JP7510697B2 - Non-invasive diffuse acoustic confocal 3D imaging - Google Patents

Description

The present technique is a method for non-invasively scanning a body or a body part to obtain a three-dimensional image in which size, position, and shape are determined. The method further allows for the determination of the condition of the body or body part. A scanning device is also provided.

The use of radiation beams to obtain information about an object by detecting the amplitude or phase of the beam is well known for scientific and medical purposes. For example, phase information of a beam passing through an object can provide information about the temperature, composition, magnetic or electrostatic fields of the object, while amplitude measurements can provide information about the opacity or density of the object. The beam is composed of waves of radiation, and the waves Φ can be mathematically expressed in terms of both amplitude A and phase θ as:
Φ=Aexp(θ) 1)
can be written as:

The information gained from this method depends on whether one is detecting the amplitude of the beam wave or both the amplitude and phase of the beam wave. If the method only measures the amplitude of the beam, as in the case of x-rays, then only concentration differences within the object are reported. This is a limitation of the technique as it does not provide information such as the temperature, composition, elasticity, strain fields, magnetic fields or electrostatic fields of the object. A further drawback of many imaging techniques, such as x-ray imaging, is the intensity of the radiation used. When used for diagnosis, the levels used can be damaging to cells in the body.

Acoustic microscopes, including ultrasound, are now widely used to image the inside of the body, such as the fetus in the womb and the blood flow in arteries and veins. These microscopes emit a collimated acoustic beam that is reflected as a diffuse and scattered beam when it comes into contact with surfaces, such as bones, and interfaces, such as the interface between the embryonic fluid and the fetus. The reflected beam is used to measure the intensity of the acoustic beam. These microscopes cannot measure the intensity and phase of the beam passing through or reflecting off soft tissues, such as muscle or embryonic fluid. These microscopes also cannot measure temperature or composition, since they only use the intensity of the acoustic beam, not its phase. Thus, the images are not suitable to provide information other than that related to the surface or interface. A further disadvantage of these microscopes is that the images produced have a significant amount of background intensity caused by diffuse scattering of the beam after hitting a surface or interface. Taking the prostate as an example, ultrasound images have difficulty identifying interfaces between the prostate and other tissues, and can identify the urethra, but not abnormalities within the prostate.

Another method of measuring the amplitude of a beam is confocal microscopy. Confocal scanning laser microscopy was developed in the 1980s to view three-dimensional objects. Confocal scanning laser microscopy produces a three-dimensional amplitude image of an object by using a laser beam passing through the object and detecting the amplitude of the beam through a pinhole aperture placed confocal with a point on the focal plane of the object.

Confocal microscopes currently find a wide range of applications in materials, biology, and medicine. As a diagnostic tool, confocal microscopes are limited to detecting only concentration differences in thin tissues and objects that result in amplitude differences in the detected beam. The beams cannot penetrate far into tissues and other materials. They do not measure phase information of objects. Therefore, confocal microscopes cannot measure the composition or temperature of objects.

If the method measures the change in phase of the beam, it can provide information about the temperature and composition of the object. An acoustic beam can be used for this. The phase of the acoustic beam varies with the refractive index of the object, which depends on the temperature and composition of the object, and is a measure of the speed of sound in the acoustic beam.

The absolute phase of an object can be measured using a confocal scanning holographic microscope, as described in U.S. Pat. No. 7,639,365. This approach cannot be used to image the inside of the human body because laser beams do not pass easily through the body. The relative phase of an object can be measured using an acoustic confocal interferometer microscope, as described in U.S. Pat. No. 8,485,034. This approach requires a complex arrangement of interfering beams, mirrors and prisms, and is not suitable for imaging the inside of the human body due to geometric constraints.

Standard interference microscopes, standard holographic microscopes, and standard holographic interference microscopes have been used to measure both the phase and amplitude of an object, giving important information such as the density, composition, and temperature of the object. These microscopes generate three-dimensional amplitude and phase images of the object by measuring both phase and amplitude. As they are optical microscopes, the three-dimensional information measured from these microscopes is only from the surface of the object and not detected at points within the object. In all cases, a reference beam and an object beam are used to collect data, which results in the formation of an image. This limits the use of these microscopes to collecting data from or about the surface of an object. Thus, in medical diagnosis, they are potentially useful for diseases of the skin, but not for diseases of internal tissues or organs.

Another means by which the amplitude and phase of an object can be measured using an acoustic beam is spatially filtered transmission ultrasound phase imaging, as disclosed in U.S. Patent Nos. 6,679,846, 6,436,046, 6,132,375, and 6,193,663. Spatially filtered transmission ultrasound phase imaging involves measuring the amplitude and phase of the emitted beam, then measuring again the amplitude and phase of the acoustic beam after it has passed through the object upon arrival at the detector. The difference in amplitude and phase is due to the object. From the source, the beam spreads outward and diffuses, resulting in unwanted background scattering. Within or around this background scattering is the image of interest. The image represents the interfaces of the object being imaged. It does not represent a three-dimensional image, nor can it identify diseased tissue within the tissue or organ of interest. Similarly, in materials, unless there are interfaces, such as between a liquid and a solid, it cannot provide a three-dimensional image, nor can it show different materials within the material, or regions with different physical properties within the material.

