WO2016161120A1 - Systems and methods for low field magnetic resonance elastography - Google Patents

Abstract

Systems and methods for performing a medical imaging process are provided. In some aspects, a method includes arranging a subject in a low-field magnetic resonance imaging (LFMRI) system configured to generate a static magnetic field about at least a region of interest (ROI) of the subject comprising materials capable of producing magnetic susceptibility artifacts, and controlling the LFMRI system to perform a pulse sequence to acquire magnetic resonance elastography (MRE) data from the subject. The method also includes performing, during the pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data, and reconstructing, using the MRE data, at least one image indicative of mechanical properties of tissues in the subject.

Description

SYSTEMS AND METHODS FOR LOW FIELD MAGNETIC RESONANCE ELASTOGRAPHY CROSS-REFERENCE TO RELATED APPLICATIONS

[0001] This application is based on, claims priority to, and incorporates herein by reference, U.S. Provisional Application Serial No. 62/142,036, filed April 2, 2015, and entitled“MAGNETIC RESONANCE ELASTOGRPAHY IN THE PRESENCE OF IRON OVERLOAD.” STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH

[0002] This invention was made with government support under W81XWH-11-2-076 awarded by the Department of Defense. The government has certain rights in the invention. BACKGROUND

[0003] The present disclosure relates to systems and methods for magnetic resonance imaging (MRI). More particularly, the present disclosure provides systems and methods for low field MRI.

[0004] Manual palpation of soft tissue is regularly performed by clinicians to detect changes in organs and glands, and remains the most efficient screening tool for breast cancer detection. Beginning in the late 1980s, new medical imaging techniques enabled remote palpation of regions not directly accessible by the hands of the examining physician. Specifically, ultrasound was the first modality to provide qualitative and quantitative evaluation of stiffness in soft tissue. With the advantage of being inexpensive and portable, ultrasound has found broad use in clinical environments. However, acoustic impedance changes at interfaces (e.g. bones) limit the utility of ultrasound to applications in superficially accessible organs. Attenuation in tissue, and in particular in fat, also impairs its use in obese patients. Finally, shear waves do not propagate in liquids, making its use in patients with ascites impossible.

[0005] Magnetic resonance elastography (MRE) is another imaging technique that has been used to quantitatively measure mechanical properties of tissues. In MRE, steady-state shear or transverse waves produced using an external driver are directed to a region of interest (ROI). Dynamic MR data is then acquired at 1.5 or 3 Tesla, and processed to extract tissue stiffness of the ROI. Specifically, pulse sequences employing motion-sensitizing gradients synchronized to the applied waves are utilized to acquire phase contrast images. The phase contrast images are processed to extract wave propagation and produce displacement field maps. An inversion algorithm is then applied to the displacement field maps to calculate the tissue properties, typically displayed in an elastogram.

[0006] Unlike ultrasound, MRE can also provide high-resolution three- dimensional anatomical images, as well as wave attenuation related to the tissue viscosity. In particular, tissue viscosity has shown promising results in the diagnosis of non-alcoholic fatty liver diseases. In addition, in rodents and subsequently in humans MRE was shown to detect the presence of non-alcoholic steatohepatitis even before the appearance of fibrosis. Therefore, despite considerably longer acquisition times compared to ultrasound, MRE has been used routinely in clinical applications. For example, MRE has shown promising results for the diagnosis of chronic liver diseases, for discriminating between benign or malignant tumors in breast cancer patients, and for the screening of prostate cancer. MRE is also under development for the investigation of brain disorders, such as Alzheimer's disease.

