CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to United States Provisional Application No. 62/344,260, filed June 1, 2016, the contents of which are hereby incorporated by reference in its entirety.

NOTICE OF GOVERNMENT SUPPORT

This invention was made with government support from the National Institutes of Health under Grant Nos. R01-EB006042 and R01-HL114358. The

Government has certain rights in the invention.

BACKGROUND

In biological tissues, an increase in stiffness can be a marker of an abnormality, and indicative of an underlying disease or condition. For example, stiffer tissue can be indicative of various tissue diseases and tumors. Thus, ultrasound elastography can be used to assess mechanical properties of soft tissue and can be of interest in the detection of solid tumors or other diseases resulting in a change is tissue stiffness as compared to surrounding or healthy tissue.

Techniques such as quasi-static ultrasound elastography have been developed to assess the strain distribution in soft tissues in two dimensions using a quasi-static compression. The technique can be limited to strain estimation in two dimensional images. However, tumors can exhibit heterogeneous shapes and a single cross-sectional strain map can lead to an erroneous estimation of the tumor volume. Thus, a three-dimensional approach can be desirable to measure the volume with improved accuracy and to remove operator dependency. A high volume rate acquisition can also be desirable, for example to avoid signal decorrelation, to perform strain mapping in 3D with a freehand compression. Thus, there remains a need for improved techniques and systems for imaging of the elasticity in human bodies in three dimensions using simple freehand scanning.

SUMMARY

The disclosed subject matter provides methods and systems for 3D ultrasound quasi- static elastography.

In certain embodiments, an exemplary method for ultrasound

elastography includes emitting at least one non-focused wave on a target, obtaining

Radio Frequency (RF) signals of the non-focused wave, beamforming 3D volumes from the RF, calculating at least two 3D displacements by comparing each 3D volume to a reference volume, and integrating the 3D displacements to create a 3D cumulative axial strain volume.

In certain embodiments, the non-focused wave can be a plane wave. The non-focused wave can be emitted at a rate of about 100 volumes/sec.

In certain embodiments, the method for ultrasound elastography can include providing a compression on the target. The compression can include a natural compression, a motorized compression, a manual compression, a compression by an external source, and a compression by an ultrasound probe.

In certain embodiments, the RF can be recorded at a frequency of from about 2.5 MHz to about 10 MHz.

In certain embodiments, the beamforming can include performing a delay-and sum beamforming. In certain embodiments, the 3D displacement comprises an axial, a lateral, and/or an elevational direction. The 3D displacement is configured to be calculated between two successive volumes.

In certain embodiments, the method for ultrasound elastography can include determining a Lagrangian strain tensor. The method for ultrasound elastography can be utilized to detect and treat a breast cancer.

In certain embodiments, an exemplary system for ultrasound elastography, includes

a 2D matrix array probe and a signal processor. The 2D matrix array probe can include an ultrasound transducer which can emit at least one non-focused wave on a target. The signal processor can obtain at least two Radio Frequency (RF) signals from the non-focused wave, beamform 3D volumes from the RF, calculate at least two 3D displacement by comparing each 3D volume to a reference volume, and integrate the 3D displacements to create a 3D cumulative axial strain volume.

In certain embodiments, the 2D matrix array probe can include a plurality of elements. The 2D matrix array can emit a non-focused wave at rate of about 100 waves per second. In certain embodiments, the ultrasound transducer can have a central frequency of about 2.5MHz. As embodied herein, the 2D matrix probe can be configured to receive at least two Radio Frequency (RF) signals from the non-focused wave and transmit the RF signals to the signal processor.

BRIEF DESCRIPTION OF THE FIGURES

Figure 1. A method of ultrasound elastography according to one exemplary embodiment of the disclosed subject matter. Figure 2. An example setup ultrasound elastography according to one exemplary embodiment of the disclosed subject matter.

Figure 3. A schematic illustration of process for the elastography with plane waves framework according to one exemplary embodiment of the disclosed subject matter.

Figure 4. Diagrams of (A) 3D B-mode volume and (B) strain volume in (C) x-z plane and (D) x-y plane.

Figure 5. Diagrams of (A) 3D B-mode volume, (B) 3D axial strain volume and (C-D) slices of the strain volume of the stiff inclusion embedded in the soft phantom.

Figure 6. Diagrams of (A) 3D B-mode volume, (B) 3D axial strain volume and (C-D) slices of the strain volume of the stiff inclusion embedded in the stiff phantom.