US Patent Application Publication No. 20040059265 discloses non-invasive focusing of acoustic energy on a mass, such as a tumor, in tissue to reduce or eliminate the mass. The presence of a mass in tissue is detected by applying acoustic energy to the material. The mass is localized to determine its location. A temporal signature is developed to drive acoustic energy on the mass. Dynamic focusing of acoustic energy on the mass to reduce or eliminate it is achieved using the temporal signature. Mass imaging is performed using stationary ultrasound imaging techniques (the use of multiple parallel beams from an acoustic source outside the patient's body that are reflected from an object of interest inside the body to provide intensity information (the reflected signals are referred to as virtual sources, but are not the source) and then either modeling or time reversal is used to drive acoustic energy that is then focused on the mass to process it. Drawbacks of this method include the fact that it is based on models and assumptions, the fact that imaging of masses is not as good as images obtained using ultrasound, how they are imaged, and the fact that the information that can be reached using this method is limited. Also, because 3D position cannot be determined, if used to process chunks, such processing relies on models and assumptions, or time reversal, to be correct.

It would be beneficial to provide devices, systems and methods that can detect both the amplitude and phase of the beam. Such devices, systems and methods can provide information about the density, temperature, composition, elasticity, strain field, magnetic or electrostatic fields of an object. This is of great importance in the medical field, as information about density, temperature and composition can be obtained, potentially diagnosing, treating and evaluating the effectiveness of treatments for diseases such as cancer. Ideally, the device would be suitable for hand-held use with a variety of differently shaped detector holders for application to various parts of the body (e.g., but not limited to, the prostate, breast, head and skin). It would be advantageous if the same beam used to detect and provide a three-dimensional image of an object could also be used to process the object.

Examples of applications where temperature and dry composition measurements are important include medical diagnostics aimed at understanding the function of organs, tissues and diseased areas within the body. Currently, medical researchers do not have good means to non-invasively measure the body's internal temperature and composition. The purpose of this technology is to provide such a capability.

This technique provides a way to create a three-dimensional image of an object of one optical density in another object of different optical density that is far superior to ultrasound methods. A coherent beam is focused on a virtual source outside the object and within the fluid or amorphous medium. A zone of interference can be created by creating an interference zone between the bypass beam and the reflected beam. This interference zone is defocused to provide a three-dimensional image.

In one embodiment, a method for imaging an object in a first material having a different optical density relative to the object is provided, comprising: focusing an acoustic coherent beam on a virtual acoustic source in a fluid or amorphous second material outside the first material; moving the virtual acoustic source in the fluid or amorphous material such that at least a plurality of scattered beams from the virtual acoustic source scan the first material, at least one scattered beam is reflected from the object to form a reflected beam, and at least one scattered beam bypasses the object to form a bypass beam, the reflected beam and the bypass beam intersecting each other to form a coherent interference zone; and defocusing the coherent interference zone to provide Fresnel fringes, the Fresnel fringes forming an image of the object.

The method may further include an acoustic detector for detecting the image of the object.

The method may further include moving the virtual acoustic source through the fluid or amorphous material such that at least one scattered beam passes through the object and becomes an object beam and is detected by the acoustic detector.

The method may further include comparing the phase of the object beam with the phase of the bypass beam to provide information about the object.

The method may further include comparing the amplitude of the object beam with the amplitude of the bypass beam to provide information about the object.

In this method, the information from the phase may be the temperature, composition, magnetic field, or electrostatic field of the object, and the information from the amplitude may be the optical density of the object.

The method may also include determining the speed of sound of the object.

In this method, the sound speed of the object may be used to identify the object.

In the method, the object may be identified as a tumor or a lesion within the first material.

In this method, the acoustic coherent beam may be focused within a fluid or amorphous second material within the patient's body.

In this method, the acoustic coherent beam may be focused on one of urine in the bladder, adipose tissue, brain tissue, peritoneal fluid, gastric fluid, gallbladder, and spleen.

In this method, the acoustic coherent beam may be focused within a fluid outside the patient's body in which at least a portion of the subject's body is immersed.

In another embodiment, a system for imaging an object within a first material of different optical density relative to the object includes a coherent acoustic beam source emitting a coherent acoustic beam, a focuser positioned to focus the coherent acoustic beam on a virtual acoustic imaging source, the virtual acoustic imaging source emitting a scattered beam, a focuser, a coherent acoustic beam source actuator in mechanical communication with the coherent acoustic beam source, a focuser actuator in mechanical communication with the focuser, and a processor in electrical communication with the coherent acoustic beam source actuator and the focuser actuator. A system is provided that includes a coherent acoustic beam source and a memory in communication with the processor having instructions for directing the processor to move the coherent acoustic beam source and the focuser such that at least a portion of the scattered beam is reflected by the object into a reflected beam and passes through the object into a bypass beam, and such that the reflected beam and the bypass beam form an interference zone, the memory being further configured to move the coherent acoustic beam and the focuser to generate Fresnel fringes in the interference zone, and an acoustic detector positioned to image the Fresnel fringes.

