impulse response can be measured by recording and digitizing the delta-function ultrasonic signal generated by short (nanosecond) laser pulses in a strongly absorbing optical medium with a high thermoelastic expansion coefficient.

One component of the impulse response is an acoustic-electrical impulse response that provides for optoacoustic or ultrasonic signal distortions due to the characteristics of ultrasonic transducers, cables and analog electronic devices. The second part of the impulse response is a spatial impulse response that provides for signal distortions associated with limited dimensions of ultrasonic transducers. In various embodiments, large transducers can incorporate ultrasonic waves incident at one angle, while point-source-type transducers can provide a perfect or near-delta-function spatial impulse response.

In one embodiment, any distortions in the acoustic-electric impulse response can be inverted from signals detected by impulse response deconvolution. However, possible distortions in the spatial impulse response can be avoided by designing transducers with small dimensions within the image plane. In one embodiment, the dimensions of the transducers are much smaller than the shortest wavelength of ultrasound waves that can be detected or emitted by the transducers.

10A-10C show the impulse response of an ultrasonic transducer with a relatively narrow sensitivity of 1010 as a function of frequency for the ultra-wide and narrow-band resonant transducers 1030, of the ultra-wide-band ultrasonic transducer 1020. Examples of impulse response, and ultrasonic spectra of transducer sensitivity are shown.

In one embodiment, the first step in processing the optoacoustic signal in an imaging system that produces two-dimensional optoacoustic images is deconvolution of the acoustic-electro-impulse response.

11A and 11B provide exemplary examples of deconvolution of the impulse response of transducers from detected optoacoustic signals 1110, where deconvolution is the original, unchanged, N-type pressure The signals 1120 are restored.

In one embodiment, the second step in processing the optoacoustic signal is signal filtering to remove noise using a signal filter. In one embodiment, the signal filter is based on wavelet transform operating simultaneously in the frequency and time domain. In one embodiment, this wavelet filter belongs to noise and can filter certain frequency components of a signal appearing at a given time, while preserving similar frequency components of useful signals appearing at different times. In one embodiment, the frequency spectrum of the wavelet filter replicates the frequency band of a conventional N-type optoacoustic signal while simultaneously providing smooth window edges that do not cause signal distortions during convolution.

In one embodiment, such a wavelet filter is useful for optoacoustic imaging in its ability to restore the original pressure profile generated in the tissue prior to pressure propagation. During propagation through tissue, the original positive pressure signal translates into a bipolar (compression / tensile) profile. Therefore, reconstruction of the image of absorbed light energy (photoacoustic image) requires a transformation starting with bipolar signals and provides for all-positive values for optoacoustic image intensities. In one embodiment, a multi-scale wavelet filter, e.g., a filter that simultaneously integrates a signal over time and provides a summation of the multiple frequency bands present in the signal, provides a positive or negative amount of positive pressure signals. It can be converted into a monopole signal expressing pressure.

12A-12C show their original rectangular pressure by summing all scales corresponding to low to blast furnace frequency ranges for 5 scales 1210, 7 scales 1220, and 9 scales 1230. Provides an exemplary example of wavelet filtered N-type optoacoustic signals reconstructed with a profile.

In various embodiments, wavelet filtering allows for enhancement of objects on the image within a certain range of dimensions. Imaging operators (ultrasound technologists or diagnostic radiologists) typically want to better visualize tumors with specific dimensions and other objects, such as blood vessels, with their specific dimensions. In one embodiment, the wavelet filter allows the operator to apply only a particular selection of scales of the wavelet filter to suppress objects of other non-critical sizes while enhancing only objects of specific sizes. In one embodiment, boundaries can be well visualized for objects of any size, so high-frequency wavelet scales are beneficial for image quality and are included in the selection of scales. In one embodiment, for a mathematically accurate tomography reconstruction, a ramp filter can be applied to the signal, which can linearly enhance the contribution of higher frequencies.