Figure 7. Diagrams of (A, E) 3D B-mode volumes, (B,F) 3D strain volumes and (C, D, G, F) slices of the strain volumes of an ex vivo canine liver before and after an HIFU ablation.

Figure 8. Diagrams of strain distribution in the 3D strain volumes (A) before ablation and (B) after ablation.

Figure 9. Diagrams of (A) 3D B-mode volume and (B-C) slices and (D) 3D strain volume and (E-F) slices of an in vivo calf muscle of a human volunteer.

Figure 10. A schematic illustration of eletromechanical model of the human heart.

Figure 11. A schematic illustration of (A) 3D simulation configuration, (B) flow chart of 3D EWI and (C) process of the 3D simulation according to one exemplary embodiment of the disclosed subject matter. Figure 12. Diagrams of (A) displacement estimation, (B)

electromechanical strain estimation, and (C) comparison between true and estimated electrical activation times.

Figure 13. A schematic illustration of (A) 2: 1 multiplexer for 3D ultrasound and (B) in vivo electromechanical strain in a canine.

DETAILED DESCRIPTION

The disclosed subject matter provides methods and systems for ultrasound elastography, including the use of ultrasound to assess the mechanical properties of tissue in a three-dimensional volume. In certain embodiments, the disclosed methods and systems provide three-dimensional (3D) ultrasound elastography using non-focused waves.

For the purpose of example and not limitation, Figure 1 provides a schematic illustration of an exemplary method for 3D ultrasound elastography. In certain embodiments, a method 100 includes emitting at least one non-focused wave to a target 101. In certain embodiments, the target can include any biological tissue, as known in the art. For example and not limitations, the non-focused wave can be emitted to soft tissues such as muscles, tendons, ligaments, fascia, connective tissue, fat, heart, skin, liver, breast, prostate, thyroid, etc.

As embodied herein, the non-focused waves can be in various forms, including plane waves, circular waves, and/or spherical waves. In certain embodiments, the non-focused waves can be emitted by a ultrasound transducer. For example and not limitation, the ultrasound transducer can be configured to emit ultrasound waves at a high volume rate, for example and as embodied herein, at a rate of from about 50 volumes/s to about 1000 volumes/s, or from about 100 volumes/s to about 500 volumes/s. In particular embodiments, the rate of ultrasound wave emission can be about 100 volumes/s or about 500 volumes/s.

As embodied herein, the ultrasound transducer can emit waves for a certain period of time. As such, a known number of wave volumes can be emitted over that certain period of time based on the volume rate. By way of example, and not limitation, an ultrasound transducer operating at a volume rate of 100 volumes/s can be configured to emit waves for 1 second or more such that at least 100 plane waves are emitted. As embodied herein, for example and not limitation, the method can include emitting at least 20 plane waves, at least 50 plane waves, at least 100 plane waves, at least 250 plane waves, or at least 500 plane waves.

As used herein, the term "about" or "approximately" means within an acceptable error range for the particular value as determined by one of ordinary skill in the art, which will depend in part on how the value is measured or determined, i.e., the limitations of the measurement system. For example, "about" can mean a range of up to 20%, up to 10%, up to 5%), and or up to 1% of a given value.

With further reference to Figure 1, the method 100 can further include recording a Radio Frequency (RF) of the non-focused wave 102. In certain

embodiments, the RF can be a backscattered echo from an ultrasound wave, e.g., a plane wave. For example and not limitation, backscattered echoes from each emitted ultrasound wave can be recorded at a frequency of from about 2.5 MHz to about 10 MHz, e.g., at about 10 MHz. As embodied herein, the recorded RF can optionally be stored in memory.

As embodied herein, the method 100 can further include beamforming 3D volumes from the RF 103. In certain embodiments, the beamforming can be performed using a delay-and-sum algorithm, as is known in the art. By way of example, and not limitation, 3D B-mode volumes can be beamformed from the RF.

As shown for purposes of example in Figure 3, multiple 3D volumes can be generated from two or more RFs. Each 3D volume can correspond to one or more RFs from a single ultrasound wave. As such, in embodiments where multiple plane waves are emitted, a 3D volume can be beamformed to correspond to each plane wave.

The method 100 can further include calculating at least one 3D displacement compared to a reference volume 103. For example, the 3D displacements, such as displacement in an axial, a lateral, and/or an elevational direction, can be determined as between two successive 3D volumes. Thus, a first 3D volume

beamformed from a first ultrasound wave can form the reference volume for a second 3D volume beamformed from a second ultrasound wave. Where the waves are emitted from the same position at different times, the 3D displacement can reflect changes in the tissue geometry over time. Alternatively, where the waves are emitted from different positions, the 3D displacement can reflect changes in the tissue geometry in space. As embodied herein, the 3D displacement can be estimated by normalized ID and/or 3D cross- correlation and/or recorrelation, as known in the art.

