Ultrasound computed tomography method and ultrasound computed tomography apparatus

Abstract

The ultrasonic computed tomography apparatus comprises an imaging area configured to contain a medium to be imaged; an array of ultrasonic elements arranged around the imaging area, the array of ultrasonic elements comprising a transmitter located on a first side of the imaging area and a plurality of receivers located on the opposite side of the imaging area; a receiving circuit for each of the plurality of receivers; and a processing circuit configured to perform computed tomography. The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area. The single-tone continuous wave signal includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency. Each of the plurality of receivers is configured to receive a sequence of single-tone continuous wave signals transmitted through the imaging area. The receiving circuit of each of the plurality of receivers is configured to mix each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signals to obtain the in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The processing circuit is configured to process the obtained in-phase and orthogonal values ​​using computed tomography to obtain an image of the imaging region.

Description

[Technical Field]

[0001] This disclosure relates to an ultrasound computed tomography (CT) imaging method and an ultrasound computed tomography apparatus, for example, for the purpose of ultrasound computed tomography imaging of the breast. [Background technology]

[0002] Breast cancer is the most common cancer among women. It accounts for one-third of all cancers diagnosed and is the third leading cause of cancer death in women. Breast cancer survival rates depend critically on early diagnosis, and decline dramatically with delayed diagnosis. Despite dedicated breast screening programs in many countries, participation rates are generally low, and as a result, many cases go undetected until later stages.

[0003] X-ray mammography is considered the gold standard for initial breast cancer screening. While the amount of radiation absorbed after conventional X-ray mammography is generally very low, women undergoing X-ray mammography may absorb unnecessary ionizing radiation. Another common problem with mammography is that cancerous lesions may be missed in women with dense breast tissue. In other words, many cancer cases go undiagnosed in women with dense breasts. Therefore, improving the detection rate of such cases is important because it could significantly reduce mortality.

[0004] It is known that ultrasound is used to perform imaging diagnostics of the breast. In handheld ultrasound, a handheld transducer transmits ultrasound signals to a part of the body and receives ultrasound signals reflected from that part of the body.

[0005] A typical handheld ultrasound system is a pulsed system using direct RF sampling. A linear array consists of, for example, 128 or 256 elements. All elements are configured to both transmit and receive ultrasound signals. Only reflected data, also called pulsed echo data, is collected. The reflected signal is collected on one side, such as one side of the breast, which is the same side from which the ultrasound signal is transmitted. Image reconstruction is often performed in the time domain using conventional delayed-sum beamforming algorithms. [Overview of the project] [Problems that the invention aims to solve]

[0006] For screening purposes, widespread adoption of handheld ultrasound is considered impractical. Handheld ultrasound is operator-dependent, requiring the operator to hold the handheld transducer and select the ultrasound imaging position and settings. Handheld ultrasound has limited footprint and may interfere with imaging the entire breast.

[0007] There is a need to provide a safe and operator-independent method for imaging breast ultrasound for the diagnosis of breast cancer. [Means for solving the problem]

[0008] According to one aspect of the present disclosure, an ultrasonic computed tomography (CT) method is provided. The ultrasonic computed tomography method comprises: an imaging area configured to contain a medium to be imaged; an array of ultrasonic elements arranged around the imaging area, each comprising a transmitter located on a first side of the imaging area and a plurality of receivers located on the opposite side of the imaging area; a receiving circuit for each of the plurality of receivers; and a processing circuit configured to perform computed tomography. The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area. The single-tone continuous wave signal includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency. Each of the plurality of receivers is configured to receive a sequence of single-tone continuous wave signals transmitted through the imaging area. The receiving circuit of each of the plurality of receivers is configured to mix each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signals to obtain in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The processing circuit is configured to process the obtained in-phase and orthogonal values ​​using computed tomography to obtain an image of the imaging region.

[0009] The medium may include breast tissue. The medium may comprise human or animal tissue. The medium may comprise at least a portion of the breast, brain, legs, arms, chest, neck, or testicles.

[0010] The transmitter and receiver may form part of an array of ultrasonic elements arranged around the imaging area where the medium is placed.

[0011] The array may be circular or partially circular.

[0012] This method may include the step of moving the transmitter and receiver by a continuous position increment relative to the medium in a direction perpendicular to the plane formed by the transmitter and receiver, the step of performing further transmission of each of the sequences of single-tone continuous wave signals for each position increment, the step of processing each of the in-phase and orthogonal values ​​obtained from the further transmissions by computed tomography to obtain a two-dimensional image of the medium, and the step of combining the two-dimensional images to obtain a three-dimensional volume.

[0013] The array may be a three-dimensional array comprising multiple two-dimensional arrays, the two-dimensional arrays may optionally be circular or partially circular, and the multiple two-dimensional arrays may optionally each have different sizes.

[0014] The array may include a first ultrasonic element having a first characteristic and a second ultrasonic element having a second characteristic different from the first characteristic. The different characteristic may consist of at least one of a different height, a different width, or a different thickness. The different characteristic may consist of a different focusing characteristic. The different characteristic may consist of at least one of a different focal point, a different beamwidth, a different frequency bandwidth, or a different transmission frequency. The first ultrasonic element is associated with a first type of acoustic lens, and the second ultrasonic element is associated with a second different type of acoustic lens having different acoustic characteristics.

[0015] The first ultrasonic element transmits in a first low frequency bandwidth, and the second ultrasonic element transmits in a second high frequency bandwidth. The first low frequency bandwidth is, for example, between 100 kHz and 500 kHz, between 200 kHz and 300 kHz, or between 300 kHz and 500 kHz, or between 200 kHz and 500 kHz, or between 200 kHz and 300 kHz, or between 300 kHz and 500 kHz. The second high frequency bandwidth can be, for example, between 500 kHz and 1 MHz, between 1 MHz and 3 MHz, between 1 MHz and 2 MHz, or between 2 MHz and 3 MHz.

[0016] Different ultrasonic elements in a two-dimensional array may have different characteristics. A first two-dimensional array is composed of first ultrasonic elements having first characteristics, and a second two-dimensional array is composed of second ultrasonic elements having second characteristics different from the first characteristics.

[0017] The array may be arranged on a convex surface, for example, a bowl-shaped surface.

[0018] The array may comprise at least one pair of opposing arrays. The opposing arrays may be two-dimensional arrays. The method may further include the step of rotating at least one pair of opposing two-dimensional arrays around a medium. The opposing arrays may be located on different sides of the medium, for example, opposite sides of the medium. The opposing arrays may face each other. The opposing arrays may be arranged such that a signal transmitted from a first array of opposing arrays is received by a second array of opposing arrays through the medium. For each pair of opposing arrays, the distance between the opposing arrays may be controlled mechanically and / or electronically. The distance between the opposing arrays may vary, for example, depending on the elevation.

[0019] At least one pair of opposing two-dimensional arrays may comprise a first pair of opposing two-dimensional arrays equipped with ultrasonic elements having a first characteristic, and a second pair of opposing two-dimensional arrays equipped with ultrasonic elements having a second characteristic different from the first characteristic. The first ultrasonic elements may transmit in a first low frequency bandwidth, and the second ultrasonic elements may transmit in a second high frequency bandwidth.

[0020] The first frequency may be between 100 kHz and 20 MHz. The second frequency may be between 100 kHz and 20 MHz. The first frequency may be between 300 kHz and 3 MHz. The second frequency may be between 300 kHz and 3 MHz. Each of the first and second frequencies may be greater than 100 kHz, and may be optionally 200 kHz, optionally 300 kHz, optionally 500 kHz, and optionally 1 MHz. Each of the first and second frequencies may be less than 20 MHz, optionally 10 MHz, optionally 5 MHz, and optionally 3 MHz.

[0021] A single-tone continuous wave signal may include a single-tone continuous wave signal having a pre-programmed set of frequencies.

[0022] Each frequency in the pre-programmed set of frequencies may be greater than 100 kHz and may be arbitrarily 200 kHz, arbitrarily 300 kHz, arbitrarily 500 kHz, and arbitrarily 1 MHz. Each frequency in the pre-programmed set of frequencies may be less than 20 MHz and may be arbitrarily 10 MHz, arbitrarily 5 MHz, and arbitrarily 3 MHz.

[0023] A sequence of single-tone continuous wave signals may include two or more single-tone continuous wave signals having a first frequency. The ultrasound computed tomography method may further include the steps of: combining the common-mode values ​​of two or more single-tone continuous wave signals to obtain a combined common-mode value of a first frequency; and combining the orthogonal values ​​of two or more single-tone continuous wave signals to obtain a combined orthogonal value of the first frequency.

[0024] The step of processing the in-phase and orthogonal values ​​by computed tomography may include model-based image reconstruction. The step of processing the in-phase and orthogonal values ​​by computed tomography may use an iterative algorithm based on numerical optimization of the cost function. The iterative algorithm may include a step of determining the difference between the measured data or obtained data and the composite data. The iterative algorithm may include a step of iteratively updating estimates of physical quantities until convergence is reached. Processing the in-phase and orthogonal values ​​by computed tomography may include a step of processing by computed tomography with or without phase encoding.

[0025] According to an aspect of the present disclosure that may be provided independently, an ultrasonic computed tomography apparatus is provided. The ultrasonic computed tomography apparatus comprises: an imaging area configured to contain a medium to be imaged; an array of ultrasonic elements arranged around the imaging area, the array of ultrasonic elements comprising a transmitter located on a first side of the imaging area and a plurality of receivers located on the opposite side of the imaging area; a receiving circuit for each of the plurality of receivers; and a processing circuit configured to perform computed tomography. The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area. The single-tone continuous wave signal includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency. Each of the plurality of receivers is configured to receive a sequence of single-tone continuous wave signals transmitted through the imaging area. Each receiving circuit of the multiple receivers is configured to mix each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signal to obtain the common-mode and quadrature values ​​of each of the single-tone continuous wave signals. The processing circuit is configured to process the obtained common-mode and quadrature values ​​by computed tomography to obtain an image of the imaging region.

[0026] An array of ultrasonic elements may comprise a plurality of transceivers, each configured to transmit and receive ultrasonic signals. The transmitter may be the first transceiver of the plurality of transceivers. The receiver may be a further transceiver of the plurality of transceivers.

[0027] The array may be circular or partially circular.

[0028] The array may be movable in a continuous increment of position perpendicular to the plane formed by the transmitter and receiver with respect to the medium.

[0029] The array may be a three-dimensional array comprising multiple two-dimensional arrays. The two-dimensional arrays may be circular. Part of the two-dimensional array may be circular. The two-dimensional arrays may be of different sizes. The two-dimensional arrays may be circular arrays having different radii.

[0030] Ultrasonic elements in different two-dimensional arrays may have different characteristics.

[0031] The array may consist of multiple two-dimensional or three-dimensional subarrays. Each two-dimensional or three-dimensional subarray is configured to transmit and receive ultrasonic signals in different frequency bandwidths.

[0032] The array may be attachable and / or provided as attachable patches. For example, the array may consist of a flexible substrate or film. Transmitters and multiple receivers are mounted or embedded in the flexible substrate or film. The array may include an adhesive layer for bonding the flexible substrate or film to a medium.

[0033] Each two-dimensional or three-dimensional subarray may consist of a corresponding flexible substrate or film. Transmitters and multiple receivers of the two-dimensional or three-dimensional subarray are mounted or embedded in the corresponding flexible substrate or film. Each two-dimensional or three-dimensional subarray may include an adhesive layer for attaching the corresponding flexible substrate or film to a medium.

[0034] The array may be arranged on a convex surface, for example, a bowl-shaped surface.

[0035] The array may consist of at least one pair of opposing two-dimensional arrays. The method may further include the step of rotating the at least one pair of opposing two-dimensional arrays around a medium.

[0036] At least one pair of opposing two-dimensional arrays may consist of a first pair of opposing two-dimensional arrays equipped with ultrasonic elements having a first characteristic, and a second pair of opposing two-dimensional arrays equipped with ultrasonic elements having a second characteristic different from the first characteristic.

[0037] The receiver may include a mixer stage configured to mix each received single-tone continuous wave signal in a sequence of single-tone continuous wave signals with each single-frequency mixer signal having the same frequency as the single-tone continuous wave signal to obtain the in-phase and quadrature values ​​of each single-tone continuous wave signal.

[0038] The receiver may further include a low-noise amplifier configured to amplify the received signal, a pair of low-pass filters to remove frequencies above a threshold frequency, and a pair of analog-to-digital converters to digitize common-mode and quadrature values.

[0039] Each analog-to-digital converter may have a bit depth of at least 16 bits. Each analog-to-digital converter may have a bit depth of at least 20 bits. Each analog-to-digital converter may have a bit depth of at least 24 bits. Each analog-to-digital converter may have a bit depth of less than 64 bits, optionally less than 32 bits, and optionally less than 28 bits. Each analog-to-digital converter may have a bit depth of 24 bits.

[0040] The sampling rate of each analog / digital converter may be between 1kHz and 1024kHz, or between 1kHz and 1000kHz. The sampling rate of each analog / digital converter may be greater than 1kHz, optionally greater than 200kHz, and optionally greater than 500kHz. The sampling rate of each analog / digital converter may be less than 1000kHz, optionally less than 750kHz, and optionally less than 500kHz. The sampling rate of each analog / digital converter may be 128kHz. The sampling rate of each analog / digital converter may be 256kHz. The sampling rate of each analog / digital converter may be 512kHz.

[0041] The ultrasound computed tomography system may further include a controller. The controller may be configured to determine and / or store for each single-tone continuous wave signal at least one of the following: each voltage, each phase, and each transmission time. The controller may also be configured to determine and / or store for each single-tone continuous wave signal at each integration time.

[0042] At least one of the following may be programmable: each voltage, each phase, and each transmission time. At least one of the following may be frequency-dependent: each voltage, each phase, and each transmission time. Each integration time may be programmable: each integration time may be frequency-dependent.

[0043] The ultrasound computed tomography system may further include a signal generator configured to produce a sequence of single-tone continuous wave signals.

[0044] According to aspects of the present disclosure that may be provided independently, a method for ultrasonic computed tomography is provided. This ultrasonic computed tomography method comprises: an imaging area configured to contain a medium to be imaged; an array of ultrasonic elements arranged around the imaging area, each array of ultrasonic elements comprising a transmitter and a plurality of receivers; a receiving circuit for each of the plurality of receivers; and a processing circuit configured to perform computed tomography. The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area. The single-tone continuous wave signal includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency. Each of the plurality of receivers is configured to receive a sequence of single-tone continuous wave signals transmitted through the imaging area. The receiving circuit of each of the plurality of receivers is configured to mix each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signals to obtain in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The processing circuit is configured to process the obtained in-phase and orthogonal values ​​using computed tomography to obtain an image of the imaging region.

[0045] The transmitter may be placed on the first side of the medium. Multiple receivers may be placed on the second side of the medium. The second side of the medium is different from the first side of the medium.

[0046] The second side of the medium may be on the opposite side from the first side of the medium, or it may optionally be on the side of the first side of the medium.

[0047] The transmitter may be positioned on a first surface in space, and the plurality of receivers may be positioned on a second surface in space. At least one of the first and second surfaces may be planar. At least one of the first and second surfaces may be curved. At least one of the first and second surfaces may be concave. The first and second surfaces may be curved toward each other. At least one of the first and second surfaces may have circular symmetry, cylindrical symmetry, or axial symmetry with respect to the transmission direction of a sequence of single-tone continuous wave signals. The first surface may define at least a portion of a first hemisphere. The second surface may define at least a portion of a second hemisphere.