Image reconstruction

In various embodiments, image reconstruction typically uses radial backprojection of signals processed and filtered to the image plane. However, due to the limited field of view available from small hand-held probes, only an incomplete data set can be obtained. As a result, 2D optoacoustic images may include artifacts that distort the shape and brightness of objects displayed on the images. In one embodiment, aperture integrated normalized radial back transmission is used to correct some of the reconstruction artifacts observed in limited aperture optoacoustic tomography.

13 provides an exemplary diagram of radial back-transmission where each transducer element opening is weighted and normalized with respect to the total opening of the transducer array.

(모든 ω_i,j,k의 합)은 그것이 전체 트랜스듀서 어레이에 의해 가시화되는 바와 같이 픽셀(i,j)에 의해 방출된 광음향 파면의 부분이며, S_{i ,j,k}는 제 k^th 트랜스듀서에 의해 측정되고 픽셀(i,j)에서의 밝기의 재구성에 사용된 광음향 신호의 샘플이다. 다양한 역전파 알고리즘들이 광음향 이미지를 정규화하는데 사용될 수 있다.In one embodiment, T _k , -T _{k +4} (1311 to 1315) are transducers 1310 in the array, and B _{i , j} is the brightness (intensity) of the pixel with coordinates (i, j). , Ω _{i, j, k} (1320, 1330) is the angular portion of the photoacoustic wavefront emitted by the pixel i, j as it is visualized by the transducer #k,

(Sum of all ω _{i, j, k} ) is the portion of the optoacoustic wavefront emitted by the pixel (i, j) as it is visualized by the entire transducer array, S _{i , j, k} is the k ^{th th} It is a sample of an optoacoustic signal measured by a transducer and used to reconstruct the brightness at pixels (i, j). Various backpropagation algorithms can be used to normalize photoacoustic images.

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

(1)

(One)

However, in at least some embodiments, aperture normalization backprojection produces good image results. In one embodiment, aperture normalization backprojection can be expressed as follows:

(2)

(2)

14A and 14B provide exemplary examples of optoacoustic tomography images 1410 and 1420 of an imaging lamella through a tumor angiogenesis model. In the first image 1410, a backpropagation algorithm, such as the first algorithm directly above, is used to normalize the image. The resulting image has strong, bright arc-like artifacts 1412 around blood vessels 1414 close to the array surface. In the second image 1420, an aperture normalization inverse projection algorithm, such as the second algorithm directly above, is used to normalize the image. As can be seen, the aperture normalization back-projection algorithm corrects image brightness and reduces arc-like artifacts.

15A and 15B show point spread as visualized with a flat linear probe using a back propagation algorithm 1510, such as the first algorithm directly above, and an aperture normalized back projection algorithm, 1520, like the second algorithm directly above. Provides illustrative examples of photoacoustic tomography images 1510, 1520 of the function. As can be seen, the aperture normalization back-projection algorithm corrects image brightness and reduces artifacts.

Image processing and display

In one embodiment, the optoacoustic image palette is equalized to reduce the effects of light distribution within the tissue. This equalization transforms the dynamic range of optoacoustic images for better visualization of both shallow and deep objects.

16A and 16B are embedded in different depths where the first image 1610 is created using one embodiment of a standard palette and the second image 1620 is created using one embodiment of a depth-normalized palette. Provides illustrative examples of photoacoustic images 1610, 1620 of phantoms with furs. As can be seen, using the depth-normalization palette enhances the visibility of deep objects in the illustrated embodiment.

In one embodiment, principal component analysis (PCA) for single optoacoustic image acquisition (different channels) is used to remove cross-correlated signal noise. Principal component analysis on a dataset of optoacoustic signals can remove correlated image clutter. Principal component analysis for optoacoustic frames can also remove correlated image clutter.