The method 100 can further include integrating the 3D displacements to create a 3D cumulative axial strain volume 104. For example, the 3D incremental axial displacements can be integrated over time to obtain 3D cumulative axial displacements, as shown in Figure 3. Then, the 3D cumulative displacements can be filtered using a 3D median filter. The 3D cumulative axial strain distributions can be estimated from the 3D cumulative displacements by applying a least-squares estimator. The final 3D axial strain volume can be displayed graphically. Optionally, more than one 3D median filter can be applied on the 3D cumulative displacement volume or/and 3D strain volume. Lagrangian axial, lateral, and elevational strain volumes can be computed from the cumulative displacement volumes. In certain embodiments, a coordinate transformation can then be applied to retrieve circumferential, radial, and longitudinal strain volumes for in vivo applications such as the heart. In some embodiments, the presently disclosed techniques and systems can further provide the strain distribution in the elevational direction. As embodied herein, anisotropy and/or elasticity values can be determined from the elevational strain.

In certain embodiments, the method can include emitting ultrasound waves to create 3D cumulative axial strain volumes at different positions in the target tissue. For example, the method can be repeated at two or more positions to obtain multiple 3D cumulative strain volumes. The multiple 3D cumulative strain volumes can be assembled and/or integrated to obtain a single graphical representation of the 3D cumulative strain.

The method 100 can optionally further include providing a compression on the target prior to the emitting. The compression can include a natural compression, a motorized compression, a manual compression, a compression by an external source, and a compression by an ultrasound probe. For example, compression can be provided manually using freehand techniques, e.g., massage, with or without the use of a tool to apply compression. Additionally or alternatively, uniaxial compression can be applied using a linear motor. In certain embodiments, up to a 3% compression at a motor speed of 3% compression per second can be applied, although a person of skill in the art will appreciate that the magnitude and speed of compression can be adjusted to the desired application. Similarly, a stepper motor can be attached to the ultrasound probe and the compression can be applied by the ultrasound probe mechanically through the motor. Alternatively, compression can be applied by manually pressing the ultrasound probe. For non-moving organs, an external compressor can be applied on the target to induce deformation. For moving organs such as the heart, natural contraction can be used as the deformation. In certain embodiments, the compression can be provided prior to or/and during emission of the non-focused wave. The compression can be applied to the target continuously, intermittently, and/or incrementally.

In certain aspects, the present disclosure further provides systems for ultrasound elastography. For example and not limitation, the presently disclosed systems can provide a device for ultrasound elastography including a 2D matrix array probe and a signal processor. The 2D matrix array probe can be coupled to an ultrasound transducer to emit at least one non-focused wave on a target. As used herein, the word "coupled" means directly connected together or connected through one or more intervening elements. The connection can be a physical connection or an operable connection, e.g., such that a signal or wave can be transmitted between the elements. For example, the 2D matrix array probe can be connected to emission and receive channels of the ultrasound transducer to emit and receive ultrasound waves and/or Radio Frequency (RF) signals. The 2D matrix probe can receive at least two RF signals from the non-focused wave and transfer the RF signals to a signal processor. The 2D matrix array probe can be fully programmable and can include a plurality of elements. In certain embodiments, the 2D matrix array probe can include at least about 100 elements, at least about 200 elements, at least about 250 elements, at least about 500 elements, or at least about 1000 elements. For example, and not limitation, the probe can include an array of elements, which can include about 256 elements (e.g., 16x16) or about 1024 elements (e.g., 32x32) to create a programmable ultrasound system with about 256 or about 1024 channels.

In certain embodiments, the system can include a multiplexer (e.g., a 2: 1 multiplexer) in transmit and/or receive channels. For the purpose of example and not limitation, the system can include a combination of two or more ultrasound scanners and the multiplexer can be used to interface the channels of the ultrasound scanners and elements of the ultrasound transducer. For example and not limitation, in an ultrasound system that has two ultrasound scanners containing 256 channels each, a 2: 1 multiplexer can be used to interface the 1024 element probe with a 512 total channels on the ultrasound scanners. In such embodiments, the 2: 1 multiplexer can switch between 2 positions to interface the elements of the transducer with channels on the scanner, e.g., elements 1 through 512 or 513 through 1024.