[0048] The transmitter may be positioned on a first line in space, and the plurality of receivers may be positioned on a second line in space. At least one of the first line and the second line may be straight. At least one of the first line and the second line may be curved. At least one of the first line and the second line may be concave. The first line and the second line may curve toward each other. At least one of the first line and the second line may be symmetrical with respect to the transmission direction of the sequence of single-tone continuous wave signals. The first line may define at least a portion of a first arc or a first semicircle. The second line may define at least a portion of a second arc or a second semicircle.

[0049] The step of processing in-phase and orthogonal values ​​obtained by computed tomography may use an iterative algorithm based on numerical optimization of the cost function. The step of processing in-phase and orthogonal values ​​obtained by computed tomography may include processing by computed tomography with or without phase encoding.

[0050] The method may further include the step of processing the obtained in-phase and orthogonal values ​​to obtain a B-mode image. Processing the obtained in-phase and orthogonal values ​​to obtain a B-mode image may include delay and sum beamforming.

[0051] According to an aspect of the present disclosure that may be provided independently, an ultrasonic computed tomography apparatus is provided. The ultrasonic computed tomography apparatus comprises an imaging area configured to contain a medium to be imaged; an array of ultrasonic elements arranged around the imaging area, the array of ultrasonic elements comprising a transmitter and a plurality of receivers; a receiving circuit for each of the plurality of receivers; and a processing circuit configured to perform computed tomography. The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area. The single-tone continuous wave signal includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency. Each of the plurality of receivers is configured to receive a sequence of single-tone continuous wave signals transmitted through the imaging area. The receiving circuit of each of the plurality of receivers is configured to mix each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signals to obtain in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The processing circuit is configured to process the obtained in-phase and orthogonal values ​​by computed tomography to obtain an image of the imaging region. The transmitter may be located on the first side of the imaging region. Multiple receivers may be located on the second side of the imaging region. The second side of the imaging region is different from the first side of the imaging region.

[0052] The second side of the imaging region may be on the opposite side from the first side of the imaging region, and may optionally be in the lateral direction.

[0053] The transmitter may be positioned on a first surface in space, and the plurality of receivers may be positioned on a second surface in space. At least one of the first and second surfaces may be planar. At least one of the first and second surfaces may be curved. At least one of the first and second surfaces may be concave. The first and second surfaces may be curved toward each other. At least one of the first and second surfaces may have circular symmetry, cylindrical symmetry, or axial symmetry with respect to the transmission direction of a sequence of single-tone continuous wave signals. The first surface may define at least a portion of a first hemisphere. The second surface may define at least a portion of a second hemisphere.

[0054] The transmitter may be positioned on a first line in space, and the plurality of receivers may be positioned on a second line in space. At least one of the first line and the second line may be straight. At least one of the first line and the second line may be curved. At least one of the first line and the second line may be concave. The first line and the second line may curve toward each other. At least one of the first line and the second line may be symmetrical with respect to the transmission direction of the sequence of single-tone continuous wave signals. The first line may define at least a portion of a first arc or a first semicircle. The second line may define at least a portion of a second arc or a second semicircle.

[0055] Features in one embodiment can be applied as features in other embodiments in any appropriate combination. For example, system features may be provided as method features, or vice versa. [Brief explanation of the drawing]

[0056] Next, embodiments of the present invention will be described using non-limiting examples, as shown in the following figures. [Figure 1A] This is a schematic diagram of a stepped frequency continuous wave ultrasound computed tomography (SFCW-USCT) system. [Figure 1B]Figure 1A is a schematic diagram of the ring of ultrasound transceiver elements in the SFCW-USCT system around the breast. [Figure 2] Figure 1A is a schematic diagram of the transmission and reception path of the SFCW-USCT system. [Figure 3] This is a plot of voltage as a function of time used to create a single-frequency continuous wave signal for use in the transmission and reception path shown in Figure 2. [Figure 4] This is a plot of the demodulated signals I and Q before and after low-pass filtering. [Figure 5] This figure shows a plot of a single-tone transmitted signal and its associated mixer waveform. [Figure 6] This figure shows a plot of a single received tone after amplification by a low-noise amplifier. [Figure 7] This figure shows plots of the demodulated signals I and Q in the time domain and frequency domain. [Figure 8] This figure shows the temporal changes of the low-pass filtered signals Ifilt and Qfilt. [Figure 9] Figure 8 shows a method for integrating the low-pass filtered signals Ifilt and Qfilt after digitization. [Figure 10] Figure 8 shows an alternative method for integrating the low-pass filtered signals Ifilt and Qfilt after digitization. [Figure 11] This figure shows the relationship between SNR and integral time. [Figure 12] This figure shows the absolute values ​​of the complex samples d(Rx, Tx; fk) measured by the side receiver and the counter receiver during single-tone transmission, and the effect of longer integration times on such counter receivers. [Figure 13] This figure shows a numerical phantom sample containing multiple inclusions and the associated measurement geometry. The phantom sample defines the spatial variation of the sound velocity in m / s units, and each inclusion brings about a spatially local variation of the sound velocity in m / s units. [Figure 14]This figure shows a numerical phantom sample defining the spatial variation of sound velocity in m / s, the development of reconstructed sound velocity images of the numerical phantom sample across five different frequency bandwidths, and the development of the associated cost function in iterative image reconstruction using frequency domain data measured using the SFCW-USCT system shown in Figure 1a. [Figure 15] This figure shows the measurement of breast shape to provide information for the imaging protocol. [Figure 16] This figure shows a first alternative shape of an ultrasonic transceiver element in which the ultrasonic transceiver element is located inside a bowl. [Figure 17] This figure shows a second alternative shape of an ultrasonic transceiver element in which the ultrasonic transceiver elements are arranged in three circular rings. [Figure 18] This figure shows a third alternative shape of an ultrasonic transceiver element having an array positioned on the opposite side of the medium. [Figure 19] This figure shows a fourth alternative configuration of an ultrasonic transceiver element, in which the ultrasonic transceiver element is arranged in two circular rings. Each ring is used for imaging in a different frequency bandwidth. [Figure 20] This is a schematic diagram of a fifth alternative arrangement of ultrasonic transceiver elements constituting a wearable patch. Each patch includes an ultrasonic transceiver array for 2D imaging. [Figure 21] This is a schematic diagram of a sixth alternative arrangement of ultrasonic transceiver elements constituting a wearable patch. Each patch includes an ultrasonic transceiver array for 3D imaging. [Figure 22] This is a schematic diagram of a hypothetical pulse system with direct RF sampling, including a voltage plot. [Figure 23] Figure 22 shows a Fourier synthesis plot related to a virtual pulse system obtained by direct RF sampling. [Figure 24] Figure 22 shows a plot of Fourier analysis related to a hypothetical pulsed system with direct RF sampling. [Figure 25]This figure shows that the operation of the virtual pulse system using direct RF sampling in Figure 22 is mathematically equivalent to the operation of the SFCW-USCT system in Figure 2. [Figure 26] This figure compares the signal-to-noise ratio of a signal measured using the virtual pulse system with direct RF sampling shown in Figure 22 with the signal-to-noise ratio of a signal measured using the SFCW-USCT system shown in Figure 2. [Modes for carrying out the invention]

[0057] Figure 1A is a simplified diagram of an ultrasonic computed tomography apparatus 100 according to an embodiment. The ultrasonic computed tomography apparatus 100 is configured to perform computed tomography, in which ultrasonic signals transmitted through a medium are processed to obtain an image.

[0058] The ultrasound computed tomography (CT) scanner 100 is a two-dimensional system. The elements are arranged in a two-dimensional array, and data is collected at each of several fixed elevation positions (elevation is in the z direction, orthogonal to the xy imaging plane of the two-dimensional array). The entire breast is reconstructed as a three-dimensional volume by superimposing the two-dimensional images.

[0059] The ultrasonic computed tomography (CT) scanner 100 consists of an array of transceiver elements 110A, 110B...110H arranged in a two-dimensional ring shape around the imaging area 120. Each of the transceiver elements 110A, 110B...110H is a transducer element configured to transmit and receive ultrasonic signals. In the example shown in Figure 1, the system has a radius of 110 mm. In other embodiments, any suitable radius can be used.

[0060] The imaging region 120 is configured to contain the medium to be imaged, which in this embodiment is breast tissue. In other embodiments, any suitable medium may be imaged, for example, any suitable type of human or animal tissue may be imaged.

[0061] The array of transceiver elements 110A, 110B, ..., 110H is movable in a direction perpendicular to the plane of the two-dimensional ring formed by the transceiver elements. Therefore, the array of transceiver elements 110A, 110B, ..., 110H is movable relative to the medium placed in the imaging region 120.

[0062] Figure 1A is a simplified diagram showing only eight transceiver elements 110A, 110B, ..., 110H. In reality, an ultrasound computed tomography system consists of hundreds or thousands of transceiver elements. For example, a single ring array may consist of 512 or 2048 transceiver elements.

[0063] Each acoustic lens 108A, 108B, ..., 108H is positioned in front of each transceiver element 110A, 110B, ..., 110H to focus the waves onto the imaging plane and receive only waves from the imaging plane.

[0064] Each transceiver element 110A, 110B, ..., 110H is configured to both transmit and receive ultrasonic signals. Therefore, each transceiver element functions as both a transmitting and receiving element.

[0065] In other embodiments, a pair of transmitting elements is used to transmit an ultrasonic signal, and a pair of receiving elements is used to receive the ultrasonic signal. The receiving elements are different from the transmitting elements. In some such embodiments, the receiving elements may have different characteristics from the transmitting elements, such as being of a different size.

[0066] The transceiver elements 110A, 110B, ..., 110H are sometimes also called ultrasonic elements, array elements, or transducer elements. Each transceiver element is configured to transmit multiple programmable frequencies. Each frequency is transmitted with a very narrow bandwidth, e.g., near zero bandwidth. Each transmitted frequency may be called a single tone or single-tone continuous wave signal or single-frequency continuous wave signal.

[0067] Each transceiver element 110A, 110B, ..., 110H consists of its respective transducer 111A, 111B, ..., 111H and its respective receiving circuit 112A, 112B, ..., 112H (not shown in Figure 1A). The receiving circuits 112A, 112B, ..., 112H consist of a low-noise amplifier 113A, 113B, ..., 113H, a mixer stage 114A, 114B, ..., 114H and 115A, 115B, ..., 115H, and a low-pass filter 116A, 116 It consists of 116H and 117A, 117B, ..., 117H, and analog-to-digital converters 118A, 118B, ..., 118H and 119A, 119B, ..., 119H (not shown in Figure 1A). The components of the receiving circuit 112E for transceiver element 110E are shown in Figure 2. Other transceiver elements 110A, 110B...110H consist of corresponding receiving circuits 112A, 112B...112H having the same components as those illustrated for transceiver circuit 110E.

[0068] The ultrasound computed tomography apparatus 100 further comprises a signal generator 130. The signal generator 130 is configured to generate a sequence of monophonic continuous wave signals. Each monophonic continuous wave signal comprises a sinusoidal continuous wave signal having a predetermined frequency, as will be further described below. The signal generator 130 steps through a plurality of different predetermined frequencies in sequence and generates each monophonic continuous wave signal for each predetermined frequency using a predetermined voltage, phase, and transmission time.

[0069] The ultrasonic computed tomography apparatus 100 further comprises a controller 140 configured to control the signal generator 130 and to control which of the transceiver elements 110A, 110B, ..., 110H is transmitting or receiving at a predetermined time. The controller 140 may consist of one or more processors. In this embodiment, the controller 140 stores a predetermined set of frequencies, voltages, phases, and transmission times used to generate a sequence of single-tone continuous wave signals. In other embodiments, the predetermined frequencies, voltages, phases, and / or transmission times may be stored in any suitable data store.

[0070] Depending on the circumstances, the frequency, voltage, phase, and / or transmission time may be selected by the user.

[0071] In this embodiment, the system 100 is configured to operate in clinical mode and research mode. In clinical mode, the frequency, voltage, phase, and transmission time are stored and may not be modified by a clinical user, such as a radiologist. For example, a clinical user may not be permitted to change the voltage level for safety reasons. For example, different imaging protocols may require different voltage levels, and the voltage level of any given imaging protocol may not be editable by a clinical user.

[0072] In research mode, the research user is permitted to fully program the system, which may include the selection of frequency and / or voltage amplitude and / or phase and / or transmission time. Furthermore, the research user may be permitted to program other parameters of the system 100, such as the analog gain of the low-noise amplifiers 113A, 113B, ..., 113H. The single-tone transmission / dwell time may also be selected. In other embodiments, any suitable mode can be used, such as clinical mode and / or research mode and / or screening mode and / or follow-up mode and / or high-frequency mode. The full programmability of the system parameters may allow for the use of multiple modes or imaging protocols.

[0073] In some embodiments, different modes use different frequencies. For example, a higher frequency mode may be used to provide higher resolution. In some embodiments, different modes use different numbers of frequency steps; for example, a research mode may acquire more frequency steps than a clinical mode.

[0074] The ultrasound computed tomography system 100 further comprises a processing resource 150 configured to perform computed tomography processing. The processing resource may consist of one or more processors.

[0075] The controller 140 and processing resources 150 can form part of one or more computing devices, such as a personal computer, server, or workstation. One or more computing devices may include one or more central processing units (CPUs) and / or graphical processing units (GPUs). For example, the controller 140 may be implemented on a CPU by a computer program having computer-readable instructions executable to perform the method of the embodiment, and / or at least part of the controller 140 may be implemented as one or more ASICs (application-specific integrated circuits) or FPGAs (field-programmable gate arrays). The processing resources 150 may be implemented by the CPU and / or GPU by a computer program having computer-readable instructions executable to perform the method of the embodiment. In some embodiments, the receiving circuits 112A, 112B, ..., 112H and / or at least part of the other electronic components of system 100 may be implemented as one or more ASICs (application-specific integrated circuits) or FPGAs (field-programmable gate arrays).

[0076] During use, the medium is placed within the imaging area 120. In this embodiment, the medium is the breast of the subject being imaged. The two-dimensional ring-shaped array of transceiver elements 110A, 110B...110H is positioned around the breast at a first position. In this embodiment, the first position is the position closest to the subject's sternum, and the ring-shaped array scans away from the subject's chest. In other embodiments, the first position is the position furthest from the chest, for example, aligned with the nipple, and the ring-shaped array scans toward the chest.

[0077] Each transceiver element 110A, 110B, ..., 110H in the two-dimensional ring array is driven sequentially as described below to transmit a series of monophonic continuous wave signals. After propagating through the medium, each monophonic continuous wave signal is recorded by another transceiver element 110A, 110B, ..., 110H. The signals are recorded in parallel by multiple elements.

[0078] The transmission by the first transceiver element 110A is described below. Note that corresponding transmissions are also performed by each of the other transceiver elements 110B, 110C, ..., 110H. Once all transceiver elements 110A, 110B, ..., 110H have transmitted, the array is rearranged, and the transmission by each transceiver element 110A, 110B, ..., 110H is repeated. Rearrangement and transmission are repeated until the entire breast is imaged.

[0079] Figure 1B shows a ring array 170 comprising multiple transceiver elements 110 arranged around a breast 180. The dotted line 172 indicates different positions of the ring array. The direction of movement of the ring array is indicated as z in Figure 1B.