17A and 17B provide exemplary examples of optoacoustic images 1710 and 1720 of a phantom of a spherical simulated tumor obtained with a flat linear probe. The first image 1710 is an original image not subject to principal component analysis processing. The second image 1720 has been subjected to principal component analysis processing together with the first principal component deconvolution. As can be seen, using principal component analysis processing enhances image quality, among other things, by reducing artifacts.

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

System features Advantages Operator-assisted boundary tracking for ultrasound and optoacoustic images Quantitative optoacoustic diagnoses can be improved by evaluating diagnostic parameters within the tumor boundaries defined for US images.
Diagnosis can be reinforced by morphological analysis of tumor boundaries. Aperture integrated normalized radial back projection Corrects some of the reconstruction artifacts observed in limited aperture optoacoustic tomography. Equalization of the optoacoustic image palette to reduce the effects of light distribution within the tissue Deforms the dynamic range of optoacoustic images for better visualization of both shallow and deep objects Principal component analysis of photoacoustic signal data (PCA) PCA for a single optoacoustic acquisition (different channels) is a fast and efficient way to remove cross-correlated signal noise.
PCA for a dataset of optoacoustic signals removes correlated image clutter.
PCA for optoacoustic frames removes correlated image clutter. Photoacoustic imaging system with quantitative evaluation of total hemoglobin, blood oxygenation, and water Cancer diagnosis based on these parameters for the mean background or single malignancy index (tHb ^* water / oxygenation) Wavelet transform that enhances images of objects within a specific dimension range -The operator can easily select the maximum size of objects to be enhanced on the image. Everything greater will be filtered out. Adaptive beam-forming for optoacoustic imaging -Allows individual reconstruction of groups of radial wavelet sub-bands

Diagnostic image reprocessing

The principles of functional diagnostic imaging can be based on tumor pathophysiology. For example, malignant tumors have enhanced concentrations of total hemoglobin and reduced levels of oxygen saturation in hemoglobin in the blood. In one embodiment, the optoacoustic images can be reprocessed, among which are converted into images of (i) total hemoglobin ([tHb]) and (ii) hemoglobin oxygen saturation ([SO2]). 18 shows an example of two breast tumors.

FIG. 18 shows a diagram illustrating tumor differentiation based on absorption coefficients at two wavelengths (755 nm (1810) and 1064 nm (1820), which is absorbed by hemoglobin (hypoxic blood) versus oxygen hemoglobin) The local maximum (757 nm) and minimum (1064 nm) ratios of absorption ratios are matched, as can be seen, while malignant tumor 1830 has absorption coefficient at 757 nm higher than benign tumor 1840. , Benign tumor 1840 has an absorption coefficient at 1064 nm higher than that of malignant tumor 1830.

19 illustrates tumor discrimination by photoacoustic images based on absorption coefficients at two wavelengths 1910 and 1920 in the phantom. At 757 nm, the model of malignant tumor (1920) is clearly visible (1922), while the model of malignant tumor (1922) is not visible at 1064 nm (1910).

20A shows an optoacoustic image of two crossed tubes filled with blood with different levels of blood ([SO2]) (98% in the left tube and 31% in the right tube). The tubes are placed in a 1% milk fat with optical properties similar to those found in the human chest. The wavelength of the laser illumination used for this image is 1064 nm. 20B shows a photograph of an experimental setup containing artificial blood vessels positioned in a milk solution and imaged using an arc-type optoacoustic probe. 20C is an ultrasound image depicting the anatomy of a body with six image panels, namely (1-top left) blood vessels; (2-top right) Photoacoustic images obtained at a wavelength of 757 nm; (3-bottom right) photoacoustic images obtained at a wavelength of 1064 nm; (4-lower left) functional image of total hemoglobin ([tHb]); (5-lower center) Functional image of blood oxygen saturation ([SO2]); (6-upper center) Only cross-matched 2D cross-section anatomical and functional images of vascular tubes showing functional images of blood oxygen saturation provided in the region of maximum concentration of total hemoglobin. The original optoacoustic images depicted in FIG. 20C in the upper right and lower right panels show different brightness of blood vessels with different levels of total hemoglobin concentration ([tHb]) and blood oxygen saturation ([SO2]). , Accurate quantitative measurements can be performed under conditions of normalized intensity of optical illumination of tissue in the body as a function of depth. These optoacoustic images are used to reconstruct functional images of total hemoglobin ([tHb]) and blood oxygenation ([SO2]). All functional images displayed in FIG. 20C are mutually matched and superimposed with the anatomical image of the tissue structure for better correlation of features.