In this system, the memory may include instructions for the processor to sharpen the image.

The system may further include a spatial filter in front of the acoustic detector.

The system may further include a conical reflector between the coherent acoustic beam source and the focuser.

The system may further include a pair of articulated arms, each having a distal end, and a second acoustic detector, each acoustic detector mounted on the arm proximate the distal end.

The system may be sized to image the ovaries.

In yet another embodiment, a system for imaging an object within a first material of different optical density relative to the object, comprising:
an apparatus including: a coherent acoustic beam source emitting a coherent acoustic beam; a focuser positioned to focus the coherent acoustic beam on a virtual acoustic imaging source, the virtual acoustic imaging source emitting a scattered beam; a coherent acoustic beam actuator in mechanical communication with the coherent acoustic beam source; a focuser actuator in mechanical communication with the focuser; and an acoustic detector;
A system is provided comprising: a computing device including a processor, a user interface, and a memory, the processor in electrical communication with the acoustic detector, the memory in communication with the processor and having instructions directing the processor to display an image in the user interface.

In this system, the memory may include instructions for the processor to sharpen the image.

The system may further include a spatial filter in front of the acoustic detector.

In this system, the focuser may be a concave curved cone-shaped reflector.

The system may further include a conical mirror surface between the concave curved cone-shaped reflector and the acoustic detector.

In the system, the processor may be in electrical communication with the coherent acoustic beam source actuator and the focuser actuator, the memory may be in communication with the processor and have instructions directing the processor to move the coherent acoustic beam source and the focuser, such that at least a portion of the scattered beam is reflected by the object and becomes a reflected beam, passes through the object and becomes a bypass beam, and the reflected beam and the bypass beam form an interference zone, the memory may be further configured to move the coherent acoustic beam and the focuser to generate Fresnel fringes in the interference zone, and the acoustic detector may be positioned to image the Fresnel fringes.

1 is a schematic diagram illustrating the operation of a diffuse acoustic confocal three-dimensional imaging device according to the present technology. FIG. 2A is a schematic diagram showing a beam used in a prior art ultrasound imaging device, and FIG. 2B is a schematic diagram showing a beam used in the present technology. 1 is an embodiment of a detection system of the present technology. 1 is an embodiment for imaging the ovaries using the present technology. FIG. 5 is a schematic diagram of the operation of a diffuse acoustic confocal 3D imager for scanning an ovary using one embodiment of FIG. 1 is an embodiment for imaging both ovaries using the present technology. 1 is one embodiment for imaging the breast using the present technology. 1 is an embodiment for imaging the brain using the present technology. 1 is an embodiment of a detection system of the present technology. 10A is a schematic diagram of a slice through the detection system of FIG. 9, FIG. 10B is a schematic diagram of a top view of a slice through a conical mirror, and FIG. 10C is a schematic diagram looking into the conical mirror. FIG. 10 is a schematic diagram showing the path of a beam in one embodiment of FIG. 12A is a side view of the detection system of FIG. 9, and FIG. 12B is a perspective view of the detection system of FIG.

Unless otherwise expressly stated, the following rules of interpretation apply in this specification (detailed description, claims and drawings): (a) All words used herein shall be interpreted as their grammatical gender or number (singular or plural) as the context requires. (b) The singular terms "a", "an" and "the" used in the specification and the appended claims include plural references unless the context clearly dictates otherwise. (c) The antecedent term "about" applied to a cited range or value indicates an approximation within the deviation in the range or value known or expected in the art from the measurement method. (d) The words "herein", "hereby", "therein", "thereto", "before" and "hereafter" and words of similar import refer to this specification as a whole and not to any particular section, claim or other subdivision unless otherwise specified. (e) The headings of the description are for convenience only and do not control or affect the meaning or construction of any part of the specification. (f) "or" and "any" are not exclusive, and "including" and "including" are not limiting. Additionally, the terms "comprise," "have," "include," and "contain" are intended to be open-ended terms (i.e., meaning "including, but not limited to") unless otherwise specified.

The recitation of ranges of values herein is intended to be a convenient shorthand method of referring individually to each separate value within the range, unless otherwise indicated herein, and each separate value is incorporated into the specification as if it were individually recited herein. When a specific range of values is provided, it is understood that each intervening value, to the tenth of the unit of the lower limit, is included within the other stated value or intervening value between the upper and lower limits of that range, and within other stated ranges. All smaller subranges are also included. The upper and lower limits of these smaller ranges are also included therein, subject to any specific excluded limits in the stated range.

Unless otherwise defined, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the relevant art. Although any methods and materials similar or equivalent to those described herein can also be used, acceptable methods and materials are described herein.