[0007] However, MRE can suffer from inherently low signal-to-noise ratio (SNR) due to strong magnetic susceptibility artifacts. Specifically, magnetic susceptibility is a measure of the extent to which substances are magnetized when placed in a magnetic field. During imaging, magnetic susceptibility gradients cause magnetic field distortions that result in variations in precessional frequency of the imaged tissues. These in turn produce signal loss from dephasing and spatial mismapping of the MR signal. Strong susceptibility gradients are typically present in patients with implanted devices, as well as patients with iron overload. In particular, iron overload is found in up to 56% of patients with chronic liver disease. As a result, the signal intensity of the acquired images can be dramatically reduced, making diagnosis difficult if not impossible. [0008] Motion-sensitizing gradients played out in an MRE sequence also increase the echo time (TE), leading to a further decrease in sensitivity to short T2 species. Signal averaging can, to some extent, compensate for the reduction of sensitivity but comes at the expense of appreciably increasing acquisition times. This problem is further exacerbated in high field scanners, namely 1.5 T and above, since artifacts from magnetic susceptibility gradients increase with increasing field strength. In addition, poor SNR can also decrease the accuracy of the calculated tissues properties. This is because lower SNR leads to a drop in phase sensitivity, resulting in underestimating the mechanical properties of the tissue under study. Hence, the accuracy of diagnosis could then be adversely affected.

[0009] As appreciated from the above, for some patients, including those with metallic implants and iron overload, measurement of tissue properties using MRI systems can be particularly challenging, if not impossible, due to strong artifacts and low SNR. Presently, measurement alternatives include ultrasound, which has reported to have up to 75% less reliability than MRE, various X-ray imaging modalities, which utilize ionizing radiation, and surgical biopsy, which is invasive and potentially risky.

[0010] Therefore, there is a clear need for systems and methods capable of accurately and non-invasively measuring tissue properties with minimized artifacts. SUMMARY

[0011] The present invention overcomes the aforementioned drawbacks by providing systems and methods for magnetic resonance elastography (MRE) imaging. As will be described, the low-field imaging approach is described that substantially reduces magnetic susceptibility artifacts, leading to an increased sensitivity to mechanical properties of tissues and other imaged objects.

[0012] In accordance with one aspect of the disclosure, a magnetic resonance imaging (MRI) system configured to perform an imaging process of a subject. The system includes a magnet system configured to generate a static magnetic field about at least a region of interest (ROI) of the subject arranged in the MRI system, a plurality of gradient coils configured to establish at least one magnetic gradient field with respect to the static magnetic field, and a radio frequency (RF) system configured to deliver excitation pulses to the subject. The system also includes a magnetic resonance elastography (MRE) driver configured to deliver an oscillatory stress to the subject, and a computer system programmed to control the plurality of gradient coils, RF system and driver system to perform a three-dimensional (3D) balanced steady-state free precession (b-SSFP) to acquire MRE data from the subject, and perform, during the 3D b-SSFP pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data. The computer system is also configured to reconstruct, using the MRE data, at least one image of the subject.

[0013] In accordance with another aspect of the disclosure, a method for performing a medical imaging process is provided. The method includes arranging a subject in a low-field magnetic resonance imaging (LFMRI) system configured to generate a static magnetic field about at least a region of interest (ROI) of the subject comprising materials capable of producing magnetic susceptibility artifacts, and controlling the LFMRI system to perform a pulse sequence to acquire magnetic resonance elastography (MRE) data from the subject. The method also includes performing, during the pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data, and reconstructing, using the MRE data, at least one image indicative of mechanical properties of tissues in the subject.

[0014] In accordance with another aspect of the disclosure, a method for performing a medical imaging process is provided. The method includes arranging a subject in a magnetic resonance imaging (MRI) system for imaging at least a region of interest (ROI) comprising materials capable of producing magnetic susceptibility artifacts, and generating, using the magnetic resonance imaging (MRI) system, a static magnetic field configured to minimize the magnetic susceptibility artifacts. The method also includes controlling the MRI system to perform a pulse sequence to acquire magnetic resonance elastography (MRE) data from the ROI, and performing, during the pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data. The method further includes reconstructing, using the MRE data, at least one image indicative of mechanical properties of tissues in the subject.

[0015] The foregoing and other advantages of the invention will appear from the following description. BRIEF DESCRIPTION OF THE DRAWINGS

[0016] The patent or application file contains at least one drawing executed in color. Copies of this patent or patent application publication with color drawing(s) will be provided by the Office upon request and payment of the necessary fee.