21A and 21B show photoacoustic signal amplitude as a function of blood oxygen saturation (with constant hematocrit) under laser illumination at wavelengths of 1064 nm in FIG. 21A and 757 nm in FIG. 21B. These plots illustrate that blood oxygen saturation can be monitored with optoacoustic imaging. Specifically, this example was used for measurements of photoacoustic signal amplitude in blood with various levels (30% to 98%) of hematocrit and oxygen saturation of hemoglobin ([tHb]) of 38 g / dL in red blood cells. Illustrate based quantitative data. As predicted by the published absorption spectra of blood, the photoacoustic signal amplitude at 1064 nm illumination increases with increased level of oxygen saturation, while the photoacoustic signal amplitude increases blood oxygenation at 757 nm illumination wavelength. Decreases with.

22 illustrates light absorption spectra of the main tissue chromophores that absorb light energy in the near-infrared range: hemoglobin, oxygen hemoglobin and water. Preferred laser wavelengths for functional imaging are 757 nm and 1064 nm matching the maximum and minimum ratios of [HHb] / [O2Hb], while the wavelength of 800 nm serves for calibration purposes through measurements of total hemoglobin ([tHb]). It is best for.

23A and 23B illustrate cross-matched functional and anatomical imaging of breast tumors in phantoms that accurately replicate the optical and acoustic properties of the average chest with tumors. 23A shows a 2D image of the model of malignant tumor morphology based on ultrasound (left), where the same anatomical image is cross-matched with a functional image (center) of total hemoglobin concentration and a functional image of blood oxygenation (right). FIG. 23B shows 2D images of model positive tumor: morphology image (left) based on ultrasound, and the same anatomical image is mutually matched with functional image (center) and blood oxygenation (right) of total hemoglobin concentration.

24A and 24B illustrate cross-matched functional and anatomical imaging of breast tumors. 24A shows 2D images of invasive breast cancer, malignant tumors with rough borders, heterogeneous morphology, high concentrations of total hemoglobin and low oxygen saturation (hypoxia). Malignant tumor morphology is based on ultrasound in the left image, and the same anatomical image is mutually matched with the functional image of blood oxygenation in the central image and the functional image of total hemoglobin concentration in the right image. 24B shows 2D images of the breast with fibroadenoma, benign tumors with relatively rounded borders, moderate concentrations of oxygen hemoglobin and relatively low total hemoglobin. Chest morphology is based on ultrasound in the left image, and the same anatomical image is mutually matched with the functional image of blood oxygenation in the central image and the functional image of total hemoglobin concentration in the right image.

conclusion

While some embodiments may be implemented in fully functioning computers and computer systems, various embodiments may be distributed as computing products in various forms, and in particular certain types of machines or computers used to effect distribution. It can be applied regardless of readable media.

At least some aspects disclosed may be embodied at least partially in software. That is, the techniques described herein are special purpose in response to its processor, such as a microprocessor, that executes sequences of instructions contained in memory, such as ROM, volatile RAM, non-volatile memory, cache or remote storage device. Or it may be implemented in a general purpose computer system or other data processing system.

Routines executed to implement the embodiments may include operating system, firmware, ROM, middleware, service delivery platform, software development kit (SDK) component, web services, or other specific application, component, program, object, module, or It can be implemented as part of a sequence of instructions called "computer programs". Calling interfaces for these routines can be exposed to the software development community as an API (Application Programming Interface). Computer programs are typically set at various times in various memory and storage devices in a computer, and when read and executed by one or more processors in the computer, the computer needs to execute elements that involve various aspects. It contains one or more instructions to perform them.