[0080] Figure 2 shows the transmit and receive paths of a stepped frequency architecture according to an embodiment. Figure 2 shows an example in which transmission is performed using a first transceiver element Tx and reception is performed using a second different transceiver element Rx located on the opposite side of the array from the first transceiver element Tx.

[0081] The controller 140 instructs the signal generator 130 to generate a series of single-tone continuous wave signals. Each single-tone continuous wave signal has a fixed frequency f within the bandwidth of the transducer 111A. k This is a continuous transmission of pure tones. Each monophonic continuous wave signal has a predetermined voltage, phase, and transmission time. In this embodiment, the controller 140 transmits at a predetermined frequency f kThe system stores the associated voltage, phase, and transmission time sets. The signal generator 130 sequentially steps through a set of different predetermined frequencies. For each predetermined frequency, the signal generator generates a single-tone continuous wave signal using the associated voltage, phase, and transmission time. The predetermined frequencies may be pre-programmed into the system. In some situations, the predetermined frequencies may be selected by the user.

[0082] In the case of transmission by the first transceiver element 110A, the controller 140 instructs the first transceiver element 110A to operate in transmit mode. The control device 140 first provides the signal generator 130 with a first single-tone continuous wave signal having a first frequency for a first transmission time. The control device 140 then provides a second single-tone continuous wave signal having a second frequency for a second transmission time after the transmission of the first single-tone continuous wave signal has finished. The control device 140 then provides any further single-tone continuous wave signals one by one, each of which has its own frequency. The control device 140 instructs the first transceiver element 112A to provide a series of single-tone continuous wave signals by providing each single-tone continuous wave signal for its respective transmission time. The transmission time is sometimes also called the residence time. Each single-tone continuous wave signal may have a different transmission time. The controller 140 instructs the remaining non-transmitting transceiver elements 110B, 110C, ..., 110H to operate in receive mode. For each monophonic continuous wave signal, the remaining transceiver elements 110B, 110C, ..., 110H receive the monophonic continuous wave signal. Each monophonic continuous wave signal has its own fixed frequency.

[0083] The signal generator 130 sequentially supplies each of a series of single-tone continuous wave signals to the first transceiver element 110A for transmission. For each single-tone continuous wave signal, the transducer 111A of the first transceiver 110A converts the single-tone continuous wave signal from an electrical signal to an ultrasonic signal and transmits it to the medium 160. The ultrasonic signal may also be called an acoustic wave.

[0084] The following description details the transmission, reception, and processing of one of a series of monophonic continuous wave signals. A monophonic continuous wave signal has a fixed frequency f k It has the same transmission, reception, and processing as each monophonic continuous wave signal transmitted from the transmitter.

[0085] A monophonic continuous wave signal is a sine wave waveform sin(2πf k It can be written as a continuous transmission having t). The transducer 111A converts a monophonic continuous wave signal from an electrical signal to an ultrasonic signal.

[0086] Figure 3 shows the frequency f k This is a voltage-versus-time plot used to generate a single-tone continuous wave signal with the peak-to-peak voltage V. pp (k) This is expressed as follows. Since the duty cycle is 1, the RF peak power and average power are equal. In the embodiment shown in Figure 3, the voltage waveform is a square wave. In some embodiments, the voltage waveform has the shape of a mathematical sine function rather than a square wave.

[0087] An ultrasonic signal is transmitted from the first transceiver element 110A, passes through the acoustic lens 108A (not shown in Figure 2) into the medium 160, and at least a portion of the ultrasonic signal is transmitted through the medium 160 for detection by one or more of the remaining non-transmitting transceiver elements 110B, ..., 110H. As will be described in more detail below, the ultrasonic signal received by one or more of the remaining non-transmitting transceiver elements 110B, ..., 110H may be used to reconstruct an image of the medium. A portion of the ultrasonic signal may be diffracted. A portion of the ultrasonic signal may be reflected.

[0088] The first transceiver element 110A may be located on the first side of the medium 160. One or more of the remaining non-transmitting transceiver elements 110B, ..., 110H used to detect ultrasonic signals may be located on the second side of the medium. The second side of the medium is different from the first side of the medium. The second side of the medium may be located approximately opposite to the first side of the medium. For example, the second side of the medium may be optionally located laterally on the opposite side from the first side of the medium. One or more of the remaining non-transmitting transceiver elements 110B, ..., 110H used to detect ultrasonic signals may be located on the opposite side of the first transmitting transceiver element 110A, and may optionally be located laterally relative to the first transmitting transceiver element 110A.

[0089] Each of the receiving transceiver elements 110B, ..., 110H receives its respective received ultrasonic signal as a result of transmission by transceiver element 110A. Figure 2 shows an example of reception by the components of transceiver element 110E facing directly towards the transmitting transceiver element 110A.

[0090] In the simplified embodiment shown in Figure 1A, when a stationary element is transmitting, all receiving elements 110B, 110C, ..., 110H receive in parallel. In practical embodiments, the system may include hundreds or thousands of receiving elements. In some such embodiments, a subset of receiving elements receives in parallel, followed by yet another subset of receiving elements receiving in parallel. In some embodiments, the receiving circuit is shared among the elements. For example, a single receiving circuit may be used to receive signals from two or more transducers. The number of channels used for receiving may be less than the number of elements, and multiplexing may be used.

[0091] Returning again to the example of FIG. 2, transducer 111E receives an ultrasonic signal that has passed through medium 160 via an acoustic lens 108E (not shown in FIG. 2). Transducer 111E converts the received ultrasonic signal into a received electrical signal that is an analog signal. The analog signal recorded by transducer 111E is input to the receiving circuit 112E of transceiver element 110E. The input to receiving circuit 112E is a sine wave form having the same frequency as the transmission signal.

[0092] The receiving circuit is composed of a low-noise amplifier (LNA) 113E, mixer stages 114E, 115E, low-pass filters 116E, 117E, and analog-to-digital converters 118E, 119E.

[0093] The received electrical signal is amplified by LNA 113A to obtain an amplified signal.

[0094] The amplified signal is passed from LNA 113E to mixer stages 114E, 115E, where the amplified signal is mixed with a single-frequency mixer signal in mixer stages 114E, 115E. The single-frequency mixer signal includes a cosine wave form and a sine wave form having the same frequency as the transmission signal. This is an example of a homodyne architecture in which the received signal is mixed with a sine wave form and a cosine wave form having the same frequency as the transmission signal.

[0095] Mixing with the cosine wave form in mixer stage 114E gives a in-phase signal (I). Mixing with the sine wave form in mixer stage 115E gives a quadrature signal (Q).

[0096] The result of each mixing operation is a signal having a single frequency that is twice the transmission frequency and a DC component. A sine wave signal of frequency f in mixed with a sine wave signal of frequency f mixer generates a signal of frequency f in ±f mixer In this case, since the received signal of frequency f k is mixed with a mixer signal of the same frequency f k , f k ±fk =2*f k , becomes 0 (DC). The mixing results are shown as I and Q in Figure 2.

[0097] In subsequent processing, only the DC component is used. Therefore, the signal obtained from the mix is ​​filtered through low-pass stages 116E and 117E. The common-mode signal is passed from the mixer stage 114E to the low-pass filter 116E, which removes the high-frequency component. The quadrature signal is passed from the mixer stage 115E to the low-pass filter 117E, which removes the high-frequency component. In the embodiment shown in Figure 2, the low-pass stages 116E and 117E have a cutoff frequency of 100 kHz. In other embodiments, any suitable cutoff frequency can be used. After low-pass filtering, the SFCW architecture collects only the two DC terms, for example, by sampling and retaining their values.

[0098] Except for initial transients, the filtered signal exhibits no temporal dynamics. In the embodiment shown in Figure 2, the filtered signal is sampled at a very low rate using ADC118E, 119E. For example, sampling rates of 128 kHz, 256 kHz, or 512 kHz may be used. The filtered signal is sampled at a high bit depth, for example, 24 bits or more. Low sampling rates are possible because the filtered signal is substantially constant with respect to time. In some situations, using a high bit depth or high bit resolution may result in higher contrast in the resulting medical image.

[0099] The filtered I signal from the low-pass filter 116E is passed to the ADC 118E, which samples the filtered I signal at a low rate and high bit depth. The ADC 118E outputs a signal Re, which is a DC value. The output signal Re is the frequency f of the transmitted signal. k It has a single value in . The output signal Re is the real part of the frequency spectrum of the signal at the receiver input.

[0100] Figure 2 shows a plot of the spectrum (real numbers) of frequency 200. Plot 200 represents the frequency f, which is the DC value Re output by the ADC118E. k Including point 201 in [location].

[0101] The filtered Q signal from the low-pass filter 117E is passed to the ADC 119E, which samples the filtered Q signal at a low rate and high bit depth. The ADC 119E outputs a signal Im, which is a DC value. The output signal Im is the frequency f of the transmitted signal. k It has a single value in . The output signal Im is the imaginary part of the frequency spectrum of the signal at the receiver's input.

[0102] Figure 2 shows a plot of the spectrum (Imag) against frequency. This plot 200 represents the frequency f, which is the DC value Im output by the ADC119E. k Including point 211 in [location].

[0103] After digitization, the two DC values ​​Re and Im are the frequency spectrum (f) of the signal at the receiver input. k This represents the real and imaginary parts of (in the given context). The two DC values ​​Re and Im are passed to processing resource 150.

[0104] Figure 4 shows several plots 220, 230, 240, and 250. Plots 220 and 230 in Figure 4 show demodulated signals I and Q plotted in arbitrary units versus time. In Figure 4, it should be understood that t=0μs corresponds to the arrival time of the signal at the receiver input. For a fixed frequency (750kHz in Figure 4) transmit tone, signals I and Q at the mixer output include a frequency twice that of the transmit tone and a DC term.

[0105] Plots 220 and 230 show the filtering results as dashed lines. Except for the initial transients caused by the filtering operation, the filtered signal is a constant signal.

[0106] Plots 240 and 250 show the amplitude spectra of I and Q plotted against frequency in arbitrary units, with the received I or Q signal and a filtered version I. filt or Q filt This shows the DC term (zero frequency). The vertical line marked DC corresponds to the DC term. Vertical line F OutMixer The vertical line FLP corresponds to the non-zero frequency at the mixer output, and the vertical line FLP corresponds to the cutoff frequency of the low-pass stage.

[0107] The explanation in Figure 2 above illustrates the processing of a single-tone continuous wave signal. In a sequence of single-tone continuous wave signals, the DC values ​​Re and Im are obtained for each single-tone continuous wave signal. Therefore, the processing resource 150 receives each pair of Re and Im values ​​from the receiving circuit 112E for each predetermined frequency used to generate the sequence of single-tone continuous wave signals. This corresponds to one complex digital sample for each combination of receiver, transmitter, and frequency. Each complex digital sample is sampled, for example, 24 bits.

[0108] In this embodiment, the sequence of monotone continuous wave signals comprises only one monotone continuous wave signal per frequency. In other embodiments, the frequencies may be repeated such that the sequence of monotone continuous wave signals consists of one or more monotone continuous wave signals per frequency. In such embodiments, the overall Re and Im values ​​for a given frequency may be obtained by combining the Re and Im values ​​for each single-tone continuous wave signal in the sequence of single-tone continuous wave signals transmitted at that frequency, for example, by averaging. As an example, if the sequence of monotone continuous wave signals consists of three monotone continuous wave signals at a given frequency, the combined Re value can be obtained by averaging the three Re values, and the combined Im value can be obtained by averaging the three Im values.

[0109] The corresponding processing is performed by the receiving circuits 112B, 112C, 112D, 112F, 112G, and 112H of the other receiving elements 110B, 110C, 110D, 110F, 110G, and 110H, and the Re and Im values ​​for each frequency are passed to the processing resource 150 by each receiving circuit 112B, 112C, 112d, 112F, 112G, and 112H.

[0110] The controller 140 instructs the signal generator 130 to sequentially transmit a sequence of single-tone continuous wave signals to each of the transceiver elements 110A, 110B, ..., 110H. When a transceiver element is transmitting, each of the other transceiver elements receives the signal and processes it to obtain a pair of Re and Im values ​​for each single-tone continuous wave signal.

[0111] When each of the transceiver elements 110A, 110B, ..., 110H transmits, the controller 140 instructs the ring-shaped array of transceiver elements 110A, 110B, ..., 110H to move by a predetermined increment in a direction perpendicular to the plane of the ring-shaped array. Depending on the situation, for example when operating in research mode, the increment may be set by the user. The controller 140 then instructs another set of signal transmission and reception by each of the transceiver elements, as described above.

[0112] The controller 140 repeats the incremental movement of the ring-shaped array and the transmission and reception of signals until, for example, the entire breast is imaged, or until a predetermined number of increments is reached, or until a limit is reached.

[0113] As already mentioned with reference to Figures 2 and 3, the transmitting circuit generates a sinusoidal waveform known as a square waveform or single tone. The single tone is transmitted over a programmable transmission time T. k Peak-to-peak voltage V Tx,pp (k) , frequency f kand phase (the latter being zero in Figure 2). The transmission time may be 400 μs. In other embodiments, different transmission times may be used, such as 200 μs or 600 μs. The peak-to-peak voltage may be 100 mV. In other embodiments, different peak-to-peak voltages, such as 1 V or 10 V, may be used. The frequency ranges from 100 kHz to 1 MHz in 50 kHz steps. In other embodiments, different frequency steps, such as 5 kHz or 100 kHz, may be used. In other embodiments, the frequency ranges from 100 kHz to 1 MHz, but the frequency steps may be of non-uniform size. The phase of a single tone may be zero. In other embodiments, the phase of a single tone may be different, such as 90 degrees or 180 degrees.

[0114] The previous parameters (transmission time, peak-to-peak voltage, frequency, and phase) are all independently programmable and may depend on the position along the scanning direction. For example, to compensate for increased signal attenuation due to larger breast volume, a higher peak-to-peak voltage may be employed in breast regions closer to the chest, combined with a longer transmission time.

[0115] As already mentioned above, in a single-ring array configuration, single tones of different frequencies are transmitted sequentially by single array elements before single tones of different frequencies are transmitted sequentially by the next single array element.

[0116] Next, referring to Figures 1A to 4, we will describe an ultrasonic computed tomography system that includes an array of 512 transceiver elements arranged in a two-dimensional ring shape around an imaging area where array element #256 is positioned opposite array element #1, similar to the one described above. Then, referring to Figures 5 to 12, we will describe the transmission and reception of a single tone.

[0117] A single tone is transmitted to a specific array element, for example, element #1. While element #1 is transmitting, all other elements receive in parallel. All frequencies within a predetermined range are transmitted sequentially. When element #1 has finished transmitting, a single tone is transmitted to array element #2. Frequencies within a predetermined range are transmitted sequentially. All array elements transmit in turn. Only one tone is transmitted at a time. The top plot in Figure 5 shows that element #1 is transmitting at f k =150kHz transmission time T k This shows an example of transmitting a single tone of =400μs.

[0118] Each array element has the same receiver chain as shown in Figure 2. The first stage is a low-noise amplifier (LNA), followed by two mixers, two low-pass filters, and two ADCs.

[0119] After transmission, as it propagates through the medium, the analog signal at the LNA input becomes a single tone, time-shifted at the same frequency as the transmitted signal. Figure 6 shows such a signal from array element #256 when array element #1 is transmitting. The arrival time may depend on the speed of sound in the inserted tissue / medium and the known distance between the Tx / Rx pairs.