In certain embodiments, the signal processor can record a RF of the non- focused wave, beamform a 3D volume from the RF, calculate at least one 3D

displacement, and/or integrate the 3D displacements to create a 3D cumulative axial strain volume. The ultrasound system can have a central frequency of about 2.5 MHz. The signal processor can include an ultrasound scanner, as described above.

For the purpose of illustration, Figure 2 provides an example schematic of the operation of a system according to the disclosed techniques. As shown in panel A of Figure 2, a 2D matrix array probe can be mounted on an axial linear motor. A compressor fitted with the footprint of the probe can be attached to the bottom of the probe. The target samples can be positioned underneath. The target can optionally be immersed in a water tank to conduct waves from the ultrasound probe.

The presently disclosed techniques and systems can be used to detect abnormalities in tissue. For example, these techniques can be used to detect various types of cancer, including breast cancer. Alternatively or additionally, the techniques can be used to monitor myocardial function, for example for ischemia and infract detections and three-dimensional myocardial ablation lesion monitoring. The presently disclosed systems and techniques can also be used to monitor radiofrequency ablation therapy for cardiac arrhythmia and high intensity focused ultrasound (HIFU) therapy for various cancers. By way of example, and not limitation, the presently disclosed techniques and systems can be utilized in the diagnosis of various diseases and disorders, including, but not limited to, atrial fibrillation, atrial flutter, ventricular tachycardia, and heart failure. In certain embodiments, the presently disclosed techniques and systems can be utilized for cardiac imaging, Doppler imaging, electromechanical wave imaging, and/or shear wave imaging.

In certain embodiments, the presently disclosed techniques and systems can decrease intra- and inter- observer variability that can occur when a 2D plane is imaged. For example, the presently disclosed techniques and systems can alleviate certain signal decorreleation between the first and the last 2D image acquired in the stack of 2D images contained in the 3D volume; or between the reconstructed 3D volumes before compression and after compression due to the short time needed to acquire 3D volumes. This decorrelation can be due to hand motion (in case of freehand

compression), normal physiologic motions induced by cardiac or respiratory activities, or to a too high strain amplitude between the two volumes. Signal decorrelation can affect a signal-to-noise ratio (S R) of the strain images. The presently disclosed techniques can acquire and process backscattered ultrasound emissions to estimate 3D axial strain volumes and prepare 3D images of axial strain in real time.

EXAMPLES

The presently disclosed subject matter will be better understood by reference to the following Examples. These Examples are provided as merely illustrative of the disclosed methods and systems, and should be considered as a limitation in any way. Example 1: 3D quasi-static ultrasound elastography with plane waves.

This Example describes one exemplary method of 3D quasi-static ultrasound elastography methods with plane waves using parallel receive beamforming that estimates axial strain distribution in vivo in entire volumes at a high volume rate, e.g., 100 volumes/sec or 500 volumes/sec.

Methods

Ultrasound system

A fully programmable ultrasound system with 256 fully programmable channels in emission and receive (Vantage, Verasonics, Kirkland, USA) was used to control an ultrasonic 2D matrix array probe of 256 elements (16-by-16 elements), with a 0.95mm ² pitch, a central frequency of 2.5 MHz, and a bandwith of 50% (Sonic

Concepts, Bothell, USA). The volume delay-and-sum beamforming and the axial strain distribution calculations were performed on a Tesla K40 GPU (Nvidia, Santa Clara, USA). 3D rendering was computed with Amira software (Visualization Sciences Group, Burlington, USA).

Experimental Setup:

A. Ex vivo study

The 2D matrix array probe was mounted on a linear motor (Figure 2, panel A). A square-plate compressor (lOcm-by-lOcm) was designed to fit the foot-print of the probe and to compress the samples with uniaxial compression. The samples were immersed inside a water tank with temperature maintained between 10°C and 15°C. A 2D real-time focused B-mode image was used to position the probe on the samples and a low axial pre-compression was applied to the samples before starting the experiments

(~i %). The samples were continuously compressed using a linear motor up to a 3% compression at a motor speed of 3% compression per second. Simultaneously, 100 plane waves were emitted from the 2D-matrix array probe at a rate of 100 plane waves per second in order to reconstruct 100 volumes for a total compression of 3% (Figure 3, panels A and B); the radio-frequency (RF) backscattered echoes from each plane wave were recorded at a frequency of 10 MHz and stored in memory.