[0017] FIG. 1 is a block diagram of an MRI system, in accordance with aspects of the present disclosure.

[0018] FIG. 2 is a block diagram of an RF system of the MRI system of FIG. 1.

[0019] FIG. 3 is a block diagram of a low-field MRI system, in accordance with aspects of the present disclosure.

[0020] FIG. 4 is a diagram of an example pulse sequence, in accordance with aspects of the present disclosure.

[0021] FIG. 5 are magnitude images comparing data acquired at high and low magnetic field, in accordance with aspects of the present disclosure.

[0022] FIG. 6 is a graph showing accumulated phase for three different encoding gradient strengths.

[0023] FIG. 7A is a graph showing simulated flip angle versus frequency offset for use in selecting optimized imaging parameters, in accordance with aspects of the present disclosure.

[0024] FIG. 7B is a graph showing simulated normalized echo amplitude versus echo number for use in selecting optimized imaging parameters in accordance with aspects of the present disclosure.

[0025] FIG. 8A are plots showing the effect on undersampling on phase accumulation.

[0026] FIG. 8B is a graph further showing the effect on undersampling on phase accumulation. [0027] FIG. 9 are maps indicating storage and loss moduli for a phantom measured in accordance with aspects of the present disclosure.

[0028] Other aspects and advantages of the present disclosure will become apparent upon consideration of the following detailed description and attached drawings.

DETAILED DESCRIPTION

[0029] Magnetic resonance elastography (MRE) is a powerful technique to assess the mechanical properties of living tissues non-invasively. However, it suffers from reduced sensitivity in regions with short T₂ and T₂ ^* relaxation times, as found in tissues with high concentrations of paramagnetic iron, or in regions surrounding implanted metals or devices. With potential for diagnosing of chronic liver diseases and staging liver fibrosis, as well as assessing other conditions, there is a need for improved methods for non-invasively measuring mechanical properties of tissues.

[0030] Therefore, the present disclosure provides systems and methods that overcome drawbacks of previous technologies. In particular, the present disclosure describes a low-field imaging approach that substantially reduces magnetic susceptibility artifacts and increases T2 and T2^*, thus leading to an increased sensitivity to mechanical tissue properties. Because magnetic resonance (MR) signals are reduced at lower magnetic fields, the present disclosure recognizes that Overhauser dynamic nuclear polarization (DNP) may be used to enhance the signals. Specifically, Overhauser DNP is well suited to low magnetic field strengths as the penetration depth of the electron saturation pulse is greater at lower frequencies, and sample heating is reduced since the specific absorption rate is proportional to the applied frequencies.

[0031] As will be appreciated from descriptions herein, the present disclosure extends MRE imaging to applications previously not conceived or thought possible using previous systems and methods. Specifically, MRE performed at low magnetic fields, in accordance with the present disclosure, can provide robust and artifact-free images in the presence of iron overload. Typically, normal iron content in the body is about 3 to 4 grams. However, in patients with iron overload, iron content can be as much as 20 grams, producing significant imaging artifacts. In addition, the present approach can also be used for patients with implants and implanted devices, that include metallic components, or for patients with unacceptably large magnetic susceptibility artifacts.

[0032] Referring particularly now to FIG. 1, an example of a magnetic resonance imaging (MRI) system 100 is illustrated. The MRI system 100 includes an operator workstation 102, which will typically include a display 104, one or more input devices 106, such as a keyboard and mouse, and a processor 108. The processor 108 may include a commercially available programmable machine running a commercially available operating system. The operator workstation 102 provides the operator interface that enables scan prescriptions to be entered into the MRI system 100. In general, the operator workstation 102 may be coupled to four servers: a pulse sequence server 110; a data acquisition server 112; a data processing server 114; and a data store server 116. The operator workstation 102 and each server 110, 112, 114, and 116 are connected to communicate with each other. For example, the servers 110, 112, 114, and 116 may be connected via a communication system 117, which may include any suitable network connection, whether wired, wireless, or a combination of both. As an example, the communication system 117 may include both proprietary or dedicated networks, as well as open networks, such as the internet.