[0120] The LNA can be applied with a gain of 20 dB (10 dB in one embodiment, 40 dB in another, -10 dB in yet another, etc.). The gain of the LNA may depend on the receiver, frequency, and scanning position. The result of LNA amplification is a time-shifted single tone with a larger peak-to-peak voltage amplitude at the same frequency as the transmitted tone.

[0121] As shown in Figure 2, the received and amplified single tone is mixed with cosine and sine waveforms (shown in the middle and bottom plots of Figure 5, respectively) at the same frequency as the transmitted single tone. As shown in Figure 5, these waveforms coincide temporally with the start of the single tone transmission.

[0122] In one embodiment, the same oscillator (not shown in Figure 2) drives both the transmitting and receiving circuits. In this case, a sinusoidal waveform is generated and passes through the first power divider. A portion of the waveform is sent to the transmitting circuit. Other portions pass through multiple power dividers and phase shifter stages (not shown in Figure 2), and these waveforms are sent to the I and Q branches of the mixer stage of each receiving element.

[0123] Since transmission is continuous, all array elements except the one transmitting will receive, for example, while array element #1 is transmitting, array elements #2 through #512 will receive.

[0124] The peak-to-peak voltage of the mixer waveform may have a value of 1 mV (100 μV in one embodiment, 10 mV in another embodiment, etc.). These peak-to-peak values ​​may depend on the transmission frequency, receiving element, and scanning position.

[0125] Mixer waveform peak-to-peak voltage, V Rx,pp (I;k) and V Rx,pp (Q;k) Even with the same receiving element, it can be different: for example, V Rx,pp (I;k) 1mV, V Rx,pp (Q;k) This can be as low as 5mV.

[0126] The mixer stage combines two sinusoidal mixer waveforms, shown in the middle and bottom plots of Figure 5, with a time-shifted LNA-amplified single tone, shown in Figure 6. The result of this operation is a transmit frequency f k Twice the output frequency f OutMixer This is a time-shifted sinusoidal waveform with a (generally) non-zero DC component. Element #1 has a frequency f k Continuing with the example of transmitting at =150kHz, the frequency of the signal at the mixer output is f in both the I and Q branches. OutMixer =2*f kThis results in 300kHz. Figure 7 shows examples of the I and Q waveforms for array element #256. Both waveforms show a non-zero DC offset (also known as DC bias) that can be observed in both the time domain and the frequency domain. Please note that the time t=0μs in Figure 7 corresponds to the start time of transmission, and the arrival time of array element #256 in Figure 7 is around t=150μs.

[0127] The receiver chain samples the DC components of both the I and Q branches at every receiving element. This is achieved by low-pass filtering the I and Q waveforms and digitizing the steady-state values ​​of these low-pass filtered analog signals.

[0128] As shown in Figure 7, the low-pass filter may have a cutoff frequency of 20 kHz. In other embodiments, the low-pass filter may have a different cutoff frequency, for example, 5 kHz or 100 kHz. The cutoff frequency may depend on the frequency of the transmitted single tone, the receiving element, and the scanning position.

[0129] Low-pass filtered signal I filt and Q filt The signal reaches its respective steady-state value after arriving at any receiving array element. For example, Figure 8 shows the low-pass filtered signal I from array elements #128 and #256 when array element #1 is transmitting. filt and Q filt This shows the trend. Two steady-state values ​​for each array element are digitized by the ADC, which is the final stage of the receiver chain.

[0130] The ADC sampling frequency is 128 kHz. In other embodiments, the ADC may have a different sampling frequency, such as 256 kHz or 512 kHz. The ADC resolution is 24 bits. In other embodiments, the ADC may have a different resolution, for example, 14 bits, 16 bits or 32 bits. The ADC sampling frequency may be programmable and may depend on the frequency of the transmitted single tone.

[0131] As shown in Figure 9, the ADC can sample from the start of transmission to the entire transmission time. The FPGA driving the ADC sums the digital samples stored in the ADC memory buffer. As a result of this calculation, for each Tx / Rx pair, a predetermined frequency f k Two real digital samples (ADC) I and ADC Q ) is obtained. These two samples are complex digital samples d(Rx, Tx; f k )=ADC I +ADC Q It is coupled to the actual ADC. I This is the result of digitization along the I branch, and the imaginary part ADC Q This is the result of digitization along the Q branch (Figures 2 and 9). This complex digital sample is transferred to a PC / GPU for further processing. In this embodiment, the integration time at all receiving elements is 400 μs, which is equal to the 400 μs transmission time shown in Figure 5. Refer to Figures 5-9 for the actual digital sample ADC described above. I and ADC Q Please refer to Figures 2-4 to understand that this is equivalent to the digital samples Re and Im described above.

[0132] In another embodiment shown in Figure 10, the ADC starts sampling at a specific arrival time until the end of the transmission time. The arrival time may be programmable and may depend on the receiving element. The arrival time can be easily estimated by assuming a known distance between Tx / Rx pairs and a known value for the speed of sound in the medium (for soft tissue, a reference value of 1540 m / s may be used). For example, in the case of a single ring array with 512 elements, when element #1 is transmitting, the element to the side (e.g., element #128) may start sampling around 100 μs after the start of transmission, and the element on the opposite side (e.g., element #256) may start sampling around 150 μs after the start of transmission. As described above, the ADC samples stored in the corresponding memory buffers are summed by the FPGA driving the ADC. The result of this operation is a complex digital sample for each Tx / Rx pair (given frequency). This complex digital sample is transferred to a PC / GPU for further processing. In this embodiment, the integration time may depend on the transmission frequency, receiving element, and scanning position. In Figure 10, array element #128 integrates for 100 μs (from 200 μs after transmission to 300 μs after transmission), while array element #256 integrates for 200 μs (from 200 μs after transmission to 400 μs after transmission). This method may yield more accurate samples because digital samples acquired during the transition from zero to steady state are discarded.

[0133] Longer integration times may result in a higher SNR, which may be particularly beneficial at higher frequencies (because higher frequency sound waves are attenuated more) and thicker / larger breast regions (because sound wave signals are attenuated more). Figure 11 shows that for a given frequency and a fixed Tx / Rx pair, longer integration times result in a higher SNR. Here, SNR is defined as the signal level (complex sample d(Rx, Tx; f) k Defined as the height in the frequency domain between the absolute value of (measured as the absolute value of) and the noise floor.

[0134] The transmission time must be the same length as the desired maximum integral time. The integral time depends on the frequency of the transmitted tone and the receiving element. The transmission time also depends on the frequency of the transmitted single tone and its peak-to-peak voltage.

[0135] All array elements transmit and receive, and all combinations are recorded. Since transmission is continuous, all array elements other than the transmitting element receive. Therefore, sample d(Rx=Tx, Tx;f k ) will not be collected. In the shape of a single ring array of 512 elements, samples d(Rx=Tx, Tx;f k It is convenient to store a single frequency dataset with a complex sample size of 512(Rx) × 512(Tx) digitized at ADC resolution, assuming that the values ​​are filled with zeros. This sample d(Rx=Tx, Tx;f k ) is not used in image formation algorithms anyway, as tomography imaging utilizes transmitted data. f If multiple frequencies are transmitted and recorded, the dataset will be 512(Rx) × 512(Tx) × N f This results in a complex sample size of [number], which is then digitized with ADC resolution.

[0136] When all combinations of Tx, Rx, and frequency are recorded at a single scanning position, the ring array moves vertically along the z-direction. The steps of elevation along the z-direction are programmable and may be 5 mm. In other embodiments, different steps such as 3 mm or 10 mm can be used.

[0137] N scans If is the number of scanning positions along the elevation angle, then a single breast dataset is 512(Rx) × 512(Tx) × N, digitized at ADC resolution. f ×N scans It has the size of a composite sample.

[0138] Figure 12 shows the complex sample d(Rx, Tx, Tx; f) measured by the side receiver (Rx128) and the opposite receiver (Rx256) when array element #1 is transmitting. k The absolute value of ) is shown. Figure 12 shows the frequencies transmitted at 100kHz to 800kHz (50kHz steps), 900kHz, and 1MHz. The data uses the same integration time and shows the peak-to-peak voltage V at all frequencies. Tx,pp (k) The data may be obtained by transmitting the same value. As shown in Figure 12, due to wave propagation and attenuation, the signal level at the opposing receiver may be lower than the signal level at the side receiver. Since tomography is enabled by transmitted data (at the side or opposing receiver), it is important to increase the signal level (and therefore the SNR) at the receiving elements involved in tomographic imaging by having a potentially uniform signal level. In the SFCW-USCT architecture, the SNR can be improved by simply adjusting the integration time of the receiving array elements to a longer value.

[0139] From the above discussion, a person skilled in the art will understand that the integral time of each receiving element is less than or equal to the transmission time of the transmitting element.

[0140] Processing resource 150 is d(Rx, Tx, Tx; f k Using all of the values, in this embodiment, a calculated transmission tomography calculation is performed to image the medium, which is the breast. In this embodiment, for each position of the ring array in the direction perpendicular to the plane of the array, the processing resource uses the d(Rx, Tx, Tx;f) obtained at that position. k Using the values, a computational transmission tomography calculation is performed to obtain a two-dimensional image for that location. Next, the processing resources combine the two-dimensional images of all locations to obtain a three-dimensional volume. In other embodiments, any suitable method of computed tomography calculation can be used.

[0141] Computerized tomography calculations are performed in the frequency domain.

[0142] The process of mapping ADC raw data to an image in Cartesian space is called image reconstruction or image formation. In medical imaging, the displayed image shows one, more, or a combination of physical and chemical quantities that contribute to the image's contrast. In ultrasound imaging, contrast is determined by the acoustic impedance of the medium. Acoustic impedance is the product of two physical parameters: mass density and sound velocity. Tissues can have different mass densities and sound velocity values, and the sound velocity value, in particular, shows a wide spread due to differences in mass density. Therefore, reconstructing images that show the sound velocity of different tissues can be useful for clinical diagnosis.

[0143] In one embodiment, image reconstruction is performed by an iterative algorithm that solves a numerical optimization problem. The latter is defined by a cost function C(θ).

number

[0144] This cost function measures the difference (e.g., the l² norm) between measured data and composite data generated according to a certain numerical rule. The vector θ is an unknown quantity (what should become the image). Solving the numerical optimization problem is equivalent to minimizing the cost function, that is, estimating the vectors θ=θ1, θ2, ... for which the cost function is minimized / locally small. In the case of a cost function that depends on a single unknown quantity θ1, it can be written as follows:

number

[0145] Here, the hatted variable θ^ represents the best estimate of the unknown variable θ1. This estimate can be achieved, for example, by known iterative techniques based on a simple gradient descent algorithm. This is the initial estimate of the unknown variable, θ1^ (0) Let's assume that this is available. By iterating through the initial values ​​along the gradient descent direction, a better estimate can be obtained.

number

[0146] Here, ∇ θ1 (kk) θ1^ is the numerical gradient of the cost function with respect to the unknown variable θ1 at kk iterations. The parameter α is the step size. Therefore, if initial estimation is available and the gradient of the cost function can be calculated, an estimated value θ1^ can be obtained. The step size can be defined by the user. The estimated θ1^ is displayed as a 2D image.

[0147] In one embodiment, image reconstruction is performed in the frequency domain using an iterative technique based on a physical model (model-based image reconstruction). These assume a physical model and iteratively minimize the difference between the measured data and the synthesized data generated through a pre-selected model. The physical model is often encoded by a partial differential equation (PDE) that describes known underlying physics. The cost function C(θ) can be written as follows:

number

[0148] The aforementioned sum is over all possible combinations of transmitting and receiving elements and over all acquired discrete frequencies. In this invention, d meas (Rx, Tx; f k The term ) is the complex sample d(Rx, Tx; f) mentioned above. k ) matches. Term d syn (Rx, Tx; f k ) is frequency f k This represents the composite data generated by the numerical model for a given combination of Tx / Rx array elements. The variable θ is a general vector θ1, θ2, ..., and each θ iθ is a parameter that describes a physical quantity. When the composite data is generated by a physical / parametric model described by a PDE, the numerical optimization problem becomes an example of a PDE-constrained optimization problem. In fact, in this case, the composite data must satisfy the constraints given by the relevant PDE, which can be written as M(θ)=0 (where M represents the model and the vector θ represents the physical parameters of this model). It is well known that the propagation of sound waves in soft tissue is well described in the time domain by the acoustic wave equation and in the frequency domain by the Helmholtz equation. The latter is a single-frequency (f) PDE.

number

[0149] Here, p(x→;f) is the pressure measured at position x→, c(x→) is the (generally) spatially varying speed of sound in the medium, and s(x→;f) is the sound source excitation (in the frequency domain) at position (x→) and frequency f (measured in Hz). ∇ 2 This is the standard Laplacian operator, where mass density is ignored. In this case, the physical model M(θ)=0 is described by the Helmholtz equation, and the model parameter is a scalar quantity θ=θ1=c (speed of sound in m / s). In this framework, after discretizing the Helmholtz equation on a numerical grid, the physical pressure p(x→;f) is given by a given pair of Tx / Rx array elements and a given frequency f=f k , and composite data d for a given estimate of the basic speed of sound syn (Rx, Tx; f k This coincides with θ=c). The iterative estimation of the speed of sound can be written as follows:

number

[0150] As mentioned above, c^ (0) This initial estimation, along with rules for calculating the numerical gradient with respect to the speed of sound, is necessary. The following describes an approach to multi-bandwidth reconstruction.

[0151] Continuing with the example of the 512-element ring array, consider the case of the phantom sample shown in Figure 13, where data is recorded at the following 13 discrete frequencies: 100kHz, 200kHz, 250kHz, 300kHz, 350kHz, 400kHz, 450kHz, 500kHz, 600kHz, 700kHz, 800kHz, 900kHz, and 1MHz. The data is generated in 3D on a grid with a radius of 110mm, a depth of focus of 55mm, a beam width of 1cm, 1MHz, a size of 1200x1200x68 pixels, and a pixel size of 0.2mm. A single array element is depicted as a rectangular shape with a focusing acoustic lens on it. The phantom sample contains multiple roughly spherical inclusions ranging in size from 5mm to 10mm. Inclusions with irregular boundaries may represent malignant tumors (spinous tumors), while those with regular boundaries may represent benign tumors or cysts. The inclusions are generated by varying the speed of sound, resulting in a different contrast against the background texture. Finally, texture was added to mimic the actual medium. Specifically, random values ​​were added at the pixel level to mimic the actual medium.

[0152] The acquired frequencies are grouped into five frequency bandwidths. Each frequency bandwidth is used to estimate the speed of sound in more detail with a smaller pixel size. Specifically, the frequencies are grouped as follows: [100kHz, 200kHz, 250kHz - 1mm], [300kHz, 350kHz - 0.8mm], [400kHz, 450kHz - 0.6mm], [500kHz, 600kHz, 700kHz - 0.4mm], [800kHz, 900kHz, 1000kHz - 0.32mm].