1 shows a schematic overview of a system, generally referred to as 8, for imaging a tissue, organ, or body part (object), generally referred to as 10. It generates and uses a scattered beam that is considered to be noise (speckle noise). A coherent acoustic source 12, such as a coherent acoustic emitter, emits a single coherent acoustic beam 14. The coherent acoustic source 12 can be moved manually or can be moved with a source actuator 16 in mechanical communication with the coherent acoustic source 12. The source actuator 16 is preferably controlled by a processor 18 under control of a memory 19, which has instructions directing the processor 18 to operate the source actuator 16. The coherent acoustic source 12 provides a coherent acoustic beam 14 having a beam frequency between and including about 0.5 megahertz and about 100 megahertz to obtain information including one or more of density, temperature, composition, elasticity, or strain fields within a mammalian body.

The coherent acoustic beam 14 has a large cross-sectional area, typically on the order of one or several centimeters. The coherent acoustic beam 14 is directed to a cone-shaped reflector 22, which is then directed to a focusing mirror or lens 24, which is reflected by a curved surface and focused into a converging beam 30 that is transmitted to a fluid or amorphous medium 32. The focusing mirror 24, under the control of a processor 18, pivots under the control of a focusing mirror actuator 26, which in turn is controlled by a memory 19 having instructions directing the processor 18 to actuate the actuator 26. The cone-shaped reflector 22 is under the control of an actuator 28, which moves it towards and away from the acoustic emitter (source) 12. The actuator 28, under the control of the processor 18, is controlled by a memory 19, which has instructions for directing the processor 18 to actuate the actuator 28. In one embodiment, the fluid or amorphous medium 32 is within the body, for example, but not limited to, urine in the bladder, fatty tissue, brain tissue, peritoneal fluid, gastric fluid, gallbladder, and spleen. The focused beam 30 converges to a point that is a virtually focused acoustic imaging source 34 at the crossover point. The processor 18 under control of the memory 19 is configured to direct the source actuator 16 to move the coherent acoustic source 12 to move the source 34 into the fluid or amorphous medium 32 and around within the fluid or amorphous medium 32. The processor 18 under control of the memory 19 is further configured to move the cone-shaped reflector 22 towards and away from the focusing mirror 24 to move the virtual source 34 towards and away from the acoustic emitter 12 to reposition the virtual source 34 in the fluid or amorphous medium 32. The virtual source 34 is located outside the object being imaged and inside the fluid or amorphous medium 32. The source 34 transmits multiple beams 36 that are scattered three-dimensionally in all directions. The scattered beams 36 pass out of the medium 32. By moving the source 34 within the fluid or amorphous medium 32, the multiple beams 36 scan around the outside of the medium 32. The beams 36 enter any object 10 that they encounter and then exit the object 10 as object beams 38, which are detected by an acoustic detector 40. The acoustic detector 40 is directed toward the source 34 so that it can detect the object beams 38. The object beams 38 refract as they pass through the object 10. The acoustic detector 40 can be moved to collect object beams 38 with a range of angular directions. The object beams 28 refract as they pass through the object 10. The acoustic detector 40 moves toward and away from the object 10 to defocus the image made by the combination of the reflected beams 37 and bypass beams 39, so that it appears photographic. The detector actuator 42 is in mechanical communication with the acoustic detector 40 and is under the control of a processor 18 in electronic communication with the detector actuator 42. Again, the processor 18 receives instructions from a memory 19. The object beam 38 also contains information about the object 10. The information transmitted by the object beam 38 is analyzed to determine its amplitude and phase according to techniques known in the art. The phase information of the object beam 38 provides information about the temperature, composition, magnetic or electrostatic fields of the object, and the amplitude measurement provides information about the opacity or concentration of the object. A spatial filter 46 reduces noise from the undesired scattered beam 36 and is positioned in front of the acoustic detector 40.

A transmission medium 60 is used between the patient and each of the coherent acoustic source 12 and the acoustic detector 40. Alternatively, the patient or relevant parts of the patient may be immersed in a fluid such that there is a fluid interface between the patient and the device 8. To observe the object 10, the source 34 is moved within the medium 32 by pivoting the focusing mirror 24 and the acoustic detector 40, by moving the microscope 6, or by repositioning the patient. A vector network analyzer is not required because the amplitude and phase information of the emitted and received intensities is not used to generate the image. However, a better intensity image can be obtained using a vector network analyzer for time filtering. An intensity image using Fresnel fringes is formed without the use of time and spatial filters, but the use of these filters improves the intensity image by being able to reduce the apparent size of the virtual source, i.e., better spatial resolution is obtained.

For the phase or velocity of sound imaging, a vector network analyzer is needed to measure the time difference for receiving the acoustic beam at each element in the detector. By measuring the time with a vector network analyzer, the velocity (m/s) can be determined, as is the path length traveled by the acoustic beam from the emitter/lens assembly to the detector.