Machine-readable media may be used to store software and data that, when executed by a data processing system, causes the system to perform various methods. The executable software and data can be stored in various places including, for example, ROM, volatile RAM, non-volatile memory and / or cache. Portions of such software and / or data may be stored on any of these storage devices. In addition, data and instructions can be obtained from centralized servers or peer-to-peer networks. Different parts of the data and instructions can be obtained at different times and at different communication sessions, or from different centralized servers and / or peer-to-peer networks in the same communication session. Data and instructions can be obtained as a whole prior to the execution of the applications. Alternatively, portions of data and instructions can be obtained dynamically, on time, when required for execution. Thus, data and instructions are not required to be on a machine-readable medium as a whole at a particular time instance.

Examples of computer-readable media are, but are not limited to, writable and non-writable tangible media such as volatile and non-volatile memory devices, read only memory (ROM), random access memory (RAM), among others. ), Flash memory devices, floppy and other removable disks, magnetic disk storage media, optical storage media (eg, compact disk read-only memory (CD ROMS), digital versatile disks (DVDs), etc.) do.

Generally, a machine-readable medium provides (stores) information in a form accessible by a machine (eg, a computer, a network device, a personal digital assistant, a manufacturing tool, any device having a set of one or more processors, etc.). Includes any mechanism.

In various embodiments, a hardwired circuit can be used in combination with software instructions to implement techniques. Thus, the techniques are not limited to any particular combination of hardware circuitry and software to any particular source for instructions executed by the data processing system.

Some of the figures illustrate multiple operations in a particular order, but operations that are not sequence dependent can be reordered and other operations can be combined or disassembled. Some reordering or other groupings are specifically mentioned, but others will be apparent to those skilled in the art and do not provide an exhaustive list of alternatives. In addition, it should be appreciated that the steps can be implemented in hardware, firmware, software, or any combination thereof.

In the foregoing specification, the disclosure has been described with reference to certain representative embodiments thereof. It will be apparent that various changes can be made to it without departing from its broader spirit and scope, as set forth in the following claims. Accordingly, the specification and drawings are to be regarded as illustrative rather than restrictive.

Claims (50)