[0153] To illustrate the basic blocks of the image processing algorithm, we will detail the steps for reconstructing images with frequency bandwidths of 100kHz, 200kHz, and 250kHz. The measured dataset for these frequency bandwidths is a complex vector of size 512(Rx)×512(Tx)×3. We consider a constant initial estimation, also known as a flat initial estimation (constant initial velocity model). This is a 2D image of size 280×280 pixels with a pixel size of 1mm. The initial value of the sound velocity is 1500m / s (speed of sound in water). We solve the Helmholtz equation for the three discrete frequencies and compute the corresponding composite data. The cost function is evaluated by taking the difference between the measured and composite data. The gradient is calculated and the sound velocity is updated. This process is repeated the desired number of times. These steps are summarized below: I. Selection of frequency bandwidth (e.g., 100kHz, 200kHz, 250kHz) II. Select the pixel size (e.g., 1mm) III. Select the size of the image / numerical grid (e.g., 280 x 280 pixels) IV. Select initial estimated values ​​(for example, fill in with data values ​​for a constant 280x280 image at 1500 m / s). V. Start the iterative loop. VI. Using the current estimate of the speed of sound (the shape of Tx / Rx is known), solve the Helmholtz equation for all elements. VII. Extract the pressure in the receiver (the shape of Rx is known). VIII. Calculate the residuals between the synthesized data and the measured data. IX. Solve the addjoint problem and calculate the gradient. X. Breaks the record for the speed of sound. XI. Repeat steps V to X the desired number of times.

[0154] After 50 repetitions, an example of a reconstructed velocity image with a bandwidth of 100-200-250 kHz is shown in Figure 14. For a frequency bandwidth of 300-350 kHz, a finer grid of 300 x 300 pixels with a pixel size of 0.8 mm yields a better estimate of the velocity image. The previous estimate of the velocity image is first interpolated from a 280 x 280 (1 mm) grid to a finer 300 x 300 (0.8 mm) grid, and the same procedure is repeated 50 times. The final estimate is saved and shown in Figure 14. The principle of multibandwidth reconstruction is simple: the final estimate of the velocity image becomes the initial estimate for reconstruction on the finer grid (after interpolation). For example, the velocity image is interpolated from a 300 x 300 (0.8 mm) grid to a finer 468 x 468 (0.6 mm) grid, and the image is reconstructed with a bandwidth of 400-450 kHz. The results of the multibandwidth reconstruction are shown in Figure 14. To avoid inverse scrying, the images are reconstructed up to a maximum frequency of 1 MHz, but with a larger pixel size (0.32 mm) compared to the actual pixel size (0.2 mm) of the true numerical phantom sample. All images are displayed on the same scale of 1450–1580 m / s. As expected from first principles (higher frequency = higher resolution = higher image quality), we can see that the image quality improves from low to high frequencies. The final image shows a degradation in image quality compared to the true numerical phantom. This is due to the fact that the ring array is a two-dimensional system, while wave propagation is a three-dimensional physical phenomenon. Out-of-plane scattering can degrade image quality in any 2D system (regardless of how the transmitter and receiver are implemented). In particular, the image quality of the xy imaging plane may plateau at the frequency bandwidth corresponding to the (minimum) beamwidth of the device. In fact, the nominal beamwidth (1 cm in this example) is reached at 1 MHz, and further improvement of the speed of sound in the xy imaging plane may not be achievable. Finally, Figure 14 shows the graph obtained when the cost function is repeated 50 times for each frequency bandwidth, for a total of 250 times. Figure 14 clearly shows a decreasing behavior within each frequency bandwidth.This suggests that the numerical optimization problem has been solved, and that the SFCW-USCT architecture can reliably produce images of improved quality by progressively inverting for higher frequencies. In the absence of noise, all inclusions were detected, albeit with varying degrees of accuracy, as would be expected from general considerations of contrast, size, boundary changes, and 2D reconstruction from 3D data.

[0155] The reconstruction method described above is highly flexible, allowing for modifications to reconstruction parameters within a given frequency bandwidth. Specifically, the frequency per bandwidth can be selected (minimum 1 frequency per bandwidth), the maximum frequency, the grid size and pixel size determined by the numerical scheme used to solve the Helmholtz equation, the step size and iteration count can be selected (numerical optimization algorithms can be similarly selected, e.g., gradient descent, conjugate gradient), and a subset of transmitters and receivers can be selected. All of these parameters are independently programmable and may depend on the scan location and clinical task. Interpolation steps with smaller pixel sizes can be achieved with any interpolation algorithm.

[0156] Referring to the aforementioned example of a planar ring array, an image is formed at each scanning position. 3D body parts (e.g., breasts) are reconstructed by superimposing 2D images onto each other. These 2D images / slice are separated by a distance equal to the elevation sampling step.

[0157] The image reconstruction algorithm is performed on a dataset with fixed scan locations (single-slice reconstruction). It is also possible to reconstruct multiple scan locations in parallel and independently.

[0158] As already described above, the tomographic image can be determined using a standard iterative technique within the general framework of model-based image reconstruction in the frequency domain. The calculated tomographic image calculation consists of an iterative algorithm based on the numerical optimization of a cost function. The cost function C measures the difference between the data d(f k ;Rx, Tx) recorded at a single frequency and the synthetic data M(f k ;θ1, ..., θ n ;Tx,Rx).

Number

[0159] Here, d(f k ;Rx, Tx) is equal to d(Rx, Tx;f k ) defined above. That is, d(f k ;Rx, Tx) is a complex digital sample in which the real and imaginary parts are respectively equal to the digitized values of the I and Q analog signals after low-pass filtering and optionally integration for a given pair of transceiver elements Tx, Rx, as described above with reference to FIGS. 2 to 10. Here, “Real”(d) is denoted as Re above, and “Imag”(d) is denoted as Im above. The sum is performed over all possible combinations of transceivers using the (implicit) assumption that tomography is enabled by the transmission data. M(f k ;θ1, ..., θ n ;Tx,Rx) should be understood as d syn (Rx, Tx;f k ;θ) described above. M(f k ;θ) is a parametric model described by a partial differential equation (pde), and θ1, ..., θ n are the physical variables of the model (e.g., the speed of sound). The estimation of the physical parameter θ^ is equivalent to the minimization of the cost function of the pde-constrained optimization problem.

Number

[0160] Synthetic data can be obtained using one or more physical models (depending on the medical application). Examples include frequency-domain acoustic wave equations (Helmholtz equations) with or without viscoelastic terms, and frequency-domain elastic wave equations with or without viscoelastic terms. These methods allow for the use of physical parameters θ. j (For example, compressive and / or shear sound velocity, acoustic and / or elastic damping, etc.) can be estimated. The cost function can be minimized by iteratively updating the estimates of the physical quantities of interest until convergence is reached. If the gradient is known, the physical quantities of interest can be iteratively updated using a simple gradient descent algorithm with a fixed step size α.

number

[0161] Here, ∇ θj (iter+1) = ∂C / ∂θ j This is the physical quantity of interest θ at each iteration step. j This is the gradient of the cost function with respect to θ. The numerical gradient can be calculated using the well-known adjoint state method. Alternatively, other optimization methods, such as line search or gradient descent using conjugate gradients, can also be employed. The initial physical model of the optimization problem (initial estimated value θ) j (iter=0)The model may be a flat / constant model. In some embodiments, the first iteration is performed in the low-frequency domain, and the updated model is fed into the iterative reconstruction algorithm in the high-frequency domain. The SNR of the low-frequency and high-frequency data can be ensured by adjusting the amplitude and integration time of the single-frequency tone (and its steps). The low-frequency and high-frequency iterative reconstruction algorithms may differ, employing different physical models and different numerical methods to solve the wave equation (for example, the partial differential equation describing a particular physical model may be solved by the standard finite difference method in one domain and by the spectral method or finite element method in another). A regularization term may be added to the cost function, for example, to account for experimental noise and / or to improve convergence and / or smoothness of the final image.

[0162] The processing resource 150 can output a 2D or 3D image to, for example, a display screen (not shown). The output may include any appropriate image rendering, such as rendering 2D slices or rendering a 3D image using any appropriate projection.

[0163] Referring to Figures 1A to 14, the method described above can provide a safe and operator-independent method for ultrasound imaging of the breast for the diagnosis of breast cancer. This ultrasound imaging method is suitable for use in screening, such as in mass breast cancer screenings.

[0164] Transmission is continuous, and data is collected directly in the frequency domain, not the time domain. Time domain data is not collected. Only frequency domain data is collected after digitization. Frequency domain data acquisition is combined with a frequency domain image reconstruction method.

[0165] By sequentially transmitting at different frequencies (programmable steps) and collecting the two DC values ​​from each transmission, the desired bandwidth can be combined.

[0166] Unlike conventional handheld ultrasound probes that utilize only pulsed echo data, SFCW-USCT systems use SFCW transmitted data. SFCW-USCT systems may also utilize SFCW diffraction data, sometimes referred to as lateral data.

[0167] Single-tone control: The amplitude, frequency, phase, transmission time, and integration time of a single tone can be controlled independently. Step frequency continuous wave ultrasound computed tomography architectures can be designed to have any desired SNR for each frequency bin / step.

[0168] In the embodiments described above, the array of ultrasonic elements is a two-dimensional array that moves in a direction perpendicular to the plane of the two-dimensional array. In some embodiments, the array of ultrasonic elements is a three-dimensional array instead of a two-dimensional array. The three-dimensional array of ultrasonic elements may consist of a stack of multiple two-dimensional arrays, for example, multiple two-dimensional ring arrays. In some such embodiments, the three-dimensional array is configured to image the entire medium, for example, the entire breast, without moving. In other embodiments, the three-dimensional array moves as described above.

[0169] The systems in Figures 1A and 1B consist of a planar ring array in which all elements function as transceivers and are equipped with acoustic lenses that focus the signals onto the imaging plane. Data is acquired by a single circular ring. The breast is scanned vertically for two-dimensional reconstruction. Reconstruction is performed slice by slice. Physical quantity θ j It is estimated to be a 3D volume formed by stacking 2D images.

[0170] That is, the physical quantity θ j It is estimated to be a complete 3D volume formed by stacking a 3D volume at the same location as the device and thin 3D volumes.

[0171] In the embodiments shown in Figures 1A and 1B, all transducer elements have the same focusing characteristics, such as the same thickness, height, and width. Furthermore, the same acoustic lens is used for each element, and the focal point on the imaging plane is the same at all positions within the z-axis. One ring scans the breast vertically and sequentially collects data in the elevation angle / z direction.

[0172] In other embodiments, the stepped frequency continuous wave signal and processing can be used in any suitable shape. In some embodiments, the acoustic lens can be omitted. The elements may constitute at least part of at least one circular ring, or they may constitute any other suitable array shape.

[0173] The SFCW-USCT system has been described above, and those skilled in the art will understand the following: The transmitted signal is a single tone of a programmable frequency; the received signal at the input of the LNA is a single tone time-shifted at the same frequency as the transmitted signal; the amplified analog signal passes through a homodyne receiver; the analog signal at the input of the ADC is a DC signal; the data is sampled in the frequency domain; and the data is processed through a model-based iterative technique that reconstructs physical parameters (e.g., speed of sound).

[0174] The SFCW-USCT systems and methods described above are purely illustrative examples, and it should be understood that modifications to the SFCW-USCT systems and methods can be made without departing from the scope of the invention as defined in accordance with the appended claims. For example, although the SFCW-USCT system 100 has been described as having 8, 512, or 2048 array elements, the SFCW-USCT system may have more or fewer array elements. Regardless of the number of array elements, at any given time, one of the array elements can transmit and another array element can receive.

[0175] Figure 16 shows an alternative configuration of the SFCW-USCT system in which ultrasonic elements 402 are arranged on the inner surface of the bowl 400. The array of elements 402 in Figure 16 may be described as a bowl array, surface array, or matrix array. Elements 204 are uniformly distributed inside the bowl and directed toward the breast 180. In the embodiment of Figure 16, the breast 180 is sensed in three dimensions and reconstructed using three-dimensional reconstruction. No mechanical movement of the array is used.

[0176] In the embodiment shown in Figure 16, all transducer elements have the same focusing characteristics, such as height and width. In other embodiments, different transducer elements may have different focusing characteristics. In further embodiments, the elements can be arranged inside any convex surface having any suitable shape.

[0177] Figure 17 shows a further alternative configuration of the SFCW-USCT system, in which transducer elements 412, 422, and 432 are arranged in three circular rings 410, 420, and 430. All transducer elements 412, 422, and 432 transmit and receive regardless of the ring to which they belong. When a transducer element in ring 410 transmits, transducer elements 412, 422, and 432 in the other three rings 410, 420, and 430 receive.

[0178] In some embodiments, the three circular rings 410, 420, and 430 may have different radii. For example, to follow the curvature of the breast, the bottom ring 430 (located further away from the sternum) may have a smaller radius, the middle ring 420 may have an intermediate radius, and the top ring 410 may have a larger radius. In the embodiment of Figure 17, the breast 180 is scanned vertically for 3D reconstruction. The three rings move as a single mechanical entity. Data is collected sequentially in the elevation / z direction, and the 3D volume is reconstructed by stacking a thin 3D volume (illustrated as a dashed rectangle / field of view 440) located at the same position as the three rings.

[0179] Other elements may use a different number of rings.

[0180] In the embodiment shown in Figure 17, all transducer elements belonging to the same ring have the same focusing characteristics. In some embodiments, ultrasonic elements in different rings may have different acoustic characteristics. For example, elements in different rings may have different focal points, beam widths, frequency bandwidths, and / or sizes of a single transducer element. Transducer elements in different rings can transmit at different frequencies. In some embodiments, different transducer spacings are used on different rings.

[0181] Figure 18 shows an embodiment in which transducers are arranged in two opposing groups of transducer arrays, each having multiple rows. In the embodiment shown in Figure 18, the first group, Group 1, includes a first two-dimensional array A1 having three rows of transducer elements 500 on the first side of the breast 180, and a second two-dimensional array B1 having three rows of transducer elements 502 on the second opposite side of the breast 180. The transducer elements of array A1 or array B1 can lie on a flat surface or on a curved or bowl-shaped surface. Arrays A1 and B1 rotate synchronously around the breast 180 so that they are always positioned on opposing sides of the breast 180.

[0182] Group 2, a second group, has a similar shape but different acoustic properties. The second group includes a third two-dimensional array A2 having three rows of transducer elements 504 on the third side of the breast 180, and a fourth two-dimensional array B2 having three rows of transducer elements 506 on the fourth side of the breast 180, opposite to the third side. The transducer elements of array A2 or array B2 may lie on a flat surface, or on a curved or bowl-shaped surface. Arrays A2 and B2 rotate synchronously around the breast 180 so that they are always positioned on opposite sides of the breast 180.

[0183] In the embodiment shown in Figure 18, each set of arrays is arranged such that signals transmitted from the first array are transmitted through the medium and received by the second array, which is positioned directly opposite the first array. In other embodiments, when the first array is positioned on the first side of the medium, the second array may be positioned at any suitable position around the medium, and it does not have to be directly opposite the first array. For example, the second array may receive signals laterally from the first array.

[0184] Elements in Group 1 may have a first characteristic, and elements in Group 2 may have a second, different characteristic. For example, elements in Group 1 may have different focusing characteristics than elements in Group 2. Transducer elements in Group 1 and Group 2 may transmit and receive in different frequency bandwidths. Group 2 may have different acoustic characteristics (e.g., height, width).

[0185] By using different acoustic characteristics, Group 1 yields data in the lower frequency range (e.g., 200kHz to 300kHz), while Group 2 yields data in the higher frequency range (e.g., 2MHz to 3MHz). Both the low-frequency and high-frequency data can be used for image reconstruction.