B. Two-layer phantom

The two-layer gelatin phantom used is shown in panel B of Figure

2Error! Reference source not found.. The first layer was made from 3% of gelatin (Bloom-275) resulting to a 4.2-kPa stiffness. The second layer was made from 12% of gelatin resulting to a 75.3-kPa stiffness. In both layers, corn stash was added to improve the backscattering properties of the gels; 1.2% in the first layer and 0.3% in the second layer to be able to see the difference on the 3D B-mode volume.

C. Stiff inclusion

A stiff gelatin inclusion (12% gelatin) with a 140 mm diameter with 0.3%) corn stash; was embedded in a softer gelatin phantom (3%> gelatin) with dimensions 60mmx40mmx40mm with 1.2% corn stash (Figure 2, panel C).

D. Soft inclusion

A soft gelatin inclusion (3%> gelatin) with a 140 mm diameter with 1.2% corn stash; was embedded in a stiffer gelatin phantom (12% gelatin) with dimensions 60mmx40mmx40mm with 1.2% corn stash (Figure 2, panel D).

E. Ex vivo canine liver

The feasibility of 3D quasi-static elastography was then demonstrated in an ex vivo canine liver before and after a High Intensity Focused Ultrasound (HIFU) ablation (Figure 2, panel Error! Reference source not found.E). 3D quasi-static elastography was first performed on the liver without ablation. The real-time focused B- mode image of the 2D-matrix array probe was used to manually position a needle in the middle of the plane-of-view. Then, the HIFU setup was positioned at the same location using the needle as a landmark. After HIFU ablation, the needle was repositioned at the same location enabling the 2D matrix array probe to return at its previous location. 30 minutes post-ablation, the 3D quasi-static elastography was performed on the ablation lesion location. The HIFU ablation was performed using a 93 -element HIFU array (H- 178, Sonic Concept Inc. Bothell WA, USA) generating an amplitude-modulated signal (fcarrier = 4.5 MHz and /yvi = 25 Hz) with an acoustic power measured in water of 5.04 W. The liver sample was sonicated for 120 s, which has been shown to generate

reproducible lesion.

F. In vivo study

The feasibility of the 3D quasi-static elastography was then demonstrated in vivo in the calf muscle of a human volunteer. The calf was set resting on a chair while the 2D matrix array probe was hand-held allowing freehand scanning.

The calf muscle was continuously and smoothly compressed using the same square compressor used previously. Similarly to the ex vivo study, 100 2D-plane waves were emitted from the 2D-matrix array probe at a rate of 100 plane waves per second in order to reconstruct 100 volumes for the total freehand compression. The RF backscattered echoes from each plane wave were recorded at a frequency of 10 MHz and stored in memory.

Image formation and 3D strain calculation

The 3D quasi-static elastography framework is depicted in Figure 3.

From RF data, a classical three dimensional delay-and-sum algorithm was used to beamform a 3D volume from each plane wave acquisition, resulting in a total of 100 volumes (Figure 3, panel B). The lateral sizes (x and y directions) of the volumes were set to 15.2 mm corresponding to the aperture of the matrix probe whereas the

reconstructed depth was set to 30 mm-60 mm depending on the application. The lateral sampling was set to 237.5 μιη in x and y directions (Figure 3, panel B). The axial sampling was set to 61.6 μιη corresponding to a 1/10 beamforming (where λ is the ultrasonic wavelength). From the RF beamformed signals, B-mode volumes were displayed using a Hilbert transform and display with the Amira software. The 3D incremental axial displacements between two successive volumes were estimated by normalized ID cross-correlation with a window size of 6.16 mm (corresponding to a \0λ window size as indicated in and a 95% overlap) (Figure 3, panel C). The 3D

incremental axial displacements were then integrated over time to obtain 3D cumulative axial displacements (Figure 3, panel D). The 3D cumulative displacements were then filtered using a 3D median filter with a pixel kernel of 3 pixel s-by-3 pixel s-by-3 pixels (195μιη^-712μιη^-712μιη). 3D cumulative axial strain distributions were computed from 3D cumulative displacements by applying a least-squares estimator with a relatively small kernel of 646 μπι (Figure 3, panel E). A second 3D median filter with a pixel kernel of 3 pixel s-by-3 pixel s-by-3 pixels was then applied on the strain volume. The final 3D axial strain volume was then displayed using Amira software.

Two slices in the x-z and y-z plane of the 3D axial strain volume were displayed for the case of the inclusion phantoms and the liver. A region of interest (ROI) was arbitrarily applied at the strain boundary of the two layers phantom, around the inclusions and the ablation lesion. The absolute strain values were averaged inside (S _in ) and outside (S _out ) the ROI with associated standard deviations and an observed strain contrast (C) was calculated according to Formula 1, below: C = 201og ₁₀ (¾^) (Formula 1)

Sin

In the case of ex vivo liver ablation a histogram presenting the strain distribution in the volume has been calculated before and after ablation to highlight the strain increase due to ablation stiffening.