[0033] The pulse sequence server 110 functions in response to instructions downloaded from the operator workstation 102 to operate a gradient system 118 and a radiofrequency (“RF”) system 120. Gradient waveforms necessary to perform the prescribed scan are produced and applied to the gradient system 118, which excites gradient coils in an assembly 122 to produce the magnetic field gradients G _x , G _y , and G _z used for position encoding magnetic resonance signals. The gradient coil assembly 122 forms part of a magnet assembly 124 that includes a polarizing magnet 126 and a whole-body RF coil 128 and/or local coil, such as a head coil 129.

[0034] RF waveforms are applied by the RF system 120 to the RF coil 128, or a separate local coil, such as the head coil 129, in order to perform the prescribed magnetic resonance pulse sequence. Responsive magnetic resonance signals detected by the RF coil 128, or a separate local coil, such as the head coil 129, are received by the RF system 120, where they are amplified, demodulated, filtered, and digitized under direction of commands produced by the pulse sequence server 110. The RF system 120 includes an RF transmitter for producing a wide variety of RF pulses used in MRI pulse sequences. The RF transmitter is responsive to the scan prescription and direction from the pulse sequence server 110 to produce RF pulses of the desired frequency, phase, and pulse amplitude waveform. The generated RF pulses may be applied to the whole-body RF coil 128 or to one or more local coils or coil arrays, such as the head coil 129.

[0035] The RF system 120 also includes one or more RF receiver channels. Each RF receiver channel includes an RF preamplifier that amplifies the magnetic resonance signal received by the coil 128, 129 to which it is connected, and a detector that detects and digitizes the I and Q quadrature components of the received magnetic resonance signal. The magnitude of the received magnetic resonance signal may, therefore, be determined at any sampled point by the square root of the sum of the squares of the I and Q components: M ^{I 2 ^ Q 2}

(_1);

[0036] and the phase of the received magnetic resonance signal may also be determined according to the following relationship: M ^{tan ^ 1 §} Q ^·

¨ ^I _¸

© _{¹ (2).}

[0037] The pulse sequence server 110 also optionally receives patient data from a physiological acquisition controller 130. By way of example, the physiological acquisition controller 130 may receive signals from a number of different sensors connected to the patient, such as electrocardiograph (“ECG”) signals from electrodes, or respiratory signals from a respiratory bellows or other respiratory monitoring device. Such signals are typically used by the pulse sequence server 110 to synchronize, or“gate,” the performance of the scan with the subject’s heart beat or respiration. [0038] The pulse sequence server 110 also connects to a scan room interface circuit 132 that receives signals from various sensors associated with the condition of the patient and the magnet system. It is also through the scan room interface circuit 132 that a patient positioning system 134 receives commands to move the patient to desired positions during the scan.

[0039] The digitized magnetic resonance signal samples produced by the RF system 120 are received by the data acquisition server 112. The data acquisition server 112 operates in response to instructions downloaded from the operator workstation 102 to receive the real-time magnetic resonance data and provide buffer storage, such that no data is lost by data overrun. In some scans, the data acquisition server 112 does little more than pass the acquired magnetic resonance data to the data processor server 114. However, in scans that require information derived from acquired magnetic resonance data to control the further performance of the scan, the data acquisition server 112 is programmed to produce such information and convey it to the pulse sequence server 110. For example, during prescans, magnetic resonance data is acquired and used to calibrate the pulse sequence performed by the pulse sequence server 110. As another example, navigator signals may be acquired and used to adjust the operating parameters of the RF system 120 or the gradient system 118, or to control the view order in which k-space is sampled. In still another example, the data acquisition server 112 may also be employed to process magnetic resonance signals used to detect the arrival of a contrast agent in a magnetic resonance angiography (MRA) scan. By way of example, the data acquisition server 112 acquires magnetic resonance data and processes it in real-time to produce information that is used to control the scan.