Figures 2A and 2B show a comparison of beams and how they generate images in prior art ultrasound (2A) and the diffuse acoustic confocal imaging device of the present disclosure (2B). A coherent beam 14 strikes an object 10 and generates a reflected beam 2 and a scattered beam 4. The reflected beam 2 generates the image (note that only interfaces are identified, as a relatively large difference in refractive index between tissue types is required to identify the interface). The scattered beam 4 generates noise that must be removed to generate a clean image.

Using a diffuse acoustic confocal imaging system, the emitted coherent beam 14 is very flat (spatial coherence) and constant (temporal coherence). The coherent beam 14 is focused into a converging beam 30 that converges to a point, a virtual source 34 that emits a scattered beam 36. The virtual source 34 can move so that the scattered beam 36 can strike the object 10 from many spots, distances and angles. The scattered beam 36 either misses the object 10, strikes the edge of the object and is reflected off the edge, or passes through the object 10.

The one reflected from the object is called the reflected beam 37. The one that passes through the object is called the object beam 38. The one that avoids the object but overlaps with the reflected beam 37 is called the bypass beam 39.

The object beam 38 can be considered for diagnostic purposes of the object 10. The emission time of the beam is measured to form a sound image velocity using temporal coherence, and then the arrival time of the beam at each pixel in the image is measured. Any differences in sound velocity across the image can be used to diagnose structures within the image.

Spatial coherence interference between the reflected beam 37 and the bypass beam 39 in the overlap is used to create the image. The image of the object 10 can be considered as an in-line hologram for diagnostic purposes of the object 10. More specifically, the principle of Fresnel diffraction is used to create the image. The bypass beam 39 overlaps with the reflected beam 37 to create an interference zone 141. This interference zone 141 between these two beams forms the Fresnel interference fringes or defocus fringes. Preservation of coherence in the bypass beam 39 and the reflected beam 37 is required to form the interference fringes or Fresnel fringes in the image. Some of the intensity of the defocus or Fresnel fringes comes from the refracted object beam 38 exiting the object 10 and interfering with the bypass beam 39, as do the bypass and reflected beams, but the refracted object beam does not play a significant role in the Fresnel fringe image formation process.

Fresnel diffraction produces fringes (Fresnel fringes) in the acoustic image when defocusing occurs. Fresnel fringes enhance the contrast between features and background, allowing imaging of soft tissues and interfaces between different soft tissues. This includes tissues with very small refractive index differences, such as, but not limited to, breast tissue and mammary glands within breast tissue, tissue lesions, and eggs in the fallopian tubes.

The width of the overlap increases with defocus, causing the width of the Fresnel fringes to increase. Defocus decreases to zero where the object and camera are on the same plane. In this state, the object disappears because there is no overlap and therefore no fringes can be produced.

In addition, spatial resolution is determined by the width of the Fresnel fringes. The smallest width of Fresnel fringes seen in the image is on the order of the size of the virtual source. The size of the virtual source is determined by the focusing ability of the emitter/lens assembly and the wavelength of the emitted beam from the emitter. The resolution can approach the wavelength of the acoustic beam, approaching the size of a cell for a 50 MHz acoustic beam (λ = frequency/velocity = 50,000,000 times per second/1500 m/s = 0.00003 = 30 microns). This is a much higher resolution than ultrasound.

Images formed with large defocus, i.e., wide fringe lines, can be processed in a processor that sharpens features in the image (i.e., Fresnel fringes) by applying an amount of defocus, delta f, and knowing the cone angle, α, of the beam such that the reduction in fringe width is delta f times ½ of the cone angle. Similarly, knowing the cone angle and the change in fringe width with known or measured defocus can be used to determine the distance or position, z, of objects in the image, making it possible to generate 3D images since the lateral dimensions x, y are already measurable in the image. This is another advantage of the transmitted images generated by this technique over reflected images (ultrasound), where z is obtained by complex reflection time measurement of the reflected beam.

The distance between the virtual source 34 and the object 10 determines the magnification of the object. The closer the virtual source 34 is to the object 10, the closer the magnification factor is to 1. The magnification of the object increases the closer the virtual source 34 is to the object 10.

The spatial filter 46 and the temporal filter 54 (see FIG. 1) provided by the software in the memory 19 limit the volume of the acoustic virtual source 34 used for the image; the smaller the volume, the better the image resolution. The spatial resolution is set in part by the size of the converging beam 30 in the focused virtual source 34. The spatial filter 46 defines the lateral x, y dimensions or angular acceptance angle of the virtual source 34. The spatial filter 46 can be smaller than the virtual source 34 and is one of the factors that determine the spatial resolution. This may be the determining factor that determines the spatial resolution in the x, y plane. The temporal filter 54 determines the z or axial spatial resolution of the virtual source 34. This determines the acceptance, start and stop times for receiving an acoustic signal.