[0186] Mechanical rotation is used to completely enclose the breast. Vertical scanning is used for 3D reconstruction (a stack of thin 3D volumes). In some embodiments, the distance between array A1 and array B1, and the distance between array A2 and array B2, can be controlled electronically and / or mechanically while scanning the breast in the elevation direction. At a fixed elevation position, the data is used to reconstruct a thin volume CO located together with two groups (illustrated as dashed rectangles / field of view 510).

[0187] In other embodiments, any suitable number of arrays can be used.

[0188] In the SFCW-USCT system configuration described above with reference to Figures 1A, 1B, 16, 17, and 18, all transducer elements function as both transmitters and receivers (with the only rule being that a transmitting element does not receive while transmitting). In other embodiments, separate transmitters and receivers may be used.

[0189] Furthermore, the system can operate in different modes, such as screening mode, high-resolution mode, study mode, and follow-up treatment mode. The final geometry design may include an imaging subsystem that transmits at a low frequency (having an SNR that allows for accurate estimation of the initial velocity model, which is important in iterative reconstruction) and another imaging subsystem that transmits at a high frequency (allowing for the recovery of high-resolution detail by further iterations for potentially clinically dependent tasks). For example, consider the geometry in Figure 18. In some embodiments, group 1 may be used to perform low-frequency imaging for screening, for example. A combination of group 1 and group 2 can be used to obtain high-resolution data for study or follow-up.

[0190] In some embodiments, the transmitter and receiver have different focusing characteristics. In some embodiments, the transmitter can be a group of elements. The group of elements can transmit together for plane wave transmission. In some embodiments, the receiver is a group of elements.

[0191] In the alternative SFCW-USCT system configuration shown in Figure 19, two rings may be present. These two rings may have different numbers of elements, different radii, and different acoustic focusing characteristics. In particular, the array elements of the low-frequency (LF) ring transmit and receive in a bandwidth of 100 kHz to 500 kHz, while the array elements of the high-frequency (HF) ring transmit and receive in a bandwidth of 500 kHz to 1000 kHz. In this configuration, there can be two independent transmit channels. That is, one element of the LF ring and one element of the HF ring can transmit simultaneously, for example, at 100 kHz and 600 kHz, respectively. During reception, the elements of the LF ring record the signal transmitted at 100 kHz, and the elements of the HF ring record the signal transmitted at 600 kHz. In terms of imaging parameters, these two rings operate as two completely independent units. The two rings are separated by a certain distance and scan the breast with mechanical movement. Slices at the same elevation are jointly registered by the two rings, and low-frequency and high-frequency images can be obtained by the two rings. LF-ring elements are used, for example, for low-frequency imaging for screening. Combinations of LF-ring and HF-ring elements can be used to obtain higher-resolution data for research and follow-up.

[0192] Please understand that the breasts or other body parts may be immersed in a water bath (a tank filled with water). The water bath represents a binding medium for ultrasound propagation.

[0193] In another embodiment, the sensing / imaging device can be implemented as a wearable patch. One drawback of wearable patches is that the power emitted is low because the drive circuit is located in the patch itself. Therefore, wearable patches for pulsed transmission may not be able to record transmitted data due to their low power (typically a few volts). On the other hand, continuous transmission with peak-to-peak voltages of several hundred mV and sufficiently long integration times may be able to record transmitted data. Wearable patches may enable 2D imaging. For example, Figure 20 shows a linear array positioned on a circular stripe attached to the breast for 2D imaging. Alternatively, wearable patches can enable 3D imaging. For example, Figure 21 shows a matrix array positioned on foil wrapped around the breast for 3D imaging. A potential drawback of some of the shapes described above is that the elevation focus is fixed. This is because the acoustic lens has a fixed depth of focus and a fixed beamwidth at a given frequency. Because there is considerable variation in breast size, a device with a fixed elevation focus at a given radius may only be suitable for some breast sizes. In the case of wearable patches, wearable stripes and / or foils with different radii, different acoustic focusing, variable array elements, etc., can be designed, enabling more personalized imaging. Wearable patches can be applied directly to the skin with or without the application of an acoustic gel as a coupling medium. In this embodiment, the wearable patch is intended to be a disposable medical device.

[0194] In the SFCW-USCT system and method described above, the same oscillator (not explicitly shown in Figure 2) drives both the transmitting and receiving circuits.

[0195] In another embodiment of the SFCW-USCT, two different synchronous oscillators can be used: Osc1 and Osc2. A single master clock is used to synchronize oscillators Osc1 and Osc2. Osc1 drives the transmitting circuit. A sine wave waveform is generated and sent to the transmitting circuit. Osc2 can drive the receiving circuit. A sine wave waveform is generated and passes through multiple power dividers and phase shifter stages. These waveforms are sent to the I and Q branches of the mixer stage of each receiving element. In this embodiment, the transmitting and receiving circuits are isolated.

[0196] In further embodiments of the SFCW-USCT, multiple synchronous oscillators can be used, where k is the number of receiving elements. All oscillators Osc1 and Osc k A single master clock is used to synchronize them. Osc1 drives the transmitting circuit. A sine wave waveform is generated and sent to the transmitting circuit. Each Osc k Each drives a different receiving circuit. k This generates a sinusoidal waveform, which passes through a power divider and a phase shifter stage before being sent to the I and Q branches of the mixer stage of a corresponding receiving element. In this embodiment, the transmitting circuit, receiving circuit, and mixer waveform are completely isolated.

[0197] In another embodiment, multibandwidth image reconstruction can be complemented by crosstalk-free multibandwidth reconstruction using phase encoding. In conventional multibandwidth reconstruction, the Helmholtz equation (or an equivalent PDE describing some physical model) is solved for each transmitting element, under the assumption that only one array element transmits at a time. In the phase-encoding embodiment, the Helmholtz equation is solved under the assumption that all array elements transmit together (as far as the synthesized data is concerned). In this case, the gradient calculation is plagued by crosstalk noise. To eliminate or reduce this crosstalk noise, a phase-encoded vector is generated at each iteration step. A phase-encoded vector is generated for each frequency and for each array element. The synthesized data is generated by solving the Helmholtz equation where all array elements are transmitted together, and each is encoded by a phase-encoded term. The measured data is encoded by the same phase-encoded term, and the residual and gradient at the receiver are calculated. The speed of sound is updated, and a new iteration can start with a different encoding vector. In this embodiment, the phase-encoded term is a pseudo-random vector obtained by subtracting a pseudo-random uniform distribution in the interval [0, 2π]. For the phase-encoded principle to function (i.e., to reduce crosstalk noise), this pseudo-random vector must be generated at each iteration step. To ensure reproducibility, this pseudo-random vector must be generated in a deterministic manner. This can be achieved by fixing the seed of the pseudo-random number algorithm in a deterministic manner (e.g., equal to the number of iterations) and using a known pseudo-random number algorithm that generates a uniform number in the interval [0, 2π] with the selected seed. These steps are summarized below: I. Select a frequency bandwidth (e.g., 100kHz, 200kHz, 250kHz). II. Select the pixel size (e.g., 1mm). III. Select the size of the image / numerical grid (for example, 280 x 280 pixels). IV. Select an initial estimated value (for example, a fixed image of 280x280 pixels, filled with 1500). V. Start the iterative loop. VI. Fix the seed (e.g., equal to the number of iterations). Subtract a pseudo-random number from a pseudo-uniform random number distribution on [0, 2π]. Encode the source term with the pseudo-random phase term. Sum the encoded sources. VII. Solve the Helmholtz equation using the current estimated value of the speed of sound (the shape of Tx / Rx is known) and the encoded sound source. VIII. Extract the pressure at the receiver (the shape of Rx is known). IX. Encoded measurement data with the same pseudo-random phase term. Sum the encoded measurement data. X. Calculate the residual between the encoded synthetic data and the encoded measurement data. XI. Solve the adjoint problem to calculate the gradient. XII. Update the speed of sound. XIII. Repeat steps V to XII for the desired number of times.

[0198] In other words, after selecting the number of iterations, the number of array elements, and the discrete frequencies, the encoding vector can be considered as a lookup table that is generated only once, even before measuring the data. It is important to emphasize that in multi-band reconstruction without crosstalk by phase encoding, the method of acquiring data does not change. Single tones are sequentially transmitted element by element for each frequency. In the reconstruction process, the synthetic data mimics the situation where all array elements transmit at once and all array elements receive together. In the reconstruction process, the acquired data is encoded by summation (as if all array elements were transmitted and received and recorded at once). The phase encoding strategy may be employed to estimate a velocity model starting from a certain initial guess, and the standard multi-bandwidth reconstruction may further improve it.

[0199] In another embodiment, the physical model of the Helmholtz equation allows for a complex speed of sound. The propagation of sound waves with energy absorption is described by the same Helmholtz equation where the speed of sound is allowed to be complex, i.e., c = c real +ic imagIt is known that it can be described by the real part c of the speed of sound. real represents the speed at which a wave propagates through a medium, and the complex number part c imag This represents frequency attenuation. In this case, the cost function has two parameters, θ1 = c real and θ²=c imag It depends on the following. Both parameters can be updated within the same iteration step by simply replacing the gradient equation with the same equation that does not take the real part. In this case, the initial estimate must be a complex vector. Reconstructing frequency attenuation in addition to the real part of the speed of sound may be clinically relevant, because different tissues may have similar values ​​for the (real part) of the speed of sound (and therefore may not show sufficient image contrast) but different frequency attenuations (and therefore the imaginary part of the speed of sound may show image contrast). It should be emphasized that this distinction exists only in the physical model employed for reconstruction and does not require any change in how SFCW-USCT data are acquired. In this embodiment, two independent images are displayed (for each slice / ring position), one representing the real part of the speed of sound and the other representing the imaginary part of the speed of sound.

[0200] In another embodiment, a generalized Helmholtz equation can be employed. In this case, the speed of sound is compressed c (compressional) and shear c (shear) There are two concepts. The generalized Helmholtz equation can allow for frequency attenuation as described above. In other words, c (compressional) =c real (compressional) +ic imag (compressional) and c (shear) =c real (shear) +ic imag (shear) In this case, the cost function is θ1 = c real (compressional) θ²=c imag (compressional) θ3=c real (shear) θ4=c imag (shear)It depends on four physical parameters. In this case, the updates of the four parameters may be performed within the same iteration step, or, as is common, they may be performed alternately. A more accurate physical model needs to include both the compressed sound velocity and the shear sound velocity. In fact, when imaging body parts containing bone / hard tissue, approximating the sound wave described by a simple Helmholtz equation with only the compressed sound velocity may not yield sufficient accuracy because the discrepancy between the physical model and the data affects the resulting image. In this embodiment, four independent images are displayed (for each slice / ring position), the first representing the real part of the compressed sound velocity, the second the imaginary part of the compressed sound velocity, the third the real part of the shear sound velocity, and the fourth the imaginary part of the shear sound velocity.

[0201] The embodiments described above can be performed with or without phase encoding. This applies to both the Helmholtz equation (when the speed of sound is real or complex) and the generalized Helmholtz equation (when the speed of sound is real or complex for compression and shear).

[0202] In another embodiment, image reconstruction is achieved by reverse time shift in the frequency domain (RTM-FD). The image reconstructed in this way is complementary to the velocity image because it also contains information about the mass density of the underlying tissue. In other words, the RTM image provides additional information / contrast because it is related to the acoustic impedance of the medium. The RTM image can be thought of as a tomographic image equivalent to a standard B-mode reflectance map. In this embodiment, one image is displayed (for each slice / ring location).

[0203] For example, RTM images are displayed according to the speed of sound (real or imaginary, compressed or sheared). It should be reiterated that multiple image formation algorithms are possible with the same SFCW-USCT measurement data.

[0204] The SFCW-USCT system enables multiple imaging protocols. For screening protocols (optimized for sensitivity, i.e., detection of benign or malignant tumors), the imaging protocol can use only the low frequencies of a single ring array, or the LF ring of the two ring systems described above. The image in Figure 14 shows that all inclusions (as small as 5 mm) can be detected within a bandwidth of 100-450 kHz. Thus, the system only needs to transmit the relevant frequency steps from 100 kHz to 450 kHz, accelerating acquisition time, which can be a limiting factor in mass screening programs. Furthermore, reconstruction time up to 450 kHz is faster because solving the Helmholtz equation at higher frequencies is computationally intensive. For imaging protocols optimized for specificity (i.e., classifying already detected tumors as benign or malignant to exclude or rule out disease), the SFCW-USCT system can employ higher frequencies and bandwidths of 500 kHz to 1 MHz. This is also applicable for treatment planning and / or follow-up (e.g., during chemotherapy). For research purposes, imaging protocols can employ finer frequency steps and longer transmission and / or integration times (the latter not necessarily coinciding with clinical scan times).

[0205] In conclusion, the SFCW-USCT architecture is highly versatile, allowing for pre-planning of the imaging protocol. Additionally, as shown in Figure 15, women undergoing breast scans can have their breast circumference measured at three locations (nipple region, mid-region, and chest region) prior to the scan. These measurements can inform the appropriate selection of imaging parameters (transmission time, integration time at the receiver and integration time along the scanning direction, peak-to-peak voltage, and frequency step). Such selection of imaging parameters can be further adjusted for combinations of breast density, fat content, and risk of developing disease in women.

[0206] Processing resources can additionally or alternatively process Re and Im values ​​to obtain a conventional B-mode image. A conventional B-mode image I can be obtained using standard delayed-sum beamforming in the frequency domain, for example, by the following equation.

number

[0207] Here, |d(f k ;Tx, Rx)| 2 is, frequency f k This is the power spectrum of the data collected, i.e., d(f k ;Tx, Rx) are complex digital samples, and their real and imaginary parts are equal to the digitized values ​​of the above I and Q analog signals after low-pass filtering and optional integration for a given pair of transmitting and receiving elements Tx, Rx, respectively. )Real(d) is denoted as Re above, and Imag(d) is denoted as Im above. W Tx and W Rx τ is a weighting function applied to the apertures of the transmitter and receiver, respectively, to control the beam shape and side lobes. TxRx This collectively includes the time-of-flight delay from a specific transmitter Tx to a point target (pixel x), and the time-of-flight delay from the point target (pixel x) to a specific receiver Rx. The first sum in the above formula is for all step frequencies, and the second sum is for all transceiver combinations involved in pulsed echo imaging, i.e., combinations that utilize reflected data.

[0208] The embodiments described above are illustrated with reference to ultrasound imaging of breast tissue. In other embodiments, apparatus and / or methods similar to those described above can be used for medical imaging of any suitable anatomical structure of any human or animal subject, which may be diagnostic imaging for any suitable clinical application. The anatomical structure to be imaged does not have to be the breast. The anatomical structure to be imaged may consist of any body part that can be at least partially enclosed by one or more geometric structures, such as the geometric structures in Figure 1A, Figure 1B, or any of Figures 16 to 21. For example, the anatomical structure to be imaged may consist of the brain, legs, arms, chest, neck, or one or two testicles. Multiple apparatuses or apparatus shapes can be designed to scan multiple parts of the body.

[0209] The primary application described so far is breast imaging. Clinical applications to other body parts such as the brain, chest, neck, legs, arms, and testicles may also be relevant.

[0210] In further embodiments, systems, apparatus, and / or methods similar to those used above can be used for non-medical applications such as non-destructive testing, seismic exploration, and the oil and gas industry.

[0211] Each feature disclosed herein and (where appropriate) in the claims and drawings may be provided independently or in any suitable combination.