Results

Two-layer phantom

3D quasi-static elastography was first applied to a two-layer gelatin phantom. The layer at the top was made about eighteenth times softer and more echogenic than the layer at the bottom. The last 3D B-mode volume is shown in panel A of Figure 4. The difference in echogeneicity is clearly visible and the two layers are distinguishable. The 3D cumulative axial strain volume following a 3% compression was formed, as described in the methods. 3D cumulative axial strain volume is shown in Figure 4, panel B wih the associated slices in x-z plane (Figure 4, panel C) and in x-y plane (Figure 4, paenl D). The softer layer exhibited in average a higher absolute strain (S _in =2.2%t0.1%) than the stiffer layer (S _OMt =0.4%±0.2%), as it was expected. The observed strain contrast was calculated and gave a value of C = 14.8 dB. The boundary between the two layers on the 3D strain volume was in agreement with the boundary found on the 3D B-mode volume.

Stiff inclusion

3D quasi-static elastography was then applied to a stiff gelatin inclusion embedded in an eighteenth times softer gelatin phantom. The results are displayed in Figure 5. The inclusion was clearly identifiable on the 3D B-mode volume as it was made with a smaller concentration of scatterer (Figure 5, panel A). The 14 mm stiff inclusion was also identifiable on the 3D axial strain volume due to the lower absolute strain estimated at the inclusion location (5 _in =0.3%±0.4%) compared to the surrounding (5 _out =1.5%±0.4%) (i.e., Figure 5, panels B, C, and D). Figure 5, panel C and Figure 5, panel D show the cross-section of the 3D axial strain volume in the x-z plane and the y-z plane respectively with an arbitrary circle of a 14 mm diameter depicting the ROI where the absolute strain estimates were calculated. The observed contrast (Eq. 1) C = 13.9 dB, was lower than the two-layer phantom case. This can be explained by the geometry of the inclusion. Indeed, a lower contrast transfer efficiency, which represents the ratio of elasticity contrast measured from strain estimation to the true contrast, was expected in the case of a spherical inclusion. In addition, some of the predicted shadowing artifacts can be observed outside the inclusions on Figure 5, panel C and Figure Error! Reference source not found.5, panel D.

Soft inclusion

The method was then applied to a soft gelatin inclusion embedded in an eighteenth times stiffer gelatin phantom. The inclusion could not be seen on the 3D B- mode volume (Figure 6A) due to the identical concentration of scatterer inside and outside the inclusion. However, the inclusion could be detected on the 3D axial strain volume (Figure 6, panel B). The higher absolute strain (S _in =1.4%±0.8%) indicates that the inclusion is softer than the surroundings (S _out =0.6%±0.7%). Figure 6, panel C and Figure 6, panel D show the cross section of the 3D strain volume in x-z plane and y-z plane respectively with an arbitrary circle of a 14 mm diameter depicting the ROI where the absolute strain values were calculated. The observed contrast was calculated and found to be equal to: C = -7.4 dB. The minus sign indicates that this is the contrast for a soft inclusion (Eq. 1), however the absolute contrast found was significantly lower than the one for the stiff inclusion. This was predicted as the elastic modulus contrast of soft inclusion fundamentally, leads to a lower contrast transfer efficiency. The homogeneity artifacts inside the inclusions could also be explained by the low contrast transfer efficiency.

Ex vivo liver

3D quasi-static elastography was obtained to an ex vivo canine liver before and after a HIFU ablation. The results are shown in Figure 7 and Figure 8.

Before ablation, the 3D B-mode volume exhibited a heterogeneous echogenicity (Figure 7, panel A). This heterogeneity is also noticeable on the 3D axial strain volume (Figure 7, panel B) and 2D strain images (Figure Error! Reference source not found.7, panels C and D), revealing the structural complexity of the liver. After ablation, the 3D B-mode volume is more echogenic at the ablation location. Moreover, only based on the 3D B- mode volume, it's not possible to precisely detect the ablation. On the 3D axial strain volume (Figure 7F), one can notice a decrease on the absolute strain at the HIFU ablation location as expected. Indeed, the effect of HIFU ablation in biological tissue is an increase in stiffness. According to the average strain inside the ablation

(S _in =0.25%±l . l%) and outside the ablation (S _out =0.8%±1.3%) the observed contrast C = 10.1 dB was calculated and was found comparable to the stiff inclusion. However, the homogeneity artifacts are higher in this case as it is showed by the higher standard deviation from the average strain calculation. These artifacts could be due to the shadowing artifacts of a spherical inclusion.