[0040] The data processing server 114 receives magnetic resonance data from the data acquisition server 112 and processes it in accordance with instructions downloaded from the operator workstation 102. Such processing may, for example, include one or more of the following: reconstructing two- dimensional or three-dimensional images by performing a Fourier transformation of raw k-space data; performing other image reconstruction algorithms, such as iterative or backprojection reconstruction algorithms; applying filters to raw k-space data or to reconstructed images; generating functional magnetic resonance images; calculating motion or flow images; and so on.

[0041] Images reconstructed by the data processing server 114 are conveyed back to the operator workstation 102 where they are stored. Real-time images are stored in a data base memory cache (not shown in FIG. 1), from which they may be output to operator display 112 or a display 136 that is located near the magnet assembly 124 for use by attending physicians. Batch mode images or selected real time images are stored in a host database on disc storage 138. When such images have been reconstructed and transferred to storage, the data processing server 114 notifies the data store server 116 on the operator workstation 102. The operator workstation 102 may be used by an operator to archive the images, produce films, or send the images via a network to other facilities.

[0042] The MRI system 100 may also include one or more networked workstations 142. By way of example, a networked workstation 142 may include a display 144; one or more input devices 146, such as a keyboard and mouse; and a processor 148. The networked workstation 142 may be located within the same facility as the operator workstation 102, or in a different facility, such as a different healthcare institution or clinic.

[0043] The networked workstation 142, whether within the same facility or in a different facility as the operator workstation 102, may gain remote access to the data processing server 114 or data store server 116 via the communication system 117. Accordingly, multiple networked workstations 142 may have access to the data processing server 114 and the data store server 116. In this manner, magnetic resonance data, reconstructed images, or other data may exchanged between the data processing server 114 or the data store server 116 and the networked workstations 142, such that the data or images may be remotely processed by a networked workstation 142. This data may be exchanged in any suitable format, such as in accordance with the transmission control protocol (TCP), the internet protocol (IP), or other known or suitable protocols.

[0044] With reference to FIG. 2, the RF system 120 of FIG. 1 will be further described. The RF system 120 includes a transmission channel 202 that produces a prescribed RF excitation field. The base, or carrier, frequency of this RF excitation field is produced under control of a frequency synthesizer 210 that receives a set of digital signals from the pulse sequence server 110. These digital signals indicate the frequency and phase of the RF carrier signal produced at an output 212. The RF carrier is applied to a modulator and up converter 214 where its amplitude is modulated in response to a signal, R _^ t _^ , also received from the pulse sequence server 110. The signal, R _^ t _^ , defines the envelope of the RF excitation pulse to be produced and is produced by sequentially reading out a series of stored digital values. These stored digital values may be changed to enable any desired RF pulse envelope to be produced.

[0045] The magnitude of the RF excitation pulse produced at output 216 is attenuated by an exciter attenuator circuit 218 that receives a digital command from the pulse sequence server 110. The attenuated RF excitation pulses are then applied to a power amplifier 220 that drives the RF transmission coil 204.

[0046] The MR signal produced by the subject is picked up by the RF receiver coil 208 and applied through a preamplifier 222 to the input of a receiver attenuator 224. The receiver attenuator 224 further amplifies the signal by an amount determined by a digital attenuation signal received from the pulse sequence server 110. The received signal is at or around the Larmor frequency, and this high frequency signal is down converted in a two step process by a down converter 226. The down converter 226 first mixes the MR signal with the carrier signal on line 212 and then mixes the resulting difference signal with a reference signal on line 228 that is produced by a reference frequency generator 230. The down converted MR signal is applied to the input of an analog-to-digital (“A/D”) converter 232 that samples and digitizes the analog signal. The sampled and digitized signal is then applied to a digital detector and signal processor 234 that produces 16-bit in-phase _^I _^ values and 16-bit quadrature _^Q _^ values corresponding to the received signal. The resulting stream of digitized I and Q values of the received signal are output to the data acquisition server 112. In addition to generating the reference signal on line 228, the reference frequency generator 230 also generates a sampling signal on line 236 that is applied to the A/D converter 232.