In the embodiment shown in FIG. 3, the fluid or amorphous medium 132 is outside the human body and is contained within the container 102. The container 102 is preferably adjustably attached to the device 110. The components of the system are as described above. The convergent beam 130 converges to a point that is a virtual convergent acoustic imaging source 134 at the crossover point. The processor 18 is under the control of the memory 19 and is configured to move the source actuator 16 to direct the coherent acoustic source 12 to move the virtual source 134 within the fluid or amorphous medium 132 and move it around within the fluid or amorphous medium 132. The virtual source 134 transmits a plurality of beams 136 scattered three-dimensionally in all directions. These scattered beams 136 pass from the medium 132. By moving the virtual source 134 within the fluid or amorphous medium 132, the multiple scattered beams 136 scan the outside of the medium 132. Depending on the angle of incidence, the scattered beams 136 will enter into any object 10 they encounter, pass through the object 10, and be detected by the acoustic detector 140 as object beams 138. Other scattered beams 136 are reflected from the object 10 and are called reflected beams 137. Scattered beams 136 that avoid the object but overlap with the reflected beam 137 are called bypass beams 139. The acoustic detector 140 is positioned to detect multiple scattered beams 136 that are focused on the source 134 and in close proximity to the object 10. The source 134 moves within the medium 132 as necessary to obtain a complete image of the object 10. The acoustic detector 140 can be moved to collect scattered beams 136 with a range of angular orientations, as shown in FIG. 1. The detector actuator 142 is in mechanical communication with the acoustic detector 140 and is under the control of a processor 144 in electrical communication with the detector actuator 142. Coherent interference between the reflected beam 137 and the bypass beam 139 at the overlap (interference zone 141) is used to create an image. The image of the object 10 can be considered an in-line hologram for diagnostic purposes of the object 10. The information carried by the object beam 136 can also be analyzed by the processor 18 to determine its amplitude and phase according to techniques known in the art. The amplitude and phase are compared to the amplitude and phase of the bypass beam 139.

To observe the object 10, the source 134 moves inside the medium 132 by pivoting the focusing mirror 124 and the acoustic detector 140, by moving the device 110, or by repositioning the patient. If the object 110 is, for example, a disease, mass, tumor, growth, etc., coherent interference between the reflected beam 137 and the bypass beam 139 in the overlap (interference zone 141) is used to create an image of the object. Fresnel fringes, as mentioned above, generate the image and are caused by defocusing.

As shown in FIG. 4, in one embodiment, a device 200 for imaging an ovary is provided. An acoustic emitter 202 is located at one end of the device and a focusing mirror 204 is located at or near the other end of the device 200. An acoustic detector 206 is aligned with the ovary. The detector 206 may be a linear array detector. A schematic diagram is shown in FIG. 5. As described above, a coherent beam 210 is emitted by the acoustic emitter 202 and focused by the focusing mirror 204 into a virtual source 212, which in turn generates a scattered beam 214 in three dimensions. The acoustic detector 206 detects an object beam 216, which is a beam passing through an object 217, in this case the ovary. The acoustic detector 206 also detects Fresnel fringes 218 generated by the bypass beam 220 and the reflected beam 222, as described above. A virtual source 212 is placed in the bladder 213 and moves around so that the scattered beam 214 scans around until it finds the ovary. The scattered beam 214 is then reflected as reflected beam 222, bypasses the ovary as beam 220, and passes through the ovary as object beam 216.

As shown in FIG. 6, a device 300 for imaging both ovaries is provided. An acoustic emitter 302 is centrally located within a vertical housing 304 that also houses a focusing mirror 306. The device 300 has two arms, a right arm 308 and a left arm 310. Proximate the ends 312 of the arms 308, 310 are located a right detector 314 and a left detector 316, respectively. The arms 308, 310 include articulated segments 320 under the control of a mechanical manipulator 322. As an additional means of detecting the beam, a right detector actuator 324 and a left detector actuator 326 are in mechanical communication with the acoustic detectors 314, 316 and are under the control of a processor 344 in electrical communication with the detector actuators 324, 326.

As shown in FIG. 7, a handheld device 400 for imaging the breast is provided. It includes an acoustic emitter 402, a focusing mirror 404, and an acoustic detector 406 arranged as described above. It is manually controlled and has a Bluetooth radio 408 for communicating with a processor 410.

As shown in FIG. 8, a system including a handheld emitter device 500 can be used to image the brain. One or more detectors 502 are positioned as needed and can be easily held in position by the patient simply tilting their head over the detector. A virtual source 504 is focused on the fatty tissue of the brain and a scattered beam 506 is emitted from the virtual source 504 in three dimensions. The beam that passes through the object 510 is the object beam 514. Again, an image is created by defocusing the interference zone 508 between the reflected beam 510 and the bypass beam 512.