In vivo calf of a human volunteer

The feasability of method was then demonstrated in vivo on the upper calf of a human volunteer. From the acquisition, B-mode volumes (Figure 9, panel A) and slices (Figure 9, panels B and C) were reconstructed. From the freehand compression and the transmission of 100 plane waves at 100 volumes/s, a 3D axial strain volume (Figure 9, panel D) were reconstructed and the softer layer (deep blue) consistent with the location of the fascia layer between the muscles of the gastrocnemius muscle and the soleus muscle (i.e., Figure 9, panel F) was detected. The two muscles exhibited similar strains.

Discussion

The objectives were to develop a new 3D axial strain method using plane waves at high volume rate. The feasibility of 3D quasi-static elastography based on three dimensions quasi-static elastography coupled with plane wave imaging at a high volume rate using a 2D matrix array probe were demonstrated. Axial strains in three dimensions in phantoms, in an ex vivo biological tissues for lesion detection, and in vivo in the calf muscle of a human volunteer can be estimated.

A 3% motorized continuous quasi-static compression was combined with 3D strain imaging with plane waves, in order to acquire 100 volumes at a volume rate of 100 volumes/s. This very high volume rate enables to compute one 3D volume of cumulative axial displacements made from 100 incremental axial displacements.

The method was validated in phantoms and showed good sensitivity. Indeed, strain differences on a two-layered phantom composed of different stiffness were detected. In addition, a 14 mm diameter stiff inclusion embedded in a soft gelatin phantom and a 14 mm diameter soft inclusion embedded in a stiff gelatin phantom were detected and visualized. The feasibility of the method in ex vivo canine liver before and after an HIFU ablation by detecting the stiffer lesion after ablation was shown. The feasibility of performing 3D quasi-static elastography with plane waves at a high volume rate in vivo was demonstrated on the calf muscle of a human volunteer by performing 3D axial strain with a simple freehand compression.

Using only one transmit plane wave to construct an entire volume enabled high volume rates and eliminated signal decorrelation occurring when multiple acquisitions from mechanical translation of a ID array were required to construct the volume. The acquisition time posed thus no longer a limitation.

For in vivo applications, such as breast cancer detection, the high volume rate used in this study will likely reduce signal decorrelations from physiologic motion such as the respiration or the heart rhythm. It also reduced artifacts from hand motions in the case of freehand scanning.

In addition, the high volume rate enabled high temporal resolution which leaded to lower strain amplitude between consecutive volumes. In this case, the Strain Filter framework predicted an improvement in the strain estimator performance by reducing signal decorrelation and increasing the correlation coefficient. In this study a volume rate of 100 volumes/s was used but the method enables the volume rate to be increased up to 3000 volumes/s, which could be used to optimize the technique based on the Strain Filter.

In addition, the high number of volumes acquired enabled the calculation of numerous incremental axial displacement volumes in order to calculate, in this study, one volume of 3D cumulative axial displacement and one 3D cumulative axial strain volume. This is a similar approach to the technique of multicompression averaging, which has been shown to increase strain signal-to-noise ratio by decreasing random noise.

This method based on axial strain estimation could also be extended to lateral and elevational strains to correct the axial strain or to obtain the full strain tensor in biological tissues, with a high temporal resolution.

Due to the associated high volume rates, this method could also be extended to moving organs such as the heart. Myocardial elastography, a method for the estimation of the strain distribution in the heart, could be measured in entire volumes at high temporal resolution.

The implementation on GPUs enabled us to obtain the 3D B-mode and the 3D axial strain in a few seconds following the acquisitions. This method has thus the potential to be implemented in real time by optimizing the processing.

Limitations to this method include the low frequency bandwith of the probe and the low number of elements in the lateral directions used, which resulted in a low signal-to-noise ratio in the strain estimation and the contrast transfer efficiency. This limitation could be ensured by using a 1024 elements probe with a larger bandwidth, controlled by two synchronized ultrasound systems with 256 fully programmable channels combined with synthetic emissions. The quality of the estimation could also be increased by increasing resolution, by implementing spatial coherent compounding in three dimensions from the transmission of multiple tilted plane waves instead of using only one plane wave performed in this study. Another limitation is the small field of view induced by the plane wave transmission from the small aperture of the probe. This issue could be addressed by implementing diverging wave transmits instead of plane waves while keeping a high volume rate.