[0047] The basic MR systems and principles described above may be used to inform the design of other MR systems that share similar components but operate at very-different parameters. Specifically, a low-field magnetic resonance imaging (LFMRI) system, in accordance with aspects of the present disclosure, may utilize much of the above-described hardware, but with substantially reduced hardware requirements and a smaller hardware footprint.

[0048] Referring particularly to FIG. 3, one embodiment of a LFMRI system 300 is shown. Specifically, the LFMRI system 300 can include a nuclear magnetic resonance (NMR) system 302, an electron spin resonance (ESR) system 304, a magnet system 306, and a driver system 308. The LFMRI system 300 also includes a controller 310, in communication with the each of the above systems, and configured to control the operation of the LFMRI system 300. In some implementations, the controller 310 may include at least one workstation or computer system in communication with various servers (not shown in FIG. 3), including a pulse sequence server, a data acquisition server, a data processing server, and a data store server, as described.

[0049] As shown in FIG. 3, the NMR system 302 is connected to various NMR coils 312 configured to acquire magnetic resonance data, including MRE data, from a subject. In one embodiment, the NMR coils 312 may be in the form of a solenoid, although other coil configurations may be possible. In a prescribed magnetic resonance pulse sequence, waveforms are generated by the NMR system 302, as directed by the controller 310. The waveforms are then amplified and transmitted to the NMR coils 312 via one or more RF transmit channels. As such, the NMR system 302 may include an RF transmitter responsive to the scan prescription to produce RF pulses of the desired frequency, phase, and pulse amplitude waveform. Each RF transmit channel may include an RF amplifier that amplifies signals produced by the NMR system 302.

[0050] Magnetic resonance signals picked up by the NMR coils 312 may then be amplified, demodulated, filtered, and digitized by the NMR system 302 before processing. As such, the NMR system 302 also includes an RF receiver for receiving induced signals using one or more RF receiver channels. Each RF receiver channel may include an RF preamplifier that amplifies the magnetic resonance signals received by the NMR coils 312 to which it is connected. In some implementations, the same channels may be utilized for the RF transmit channels and RF receiver channels. To this end, various switching components may be utilized to select the mode of operation, that is transmitting or receiving RF signals. By way of example, the RF transmitter and RF receiver may be configured to operate in a frequency range between 1 kHz and 500 MHz, although other frequencies may also be possible.

[0051] The LFMRI system 300 also includes an ESR system 304 connected to one or more ESR coils 314, as shown in FIG. 3. The ESR coils 314 are configured to transmit RF signals that can enhance MR signals induced in the NMR coils 312, via Overhauser DNP, as described. In one embodiment, the ESR coils 314 may be configured as an Alderman-Grant resonator, although other coil configurations may be possible. By way of example, the ESR coils 314 may be tuned to a low energy transition of approximately 140 MHz of a nitroxide radical. As such, the ESR system 304 may also include a tuning or matching circuit connected to the ESR coils 314.

[0052] Although the NMR system 302 and ESR system 304 are shown in FIG. 3 as separate systems, it may be appreciated that they could be combined into a single RF system configured to generate, transmit and receive signals at various frequencies using RF channels, as described.

[0053] As shown, the LFMRI system 300 also includes a magnet system 306 in communication with a magnet assembly 316 that includes a biplanar electromagnet 318 and gradient coils. In some implementations, the gradient coils include z gradient circular coils 320, and x and y gradient rectangular coils 322, as shown in FIG. 3. The magnet system 306 controls the biplanar electromagnet 318 to generate a polarizing field about a subject, which in accordance with aspects of the present disclosure can be a low-field static magnetic field less than 10 mT, for example. During a pulse sequence, gradient waveforms are produced and applied by the magnet system 306, as directed by the controller 310. The gradient waveforms energize the gradient coils in the magnet assembly 316 to produce the magnetic field gradients G _x , G _y , and G _z used for position and motion encoding magnetic resonance signals.