9, in one embodiment, the device generally referred to as 600 has a concave curved cone-shaped reflector 610 replacing the cone-shaped reflector and a cone mirror 612 replacing the focusing mirror. Functionally, the concave curved cone-shaped reflector 610 is a focusing lens/mirror. An acoustic emitter 614 is located at a base 615 of the interior 617 of the cone mirror 612 and is aligned with the concave curved cone-shaped reflector 610. A support base 616 and a support member 618 hold the concave curved cone-shaped reflector 610 within the interior 617 of the cone mirror 612. A detector 620 may also be held by the support base 616. A housing 622 holds the cone mirror 612. The device 600 includes a virtual aperture 624 that is created by two features of the device 600. First, there is essentially a low-pass filter created by the concave curved cone-shaped reflector, which reflects the intensity of the planar (coherent 640) beam from the optical axis 648 (see FIG. 11) to the edge of the concave curved cone-shaped reflector 610, i.e., to its maximum diameter. Acoustic intensity with a diameter larger than the diameter of the concave curved cone-shaped reflector 610 is not reflected by it and passes through the device 600. It has an outer limit of passage, determined by the radial distance to the edge of the cone mirror 612, i.e., the hole made by the cone mirror 612 through which the acoustic beam must pass to focus at the virtual source 646 location. Another obstacle to the intensity reaching the virtual source 646 location is the cutoff of the reflected intensity from the cone mirror 612 reaching the virtual source 646 location by the lower part 632 of the concave curved cone-shaped reflector 610 (see FIG. 10A).

The path of the acoustic beam from the acoustic emitter 614 to the conical mirror 612 is shown in Figures 10A-C. A coherent beam 640 is emitted from the acoustic emitter 614 and hits the concave curved conical reflector 610. Figure 10A is a thin slice through the concave curved conical reflector 610 and the conical mirror 612. As can be seen from the slice, the sides of the conical mirror 612 are flat, while the sides of the concave curved conical reflector 610 are concave. The coherent beam 640 reflects from the concave curved conical reflector 610 and converges, thus generating a convergent beam 642. The convergent beam 642 hits the flat surface 626 of the conical mirror 612. As shown in Figure 10B, the convergent beam 642 from the slice of Figure 10A hits the flat wall of the conical mirror 612 in one plane. As shown in FIG. 10C, the converging beam 642 from the fragment of FIG. 10A forms a ring 628 on the flat surface 626 of the conical mirror 612. As known to those skilled in the art, the coherent beam 640 impinges on the concave surface of the concave curved conical reflector 610, around the concave curved conical reflector 610, and 360 degrees from the top 630 to the bottom 632 of the concave curved conical reflector 610.

The beam paths are also shown in FIG. 11, where coherent beams 640 emanate from an acoustic emitter 614 and strike a concave curved conical reflector 610. The coherent beams 640 reflect and converge from the concave curved conical reflector 610, and are therefore converging beams 642. The converging beams 642 strike the flat surface 626 of the conical mirror 612 and are reflected as converging beams 644 that converge at the crossover point to a point that is a virtual converging acoustic imaging source 646. The virtual source 646 is at or near the optical axis 648 of the conical mirror surface 612. Note that the curvature of the concave curved conical reflector 610 is selected such that the virtual source 646 is at or near the optical axis 648.

As mentioned above, from the virtual source 646, the acoustic beam diffusely scatters the acoustic beam in all directions. The acoustic detector 620 can be placed on the end of the concave curved cone-shaped reflector 610 or anywhere around the virtual source 646 pointing to the virtual source 646. Without being bound by theory, the design of Figs. 9-12B has the following advantages: The concave curved cone-shaped reflector 610 and the curved focusing mirror provide the same functionality, but the concave curved cone-shaped reflector is easier to manufacture than the curved focusing mirror. The concave curved cone-shaped reflector 610 can be easily replaced with another concave curved cone-shaped reflector 610 with different dimensions and curvatures that allow the focal depth of the virtual source 646 to be changed. The concave curved cone-shaped reflector 610 can be moved back and forth on the optical axis 648 to change the focal depth of the virtual source 646. The design increases spatial resolution and allows for smaller virtual source sizes. The concave curved cone-shaped reflector 610 reflects the convergent beam 642 to the cone mirror 612, where the surface of the cone mirror 612 is at a constant angle to the convergent beam 642 as the concave curved cone-shaped reflector 610 moves back and forth within the cone mirror 612. As described above, this allows the probe/virtual source 648 to move back and forth. This design increases the device's ability to focus the beam and therefore move the virtual source within the body.

An exemplary device is shown in Figures 12A and 12B. As shown in Figure 12A, an acoustic emitter 614 is at one end of a housing 622. As shown in Figure 12B, a concave curved cone shaped reflector is located at the interior 617 of the housing 622 and is held by a support member 618. A cone mirror surface 612 is inside the housing 622. The acoustic emitter 614 is at the other end of the housing 622. The device 600 may be handheld and adjusted and operated by a user, or may be under the control of a processor and memory within a computing device, as described above. In either case, the computing device includes a user interface on which images are displayed.

In all embodiments, the method includes an acoustic emitter that emits a coherent beam that is reflected and focused, or focused and reflected, to provide a point that is a virtual source within the amorphous material or fluid. The virtual source travels to emit a scattered beam, some of which passes through the object, some of which bypasses the object, and some of which reflects off the object. The bypass and reflected beams interfere with each other in an interference zone to defocus the image, providing an image that is seen as Fresnel fringes.