Conclusion

3D quasi-static elastography with plane waves at high volume rate were developed and operated. By transmitting 100 plane waves at a volume rate of 100 volumes/s, the axial strain distribution in a two-layer gelatin phantom of two different stiffness, a soft inclusion embedded in a stiff gelatin phantom, and a stiff inclusion embedded in a soft gelatin phantom throughout the entire volume in three dimensions were detected. Moreover, the 3D axial strain in an ex vivo canine liver before and after HIFU ablation were estimated and the axial strain distribution in three dimensions to detect the lesion were mapped. Finally, in vivo feasibility of 3D quasi-static

elastography with plane waves was demonstrated on the calf of a human volunteer with a simple freehand scanning at a high volume rates. Due to the high volume rate, 3D elastography with plane waves for in vivo applications could have a major impact for three dimensional elasticity imaging.

Example 2: Three-dimensional cardiac electromechanical activation mapping with In Silico validation.

This Example describes one exemplary method of 3D electromechanical waves imaging (EWI) with 3D ultrasound in a single heartbeat In silico.

Methods

Experimental Setup:

A. In Silico study

The ventricular geometry was obtained from MRI images of a healthy volunteer. Finite element simulation of an electromechanical heart model during normal sinus rhythm was performed to design electromechanical model of the human heart as shown in Figure 10.

For 3D ultrasound simulation with Field II software, a 3D ultrasound probe with 32-by-32 transducer elements, a 9.6 mm ² foot-print and a central frequency of 3 MHz was modeled. During the stimulation with the Field II software, a simulated numerical phantom with a ventricular mask having a 170x170x120mm volume with lmm voxel size were added as inputs. Imaging sequences were set to emit 563 spherical waves at a rate of 556 volumes per second in order to reconstruct 3D volumes as shown in Figure 11, panel A. The radio-frequency (RF) channel data was recorded and stored in memory. The flow chart of 3D electromechanical wave imaging (EWI) analysis is depicted in Figure 11, panel B. From RF data, images were reconstructed. The interframe displacements were estimated between two successive interframes to calculate electromechanical interframe strain and EWI activation time.

B. In vivo study

An in vivo study was also performed on the heart of an open-chest canine with normal sinus rhythm. A fully programmable ultrasound system with fully programmable channels in emission and receive was used to control a 3D ultrasound probe of 32-by-32 transducer elements. A 2: 1 multiplexer was used for transmit and receive channels as shown in Figure 13. 500 spherical waves were emitted from the 2D- matrix array probe at a rate of 500 volumes per second in order to reconstruct 3D volumes; the radio-frequency (RF) backscattered echoes from each wave were recorded and stored in memory.

Results

The electromechanical wave was successfully imaged with 3D ultrasound in a single heartbeat in silico. As Figure 12 shows, the ventricular displacement and strain at a high volume rate in a realistic human heart model were accurately estimated when compared to an electromechanical model of a human hart with true electrical activation times and incremental displacements. The estimated electromechanical and true electrical activation times were strongly correlated. For example, as shown in Figure 12, correlation value between estimated electromechanical and true electrical activation times is R ² =0.86. As shown in Figure 13, panel B, in vivo 3D

electromechanical mapping of the heart was feasible during a single heartbeat with high frame rate and with high temporal resolution.

Discussion As demonstrated in this Example, the disclosed systems and techniques can be used for 3D electromechanical mapping of the heart, e.g., during a single heartbeat, and can provide validation of 3D EWI in arrhythmic patients. The 3D imaging of electromechanical strain at high volume rate can be used to show an earlier contraction in the atria than in the ventricles. Moreover, the 3D electromechanical mapping can be performed in vivo to study the electromechanical strain in a heart.

* * *

In addition to the various embodiments depicted and claimed, the disclosed subject matter is also directed to other embodiments having other combinations of the features disclosed and claimed herein. As such, the particular features presented herein can be combined with each other in other manners within the scope of the disclosed subject matter such that the disclosed subject matter includes any suitable combination of the features disclosed herein.

The foregoing description of specific embodiments of the disclosed subject matter has been presented for purposes of illustration and description. It is not intended to be exhaustive or to limit the disclosed subject matter to those embodiments disclosed.

It will be apparent to those skilled in the art that various modifications and variations can be made in the methods and systems of the disclosed subject matter without departing from the spirit or scope of the disclosed subject matter. Thus, it is intended that the disclosed subject matter include modifications and variations that are within the scope of the appended claims and their equivalents.