[0212] Comparison of the SFCW-USCT system with a virtual pulse system using direct RF sampling.

[0213] Here, we compare a system using step frequency continuous wave transmission as described above, for example with Figures 1A to 14, to a hypothetical pulse system using direct RF sampling as shown in Figure 22. This hypothetical pulse system assumes that 2048 independent transceiver elements are arranged in a two-dimensional ring array around the medium being imaged, with fixed elements transmitting and the remaining elements receiving in parallel. The hypothetical pulse system can transmit at a frequency of 3 MHz.

[0214] When using direct RF sampling, data must be collected at at least twice the Nyquist frequency to prevent aliasing. However, in actual equipment, due to noise, the sampling frequency is never equal to the Nyquist, and is usually at least four times the center frequency. Therefore, we assume that the system has the same number of ADCs as the number of independent channels sampling at 12 MHz, and the bit depth is 14 to 16 bits or less.

[0215] Typical drive voltage waveforms 320, 322, and 324 are shown in Figure 22. Various possible options for pulse drive voltage waveforms are shown, including unipolar drive waveform 320, bipolar drive waveform 322, and center frequency F c The drive waveform 324 includes one or more cycles. The voltage amplitude and pulse width must be large enough to cause the transducer element to resonate within its bandwidth. Alternatively, F c By driving the transducer for one or more cycles, the center frequency and bandwidth of the transmitted pulse can be controlled more precisely.

[0216] The ADC typically turns on immediately after transmission is complete. For the wave to travel through the breast and reach the opposing element, the drive peak-to-peak voltage value can be approximately 100Vpp or higher. This voltage, coupled with the system's pulsed nature, can limit the RF peak power absorbed by the transmitting circuit. Finally, transmission must be within the ISPTA (Spatial Peak Time Average Intensity) and other related acoustic output limits required by regulatory bodies.

[0217] A typical direct RF sampling architecture based on pulse transmission is shown in FIG. 22. F c The pulse transmission 300 of F is transmitted by the transducer 302 through the medium 304. The signal received by the transducer 306 is usually amplified by an LNA (low noise amplifier) 308. The LNA is one of the first stages of the analog front end of any receiver because the amplitude of the input signal is too small to be properly digitized. The signal is then band-passed by a band-pass filter (BPF) 310 with a bandwidth that depends on the bandwidth of the transducer and the drive voltage waveform. Finally, the signal is sampled by an analog-to-digital converter (ADC) 312, and the digital time series 314 is stored. Image reconstruction (e.g., speed map sound) can be performed in the time domain or the frequency domain after FFT.

[0218] When using the RF sampling method as described above, to collect data for the entire fixed-position ring array, N Tx transmissions are required, and since the system records all Tx / Rx combinations, N Tx *N Rx *N t results in N t actual digital samples, where N Tx is the number of digital samples saved per transmission. N Rx = N T = 2048, N

[0219]

[0219] = 2000 (assuming sufficient time is recorded for the wave to travel from one transceiver to the opposite transceiver), then for each position of one ring, it is 2048×2048×2000 (real samples of discrete time) × 16 bits ~ 16 GB. When collecting data for a breast in a range of about 10 cm from the chest with a spatial sampling of 1 mm, there are 100 independent scanning positions for one breast, and the total data volume is ~ 1.5 TB. In practice, data may also be collected from both breasts.When reconstructing breast images in the frequency domain, a limited set of frequencies, e.g., 10 frequencies, can be used. However, in RF sampling architectures, a large amount of time-domain data is collected before conversion to the frequency domain, so using a limited number of frequencies typically does not reduce hardware requirements. In contrast, when data is collected in the frequency domain, using fewer frequencies may reduce hardware demands.

[0220] The step frequency principle relies on the fundamental properties of the Fourier transform (analysis and synthesis). These properties are briefly explained below.

[0221] It is well known that the Discrete Fourier Transform (DFT) of a discrete-time domain signal x(n) of length N is given by the following equation:

number

[0222] The DFT is usually performed using a well-known algorithm for the FFT (Fast Fourier Transform). The result is a complex vector of length N^. This vector has several notable properties, in particular the real part of the FFT is the Nyquist frequency (F Nyq The FFT is symmetric around (=Fs / 2), so Re(X(2+j))=Re(X(Nfft-j)), and the imaginary part of the FFT is inversely symmetric, so Im(X(2+j))=-Im(X(Nfft-j)). Nfft is the number of points on which the FFT operation is performed. Here, we assume Nfft=N.

[0223] Fourier analysis refers to the study of the output of the DFT and the spectral content of the signal in the frequency domain. The inverse DFT (IDFT) is given by the following equation.

number

[0224] Each term in the previous sum is a complex exponent at a fixed frequency. So-called Fourier synthesis is equivalent to the previous sum, and by summing all the complex exponential terms at fixed frequencies, the time-domain signal can be completely reconstructed. Next, the signal is decomposed into Fourier components. The Fourier component at a fixed frequency k is defined as a discrete real-time domain signal.

number

number

[0225] Finally, we define a (pure) tone at a fixed frequency k as a sinusoidal waveform of the same frequency with frequency-dependent amplitude and zero phase ψ0=0.

number

[0226] For practical use, the phase value of a monophonic continuous wave signal can be any value and does not need to be zero.

[0227] The concept is shown in Figure 23. As shown in Plot 330, consider a time-domain signal with an RF pulse shape and a center frequency Fc = 750 kHz. A fixed-frequency Fourier component and a pure tone (B) with the same amplitude as the corresponding Fourier component at the same frequency are considered. k =a kThis is shown in plot 340. Also note the difference in positive / negative peak amplitudes between the RF pulse and the Fourier component / pure tone. The RF power and voltage amplitude required for single tone transmission can be much smaller than the RF power and voltage amplitude used for RF transmission. The original signal can be reconstructed by summing all the Fourier components (plot 332). The tone amplitude is the same as the Fourier component amplitude, but the phase is not added (coherently) to reconstruct the signal. Plots 334 (Fourier component) and 344 (pure tone) show the results of Fourier synthesis considering only a limited number of frequencies, specifically 10 frequencies around Fc. In this case, the reconstruction by summing only a subset of the Fourier components (plot 334) is inaccurate, and uncanceled oscillations are observed.

[0228] As shown in Figure 23, accurately reconstructing a time-domain profile requires collecting a large number of frequencies. Collecting a limited subset of frequencies and performing reconstruction in the time domain results in artifacts in image space. On the other hand, collecting only a few discrete frequencies within the bandwidth of interest and performing reconstruction in the frequency domain (e.g., reconstruction using the output of Figure 2 above) may yield an artifact-free image.

[0229] Figure 24 shows the Fourier analysis. The real and imaginary parts of the FFT of the RF pulse are shown around the center frequency (real part in plot 360, imaginary part in plot 362). The real and imaginary parts of the FFT of each Fourier component are consistent with the FFT values ​​of the RF pulse, as expected from theory. Plot 364 shows the amplitude spectrum.

[0230] For visualization purposes, the phase spectra of the Fourier components (shown with asterisks in plots 360 and 362) and the phase spectra of the timbre (shown with crosses in plots 360 and 362) only show values ​​corresponding to the fundamental frequency, while the amplitude spectra show the complete curves of both.

[0231] Since the pure tone has zero phase, its phase spectrum does not match that of the RF pulse. In particular, following the FFT coding convention (exp(-iθ)=cos(θ)-isin(θ)) and defining the pure tone as a zero-phase sine function, it is expected that after projection along the Fourier axis, the frequency spectrum of each tone will have a real part of zero and an imaginary part of negative. However, since amplitudes equal to those of the Fourier component and the pure tone were chosen, the amplitude spectra of the RF pulse, the Fourier component, and the pure tone coincide.

[0232] Turning back to the step-frequency continuous wave architecture, the desired bandwidth can be synthesized by sequentially transmitting at different frequencies (in programmable steps) and collecting two DC values ​​for each transmission. The result of this synthesis is shown in Figure 25.

[0233] center frequency F c Consider pulse transmission. After wave propagation, the signal is received and digitized according to a direct conversion architecture as shown in Figure 22. The phase and amplitude spectra of the digitized signal are plotted in 370, 372, and 374.

[0234] Next, F, which has the same amplitude as the Fourier component of the pulse signal. c Consider multiple consecutive transmissions at 10 different frequencies (pure tones) in the vicinity. The received signal of each transmission passes through the SCFW receiving path as shown in Figure 2, and two DC values ​​are stored for each transmission. The amplitude spectra of the pure tones (indicated by the crosshairs) and pulse transmission match as expected.

[0235] Figures 24 and 25 show the FFT spectra of the Fourier components when the amplitude and phase are determined by the Fourier transform of pulse transmission and transmitted, then digitized via the SCFW architecture. In this case, the phase spectra overlap, demonstrating that the SFCW architecture correctly extracts the frequency spectrum (amplitude and phase). The amplitude spectrum of the pure tone collected by the SFCW architecture matches the amplitude spectrum of an equivalent pulse-based transmission system obtained by direct RF sampling.

[0236] A set of values ​​at a fixed frequency (indicated by a cross) corresponds to the Re and Im values ​​in the ADC output of Figure 2.

[0237] Figure 25 shows the output of a step frequency continuous wave (SFCW) architecture. In this case, the real and imaginary parts of the spectrum can be directly extracted (two real samples in the discrete case). Therefore, plot 374 in Figure 25 shows only the dots corresponding to the timbre spectrum.

[0238] The stepped frequency continuous wave architecture (SFCW-USCT) for ultrasonic CT is mathematically equivalent to a pulse-based ultrasonic CT system with direct RF sampling, but the SFCW-USCT system offers superior performance (e.g., lower emission power and lower noise floor) and may be more versatile than a pulse-based ultrasonic CT system with direct RF sampling. For comparison, we assume the shape of a single ring array where all elements function as both transmitters and receivers.

[0239] In pulse-based ultrasound CT systems, all transmissions are driven by the same drive voltage waveform (single negative excitation, fixed center frequency, one or more cycles, similar). Peak-to-peak voltage V pp Pulsed It is fixed. Transmission is pulsed, meaning the drive voltage is not zero for a limited time. RF peak power and average power are different.

[0240] In SFCW-USCT, the drive voltage waveform is a sine wave, and its frequency can be stepped within a predetermined range. Voltage amplitude V pp SFCW(k) The phase ψ(k) is programmable and frequency-dependent. Transmission is continuous. Since the duty cycle is 1, the peak power and average power of the RF are equal.

[0241] In pulse-based ultrasound CT systems, once the transducer dimensions, material, acoustic lens, and drive voltage waveform are fixed, the pulse shape cannot be changed (for example, when measuring with a hydrophone underwater). In other words, the frequency content and relative frequency amplitude cannot be changed.

[0242] With SFCW-USCT, after selecting the transducer dimensions, material, and acoustic lens, you can program the frequency content and relative frequency amplitude.

[0243] In pulse-based ultrasound CT systems, as soon as transmission is complete, all elements (including the one that just finished transmitting) record the received signal. This can be achieved by having as many independent ADCs as there are physical transducer elements. The ADCs are turned on when transmission is complete.

[0244] In SFCW-USCT, all elements (except the transmitting element) record the received signal during transmission. This can be achieved by having the same number of independent ADCs as the number of physical transducer elements. The ADCs may be turned on from the start of transmission, as shown in Figure 9.

[0245] In pulse-based ultrasound CT systems, multiple pulses from the same transducer are possible within the range that increases the signal-to-noise ratio (SNR). In this case, the recorded time series of the first and second pulses are averaged, and the same applies to multiple pulses.

[0246] In SFCW-USCT, multiple consecutive transmissions (with the same amplitude, frequency, phase, and integration time) are possible from the same transducer to improve the signal-to-noise ratio (SNR). In this case, the two DC values ​​recorded in the first and second transmissions are averaged, and the same applies to multiple transmissions.

[0247] In SFCW-USCT, the frequency is stepped through all desired values. Amplitude, phase, and integration time can be varied.

[0248] In pulse-based ultrasound CT systems, after data is collected from the same transducer element (one or more pulses), another transceiver begins transmitting, collecting data in the same way as before, and the process continues until all Tx / Rx combinations are exhausted. The duration of each transmit / record (sometimes known as the PRF, pulse repetition frequency) is fixed and must be such that the wave reaches all elements in the array. The time between consecutive transmits is adjustable and must be such that echoes from the previous transmit attenuate below the receiver's noise floor.

[0249] In SFCW-USCT, data is collected from the same transducer element (one or more transmissions), and after all frequency steps have passed, another transceiver begins transmitting, collecting data as before, and the process continues until all Tx / Rx combinations have been explored (the rule is that all elements except the transmitting element receive). The temporal duration of each consecutive transmission is equal to the integral time, which is sometimes called the residence time. The integral time may be frequency-dependent, allowing the wave to reach all elements in the array. The time between consecutive transmissions is adjustable, allowing echoes from previous transmissions to attenuate below the receiver noise floor.

[0250] In the case of a single array, the device is then moved to another location along a direction perpendicular to the chest, and this process is repeated until all locations have been scanned.

[0251] In pulse-based ultrasound CT systems, data is digitized at (at least) four times the center frequency. For example, F c If =3MHz, the sampling frequency is F s = 12MHz. Even if frequency inversion is performed in the 1MHz range, low frequencies may be lost due to aliasing, so 4×F c Note that it is sampled as follows: N per receiver (in the most preferred scenario) sampled at 16 bits. t Each (real) time sample is sent.

[0252] In SFCW-USCT, since the analog signal after mixing and low-pass filtering is a pure DC signal, a low sampling rate ADC with high bit depth in the kHz domain, such as 24 bits, can be employed. The sampling frequency and bit depth of the ADC do not depend on the transmission frequency. Digitizing the DC values ​​of transmission frequencies of 1 MHz and 3 MHz can be achieved with the same ADC. For each continuous transmission at a fixed frequency, there are 2 (real) frequency samples per receiver, sampled at 24 bits. For each transmitter, there are 2 (real) frequency samples per receiver, sampled at 24 bits, with a frequency step count of N. fSteps It becomes twice that amount.

[0253] In pulse-based ultrasound CT systems, when evaluating the system's SNR, the signal level (shown as an asterisk in Figure 26) is basically the V of the drive voltage waveform. Tx,pp Pulsed The noise figure of the system (shown as a dashed line in Figure 26) is essentially determined by the entire bandwidth (MHz range) of the receiver. The dashed line connecting the asterisks in Figure 26 indicates that these data were obtained after the FFT of the digital time series.

[0254] In SFCW-USCT, when evaluating the system's SNR, the signal level (shown as a cross in Figure 26) is the V of the drive voltage waveform. Tx,pp SFCW(k)The noise figure of a system (shown as a dashed line in Figure 26) is essentially determined by the entire bandwidth (kHz domain) of the receiver, although this can vary with frequency. The SNR per frequency (i.e., the ratio in the frequency domain between the signal level and the noise floor at a particular frequency) can be selected by adjusting the amplitude of different tones at different frequencies and / or by adjusting the integral / dwell time of a single tone, and thus can be much better than the SNR of a pulse-based USCT system.