Although the exemplary embodiments have been described in connection with what are currently believed to be examples of the most practical and/or suitable embodiments contemplated, it should be understood that the description is not limited to the disclosed embodiments, but is intended to encompass various modifications and equivalent arrangements that are within the spirit and scope of the exemplary embodiments. Those skilled in the art will recognize, or be able to ascertain using no more than routine experimentation, many equivalents to the specific embodiments specifically described herein. Such equivalents are intended to be encompassed within the scope of the claims appended hereto or subsequently filed.

Claims (19)

1. A method for imaging an object within a first material having a different absorbance relative to the object, comprising: focusing an acoustic coherent beam on a virtual acoustic source within a fluid or amorphous second material outside the first material; moving the virtual acoustic source within the fluid or amorphous material such that at least a plurality of scattered beams from the virtual acoustic source scan the first material, at least one scattered beam being reflected from the object to form a reflected beam and at least one scattered beam bypassing the object to form a bypass beam, the reflected beam and the bypass beam intersecting each other to form a coherent interference zone; and defocusing the coherent interference zone to provide Fresnel fringes, the Fresnel fringes forming an image of the object. The method of claim 1, further comprising an acoustic detector for detecting the image of the object. The method of claim 2, further comprising the step of moving the virtual acoustic source in the fluid or amorphous material such that at least one scattered beam passes through the object and becomes an object beam and is detected by the acoustic detector. The method of claim 3, further comprising comparing the phase of the object beam with the phase of the bypass beam to provide information about the object. The method of claim 4, further comprising comparing the amplitude of the object beam with the amplitude of the bypass beam to provide information about the object. The method of claim 5 , wherein the information from the phase is the temperature, composition, magnetic field, or electrostatic field of the object, and the information from the amplitude is the absorbance of the object. The method of claim 6, in which the speed of sound of the object is determined. The method of claim 7, wherein the sound speed of the object is used to identify the object. The method of claim 1, wherein the object is identified as a tumor or a lesion within the first material. The method of claim 9, wherein the acoustic coherent beam is focused within a fluid or amorphous second material within the patient's body. The method of claim 10, wherein the acoustic coherent beam is focused on one of urine in the bladder, adipose tissue, brain tissue, peritoneal fluid, gastric fluid, gallbladder, and spleen. The method of claim 9, wherein the acoustic coherent beam is focused within a fluid outside the patient's body in which at least a portion of the subject's body is immersed. 1. A system for imaging an object within a first material of differing absorbance relative to the object, comprising: a coherent acoustic beam source emitting a coherent acoustic beam; a focuser positioned to focus the coherent acoustic beam onto a virtual acoustic imaging source, the virtual acoustic imaging source emitting a scattered beam; a coherent acoustic beam source actuator in mechanical communication with the coherent acoustic beam source; a focuser actuator in mechanical communication with the focuser; a processor in electrical communication with the coherent acoustic beam source actuator and the focuser actuator; and a processor for focusing the scattered beam. a memory having instructions in communication with the processor directing the processor to move the coherent acoustic beam source and the focuser such that at least a portion of a coherent acoustic beam is reflected off the object into a reflected beam and bypasses the object into a bypass beam, and such that the reflected beam and the bypass beam form an interference zone, the memory being further configured to move the coherent acoustic beam and the focuser to generate Fresnel fringes in the interference zone; and an acoustic detector positioned to form an image of the object by imaging the Fresnel fringes. The system of claim 13, wherein the memory includes instructions for the processor to sharpen the image. The system of claim 14, further comprising a spatial filter in front of the acoustic detector. The system of claim 15, further comprising a conical reflector between the coherent acoustic beam source and the focuser. The system of claim 16, further comprising a pair of articulated arms, each having a distal end, and a second acoustic detector, each acoustic detector mounted on an arm proximate the distal end. The system of claim 17, sized for imaging the ovaries. 1. A system for imaging an object within a first material of different absorbance relative to the object, comprising:
an apparatus including: a coherent acoustic beam source emitting a coherent acoustic beam; a focuser positioned to focus the coherent acoustic beam on a virtual acoustic imaging source, the virtual acoustic imaging source emitting a scattered beam; a coherent acoustic beam source actuator in mechanical communication with the coherent acoustic beam source; a focuser actuator in mechanical communication with the focuser; and an acoustic detector;
a computing device including a processor, a user interface, and a memory, the processor in electrical communication with the acoustic detector, the memory having instructions in communication with the processor directing the processor to display an image on the user interface, the processor in electrical communication with the coherent acoustic beam source actuator and the focuser actuator, the memory having instructions in communication with the processor directing the processor to move the coherent acoustic beam source and the focuser such that at least a portion of the scattered beam is reflected off the object into a reflected beam and bypasses the object into a bypass beam, and the reflected beam and the bypass beam form an interference zone, the memory further configured to move the coherent acoustic beam and the focuser to generate Fresnel fringes in the interference zone, and the acoustic detector is positioned to image the Fresnel fringes;
A system comprising:

Images