In the imaging system for the visualization of flakes to the depth of the tissue of all or part of the body,
A hand-held imaging probe including a light emitting portion and an array of ultrasonic transducers;
In the light emitting portion, the first optical fiber bundle and the first optical diffuser are configured to form a first optical beam, and the second optical fiber bundle and the second optical diffuser are configured to form a second optical beam, and the first The optical fiber bundle and the second optical fiber bundle are disposed on opposite sides of the array of ultrasonic transducers,
A processing system that receives data originating from the hand-held imaging probe and processes three or more independent images based on all or part of the data, wherein the three or more independent images are:
A first functional image reflecting the distribution of total hemoglobin concentration;
A second functional image reflecting the distribution of blood oxygen saturation; And
Comprising the processing system with a morphological image of tissue structures,
The processing includes deconvolution of an impulse response associated with the array of ultrasonic transducers, the impulse response being based on a delta-function ultrasound signal generated by laser pulses in an absorbing optical medium. According to claim 1,
And a signal filter for removing noise of the delta-function ultrasound signal and in communication with the processing system. According to claim 2,
The signal filter is a wavelet filter based on wavelet transform operating simultaneously in the frequency and time domain, the imaging system. The method of claim 3,
The wavelet filter filters the frequency components of the delta-function ultrasound signal that are noisy at the first time and preserves the frequency components of the delta-function ultrasound signal at the second time. The method of claim 3,
The frequency spectrum of the wavelet filter replicates the frequency band of the N-type photoacoustic signal, an imaging system. The method of claim 5,
The wavelet filter provides smooth window edges that do not cause signal distortion upon convolution. The method of claim 5,
The wavelet filter converts anode pressure signals into unipolar signals, an imaging system. According to claim 1,
The deconvolution of the impulse response recovers an N-type pressure signal. According to claim 1,
The hand-held imaging probe has a curved concave arc shape. delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete delete An imaging method performed by an imaging system for mutually matched functional and anatomical mapping of tissues of part or all of a body,
a) transmitting ultrasound pulses to the tissue and detecting backscattered ultrasound signals reflected from structural tissue boundaries associated with body morphology;
b) transients resulting from the selective absorption of each of the two or more optical pulses in the hemoglobin and oxygen hemoglobin of blood contained in the tissues and sequentially delivering two or more optical pulses with different spectral bands of electromagnetic energy Detecting ultrasonic signals;
c) return signal changes during signal propagation through tissue and through system components that detect the backscattered ultrasound signals and the transient ultrasound signals, to remove noise, the temporal shape of the original signals and the ultrasound spectrum Processing the detected ultrasonic signals to restore the;
d) performing image reconstruction and further processing to cross-match partially transparent functional images reflecting total hemoglobin concentration and blood oxygen saturation and to generate morphological images of overlapping tissue structures; And
e) repeating steps a) to d) at a video frame rate so that real-time images display tissue functional and morphological changes when substantially changes occur. The method of claim 37,
Three or more optical pulses, each with electromagnetic radiation of different spectral bands, are sequentially delivered to the tissues to produce functional images with improved accuracy reflecting significant molecular chromophones of the tissue. The method of claim 38,
The molecular chromophore of the tissue comprises water, an imaging method. The method of claim 38,
The method of imaging, wherein the molecular chromophore of the tissue comprises lipids. The method of claim 38,
An imaging method in which each of the four optical pulses, each having electromagnetic radiation of different spectral bands, is sequentially delivered to the tissue. The method of claim 37,
The spectral bands of the two optical pulses are selected, one of the spectral bands of the two optical pulses coincides with the local maximum peak of hemoglobin absorption around 757 nm and the other band is the optical absorption band of oxygen hemoglobin. Imaging method, consistent with the spectral range around 1064 nm corresponding to the maximum ratio in absorption of hemoglobin. The method of claim 37,
And displaying, by the imaging system, tumor discrimination based on all or part of absorption coefficients measured in the tumor at first and second wavelengths. The method of claim 43,
Wherein the first wavelength comprises 757 nanometers. The method of claim 43,
Wherein the second wavelength comprises 1064 nanometers. The method of claim 43,
The stage of tumor discrimination is:
a. The process of displaying the tumor in a relatively smooth form superimposed with a relatively low increase in the concentration of total hemoglobin and normal blood oxygen saturation to indicate a benign tumor, or tissue surrounding the tumor, or
b. Imaging method comprising displaying a superimposed, coarse form of tumor, or tissue surrounding the tumor, with a high increase in total hemoglobin and low blood oxygen saturation to indicate a malignant tumor. The method of claim 37,
In order to make objects appear larger with greater depth, the relative brightness of image pixels corresponding to the surface of the tissue is reduced, thereby amplifying the relative brightness of pixels corresponding to larger depths in the tissue. And renormalizing the display palette. The method of claim 37,
Processing the detected ultrasonic signals to return the signal changes is hardware delivered to obtain a profile of intrinsic photoacoustic amplitude and distribution of the detected ultrasonic signals and optical absorption coefficient in the body. An imaging method comprising deconvolution of a function. The method of claim 37,
An imaging method of imaging part or all of the body in which optoacoustic and ultrasonic contrast agents have been used to enhance the visualization of contrast and tissue morphology of functional information within a portion of the body to be imaged. The method of claim 37,
An imaging method of imaging part or all of the body in which optoacoustic and ultrasonic contrast agents have been used to enhance the characterization of the distribution of molecules, cells, or tissues of predetermined types in the body.

Images