[0255] Figure 26 compares the SNR of a pulse-based system using direct RF sampling with the SNR achievable with the SFCW-USCT system. Three scenarios are shown. In the scenario shown on the left of Figure 26, the tone is transmitted with the same amplitude as the Fourier component of the equivalent RF pulse transmission. In the scenario shown in the center of Figure 26, the tone is transmitted with the same amplitude as the Fourier component (center frequency) of the equivalent RF pulse transmission. In the scenario shown on the right of Figure 26, the tone is transmitted with increasing amplitude while stepping through the frequency to accommodate, for example, tissue attenuation or transducer bandwidth. In all cases, SNR gain and data fidelity are ensured.

[0256] The SFCW-USCT architecture offers the flexibility to independently program single-tone peak-to-peak voltage levels for each frequency and to select different integration times for each receiver. Thus, the SNR of the SFCW-USCT architecture can be shaped to the desired form. In effect, the SFCW-USCT architecture may enable imaging at frequencies below 400 kHz, whereas pulse / RF devices have poor SNR performance below 400 kHz and may therefore be unable to image at frequencies below 400 kHz.

[0257] Using SFCW-USCT may increase the detection rate of inclusions. Another benefit is that with SFCW-USCT, the irregular edges of irregular inclusions (spinous processes of cancerous tumors in clinical settings) become more prominent, and the distinction between round and irregular inclusions becomes stronger. In summary, because SFCW-USCT allows for flexible adjustment of the SNR, it can improve sensitivity, specificity, and image quality compared to pulsed devices that perform direct RF sampling.

[0258] Compared to pulsed devices that perform direct RF sampling, the SFCW-USCT (step-frequency continuous-wave ultrasound computed tomography) architecture may allow for greater flexibility in data acquisition in the frequency domain. SFCW-USCT systems may have a higher SNR and higher data fidelity, lower storage requirements, and / or lower demands on the absorbed RF peak power of the transmitting circuit. In tomography settings, the SFCW architecture may improve SNR and data fidelity.

[0259] SFCW systems have lower peak-to-peak voltages than pulsed systems. SFCW systems may have lower absorbed RF power than pulsed systems.

[0260] For example, SFCW-USCT uses 24 bits or more, while pulse / RF uses 14 bits. This also means improved data fidelity.

[0261] SFCW systems may require less storage than pulsed systems, for example, several Gb per breast compared to 1 Tb for pulsed / RF systems.

[0262] The SFCW-USCT system is expected to have a much higher SNR, and because the frequency step and voltage amplitude can be programmed, in some situations the frequency required for inversion can be reduced to achieve image quality equivalent to or better than that of a pulsed USCT system. In some embodiments, a different geometry can be used than that of a pulsed USCT system. In some embodiments, fewer transmitters and / or receivers may be used than in a pulsed USCT system.

[0263] Single-tone control is provided. The amplitude, frequency, phase, transmission time, and integration time of a single tone can be controlled independently. A step-frequency continuous-wave ultrasound computed tomography system can be designed to have any desired SNR for each frequency bin / step. Such precise control of individual frequency bins, amplitudes, and steps is impossible with pulsed systems.

[0264] The described SFCW-USCT system and method can be used to obtain artifact-free images without reverse climb, starting from constant / flat initial estimations, in a bandwidth of 100 kHz to 350 kHz. In the case of breast tissue (whose sound velocity values ​​are widely known to be in the range of 1450 to 1580 m / s), the solution disclosed herein demonstrates that the SFCW-USCT architecture can already provide robust images at low frequencies.

[0265] In pulse-based ultrasound CT systems, when image reconstruction is performed in the frequency domain, it is necessary to take the FFT of all Tx / Rx combinations (after potential digital filtering operations on the raw data) and extract the discrete frequencies of interest. The frequency bin / resolution is determined by the sampling frequency of the ADC (which is fixed) and the number of digitized time samples N. t This is determined by (also fixed).

[0266] In SFCW-USCT, data is acquired in the frequency domain (with programmable frequency steps), so the system is naturally designed to perform inverse transforms in the frequency domain. No additional operations on the data are required.

[0267] Single-tone control is important. In pulse transmission, the pulse cannot be easily shaped, and the relative frequency content and SNR per frequency follow suit. c When transmitting several cycles at 400kHz, the power spectrum at 1MHz will be lower than the corresponding value at 400kHz. Resolution improves with higher frequencies, so the signal-to-noise ratio (SNR) increases with higher frequencies (making the computational cost worthwhile). When transmitting several cycles at 1MHz, the SNR at 400kHz worsens, which can affect the quality of the initial velocity model at lower frequencies and thus the reconstruction at higher frequencies.

[0268] The SFCW-USCT architecture can be designed to have any SNR per frequency bin / step, which is not possible in pulsed USCT systems, regardless of how the receiving path is implemented.

[0269] The primary application to date has been breast imaging. However, clinical applications to other body parts such as the brain, chest, neck, legs, arms, and testicles are also possible. The chest and brain, being larger in size and containing bone and hard tissue, can present greater acoustic mismatch and make imaging more difficult. Due to the increased size, body complexity, signal attenuation, and wave scattering, the peak-to-peak voltage of pulsed transmission must be 100V to ensure the signal level of the transmitted data is above the noise floor. ppIt can be much higher. This has many drawbacks. Firstly, medical transducers commonly used in diagnostic imaging may not be able to withstand such high pulse voltages. Secondly, very high pulse voltages may violate diagnostic imaging safety guidelines in terms of parameters such as mechanical index and ISPTA. Thirdly, because the noise floor of receivers sampling in the MHz range is generally higher than that of receivers sampling in the kHz range, the signal level of transmitted data may be captured with a very low number of bits. Taken together, pulse architectures that directly sample RF may be completely impractical for collecting transmitted data. Thanks to the flexibility of SNR due to the combination of peak-to-peak voltage, integration time, and frequency step, the SFCW-USCT architecture can achieve more penetration than conventional pulse architectures with direct RF sampling, and subsequently enable ultrasound tomography imaging of body parts that are inaccessible with pulse / RF schemes.

Claims

1. An ultrasound computed tomography method, The steps include transmitting a sequence of single-tone continuous wave signals through the medium using a transmitter located on the first side of the medium, The steps include: receiving a sequence of single-tone continuous wave signals transmitted through the medium by each of a plurality of receivers located on the opposite side of the medium; For each of the plurality of receivers, the steps include mixing each of the received sequences of single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signal to obtain the common-mode and quadrature values ​​of each of the single-tone continuous wave signals. The steps include processing the obtained in-phase and orthogonal values ​​by computed tomography to obtain an image of the medium, Includes, The ultrasonic computed tomography method is characterized in that the sequence of single-tone continuous wave signals includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second frequency different from the first frequency.

2. The ultrasound computed tomography method according to claim 1, characterized in that the medium includes breast tissue.

3. The ultrasonic computed tomography method according to claim 1 or 2, characterized in that the transmitter and receiver form part of an array of ultrasonic elements arranged around the imaging region in which the medium is placed.

4. The ultrasonic computed tomography method according to claim 3, characterized in that the array is circular or partially circular.

5. The steps include moving the transmitter and receiver in a direction perpendicular to the plane formed by the transmitter and receiver by a continuous increment of position relative to the medium, For each of the position increments, the step of performing further transmission of each of the sequences of the single-tone continuous wave signals, The steps include processing each of the in-phase and orthogonal values ​​obtained from the further transmission by computed tomography to obtain a two-dimensional image of the medium, The steps include: obtaining a three-dimensional volume by combining the aforementioned two-dimensional images; An ultrasonic computed tomography method according to any one of claims 1 to 4, characterized by including the following:

6. The ultrasound computed tomography method according to claim 3, characterized in that the array is a three-dimensional array comprising a plurality of two-dimensional arrays, wherein the two-dimensional arrays are optionally circular or partially circular, and further optionally the plurality of two-dimensional arrays each have different sizes.

7. The ultrasonic computed tomography method according to claim 6, characterized in that different two-dimensional array ultrasonic elements have different characteristics.

8. The ultrasonic computed tomography method according to claim 3, characterized in that the array is arranged on a convex surface, for example, on a bowl-shaped surface.

9. The array comprises at least one pair of opposing two-dimensional arrays, The ultrasonic computed tomography method further includes the step of rotating at least one pair of opposing two-dimensional arrays around the medium, The ultrasonic computed tomography method according to claim 3, wherein optionally, the at least one pair of opposing two-dimensional arrays comprises a first pair of opposing two-dimensional arrays equipped with ultrasonic elements having a first characteristic, and a second pair of opposing two-dimensional arrays equipped with ultrasonic elements having a second characteristic different from the first characteristic.

10. The first frequency is between 100 kHz and 20 MHz. The ultrasonic computed tomography method according to any one of claims 1 to 9, characterized in that the second frequency is between 100 kHz and 20 MHz.

11. The ultrasonic computed tomography method according to any one of claims 1 to 10, characterized in that the single-tone continuous wave signal includes a single-tone continuous wave signal having a pre-programmed series of frequencies.

12. The sequence of single-tone continuous wave signals includes two or more single-tone continuous wave signals having a first frequency. The ultrasound computed tomography method is The steps include: combining the common-mode values ​​of two or more single-tone continuous wave signals to obtain a combined common-mode value of the first frequency; The steps include: combining the orthogonal values ​​of two or more single-tone continuous wave signals to obtain the combined orthogonal value of the first frequency; The ultrasonic computed tomography method according to any one of claims 1 to 11, further comprising:

13. The ultrasound computed tomography method according to any one of claims 1 to 12, characterized in that the step of processing each of the in-phase values ​​and the orthogonal values ​​by computed tomography uses an iterative algorithm based on numerical optimization of the cost function.

14. The ultrasound computed tomography method according to any one of claims 1 to 13, characterized in that the step of processing each of the in-phase values ​​and the orthogonal values ​​by computed tomography includes the step of processing them by computed tomography with or without phase encoding.

15. An ultrasound computed tomography (CT) scanner, An imaging region configured to accommodate the medium to be imaged, An array of ultrasonic elements arranged around the imaging region, comprising a transmitter located on a first side of the imaging region and a plurality of receivers located on the opposite side of the imaging region, A receiving circuit for each of the aforementioned plurality of receivers, A processing circuit configured to perform computed tomography, Equipped with, The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area, the single-tone continuous wave signal comprising at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency, Each of the plurality of receivers is configured to receive the sequence of the single-tone continuous wave signal transmitted through the imaging area. Each receiving circuit of the plurality of receivers is configured to mix each of the sequences of received single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signal to obtain the in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The ultrasonic computed tomography apparatus is characterized in that the processing circuit is configured to process the obtained in-phase and orthogonal values ​​by computed tomography to obtain an image of the imaging region.

16. The array of ultrasonic elements comprises a plurality of transceivers, each configured to transmit and receive ultrasonic signals. The transmitter is the first transceiver of the plurality of transceivers, The ultrasonic computed tomography apparatus according to claim 15, characterized in that the receiver is a further transceiver of the plurality of transceivers.

17. The ultrasonic computed tomography apparatus according to claim 15 or 16, characterized in that the array is circular or partially circular.

18. The ultrasonic computed tomography apparatus according to claim 15 or 16, characterized in that the array is movable in a continuous increment of position perpendicular to the plane formed by the transmitter and the receiver with respect to the medium.

19. The ultrasound computed tomography apparatus according to any one of claims 15 to 18, characterized in that the array is a three-dimensional array comprising a plurality of two-dimensional arrays.

20. The ultrasonic computed tomography apparatus according to any one of claims 15 to 19, characterized in that the receiver comprises a mixer stage configured to mix each received single-tone continuous wave signal in the sequence of single-tone continuous wave signals with each single-frequency mixer signal having the same frequency as the single-tone continuous wave signal to obtain the in-phase value and the orthogonal value of each of the single-tone continuous wave signals.

21. The ultrasonic computed tomography apparatus according to claim 20, further comprising: a receiver, a low-noise amplifier configured to amplify the received signal; a pair of low-pass filters that remove frequencies exceeding a threshold frequency; and a pair of analog-to-digital converters that digitize the common-mode value and the quadrature value.

22. The ultrasonic computed tomography apparatus according to claim 21, characterized in that each of the pair of analog-to-digital converters has a bit depth of at least 24 bits.

23. The ultrasonic computed tomography apparatus according to claim 21 or 22, characterized in that the sampling rate of each of the pair of analog-to-digital converters is between 1 kHz and 1000 kHz.

24. The ultrasonic computed tomography apparatus according to any one of claims 15 to 22, further comprising a controller configured to determine and / or store for each of the single-tone continuous wave signals at least one of the following: each voltage, each phase, each transmission time, and each integration time, wherein optionally, at least one of the following is programmable and frequency-dependent.

25. The ultrasonic computed tomography apparatus according to any one of claims 15 to 24, further comprising a signal generator configured to generate a sequence of single-tone continuous wave signals.

26. An ultrasound computed tomography method, The transmitter transmits a sequence of single-tone continuous wave signals through a medium, The steps include: each of the multiple receivers receives a sequence of single-tone continuous wave signals transmitted through the medium; For each of the plurality of receivers, the steps include mixing each sequence of transmitted single-tone continuous wave signals with each single-frequency mixer signal having the same frequency as the single-tone continuous wave signal to obtain the common-mode and quadrature values ​​of each of the single-tone continuous wave signals. The steps include processing the obtained in-phase and orthogonal values ​​by computed tomography to obtain an image of the medium, Includes, The ultrasonic computed tomography method is characterized in that the sequence of single-tone continuous wave signals includes at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second frequency different from the first frequency.

27. The transmitter is positioned on the first side of the medium. The plurality of receivers are arranged on the second side of the medium. The second side of the medium differs from the first side of the medium, The second side of the medium is on the opposite side from the first side of the medium and optionally on the side of the first side of the medium. The transmitter is positioned on a first surface in space, and the plurality of receivers are positioned on a second surface in space, or The ultrasonic computed tomography method according to claim 26, characterized in that the transmitter is arranged on a first line in space and the plurality of receivers are arranged on a second line in space.

28. An ultrasound computed tomography (CT) scanner, An imaging region configured to accommodate the medium to be imaged, An array of ultrasonic elements arranged around the imaging area, comprising a transmitter and a plurality of receivers, A receiving circuit for each of the aforementioned plurality of receivers, A processing circuit configured to perform computed tomography, Equipped with, The transmitter is configured to transmit a single-tone continuous wave signal through the imaging area, the single-tone continuous wave signal comprising at least one single-tone continuous wave signal having a first frequency and at least one single-tone continuous wave signal having a second different frequency different from the first frequency, Each of the plurality of receivers is configured to receive the sequence of the single-tone continuous wave signal transmitted through the imaging area. Each receiving circuit of the plurality of receivers is configured to mix each of the sequences of received single-tone continuous wave signals with each of the single-frequency mixer signals having the same frequency as the single-tone continuous wave signal to obtain the in-phase and quadrature values ​​of each of the single-tone continuous wave signals. The ultrasonic computed tomography apparatus is characterized in that the processing circuit is configured to process the obtained in-phase and orthogonal values ​​by computed tomography to obtain an image of the imaging region.

29. The transmitter is positioned on the first side of the imaging area. The plurality of receivers are arranged on the second side of the imaging area. The second side of the imaging region differs from the first side of the imaging region, The second side of the imaging region is on the opposite side from the first side of the imaging region, and is optionally located to the side of the first side of the imaging region. The transmitter is positioned on a first surface in space, and the plurality of receivers are positioned on a second surface in space, or The ultrasonic computed tomography apparatus according to claim 28, characterized in that the transmitter is arranged on a first line in space and the plurality of receivers are arranged on a second line in space.

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