[0054] In accordance with aspects of the present disclosure, the LFMRI system 300 also includes a driver system 308 connected to a driver 324 connectable to an imaged subject. The driver 324 is configured to induce vibrations in the subject during a magnetic resonance pulse sequence. For instance, the driver 324 may be configured to produce steady-state shear or transverse waves in an imaged subject with frequencies in a range between 50 to 500 Hz, although other frequencies may be possible. In some implementations, the driver 324 may be configured to couple to the subject in a fashion that minimizes signal attenuation. By way of example, the driver 324 may be a pneumatic driver, a piezoelectric driver, an electromechanical transducer, or an acoustic driver. Hence, the driver system 308 may include a variety of hardware and components for modulating the driver 324. In one embodiment, the driver system 308 may include a loudspeaker connectable to an acoustic driver 324 via an acoustic waveguide.

[0055] During operation, the controller 310 may be configured to control the NMR system 302, ESR system 304, magnet system 306 and driver system 308 to perform a pulse sequence configured to acquire MRE, and other data, from the subject at low magnetic fields. For example, as will be described, the pulse sequence may include a 3D balanced steady-state free precession (b-SSFP) sequence combined with fractional encoding and an undersampling scheme with a variable density Gaussian pattern. The controller 310 may then process the acquired MRE data to generate a report indicative of mechanical properties of tissues imaged in a subject. As such, the controller 310 may reconstruct various images, using MRE and other data, such as phase contrast images, anatomical images, and so forth. In some aspects, the reconstructed images are processed to provide information related to the mechanical properties of the subject. For instance, the controller 310 may perform phase unwrapping of the phase contrast images using a Laplacian unwrapping algorithm and generate one or more displacement field maps, viscoelastic maps or elastograms associated with an ROI of a subject using the unwrapped phase contrast images. In some aspects, the controller 310 may also correct for B₀ drift.

[0056] The NMR system 302, ESR system 304, magnet system 306 and driver system 308 of the LFMRI system 300 are shown in FIG. 3 to be outside of a Faraday cage 326, which is configured to provide RF-shielding. However, it may be appreciated that various components of the above systems may be located inside the Faraday cage 326. For example, resonances boxes, transmit/receive switches, pre-amplifiers, and other components may be advantageously located inside the Faraday cage 326. In addition, in some implementations, the LFMRI system 300 may further be configured to be portable, providing increased flexibility compared to traditional MRI systems.

[0057] Referring now to FIG. 4, a diagram of a pulse sequence 400 in accordance with aspects of the present disclosure is shown. The pulse sequence 400 is based on a b-SSFP-based sequence that is modified for acquiring MRE data at low magnetic fields, hereafter referred to as an MRE-bSSFP sequence. Specifically, the pulse sequence 400, carried out during a steady-state wave 402, begins with an NMR excitation pulse 402 that is followed by a first ESR pulse 404 played out during application of at least one motion encoding gradient (MEG) 406 along various gradient directions. By way of example, a pulse frequency of the NRM excitation pulse 402 can be 276 kHz, with flip angle values approximately between 70° and 90°, although other frequencies and flip angles may be possible. In some implementations, a frequency of the MEG 406 is different than the steady-state wave 402, and more specifically higher to significantly reduce the time spent per TR. For example, the steady-state wave 402 may be driven at 103 Hz while the frequency of the MEG 406 is 206 Hz, although other frequencies may possible.

[0058] A second ESR pulse 408 and a third ESR pulse 410 are further played out during the application of phase encoding gradients 412 and phase rewinder gradients 414, respectively. As described, application of ESR pulses allows for substantial enhancement of the measured signals 416 via Overhauser DNP. By way of example, the applied ESR pulses may be at a frequency of approximately 140 MHz, although other values may be possible. The above pulse sequence 400 may be repeated for a number of TR to acquire sufficient MRE data for a target SNR. In some aspects, a full sampling or undersampling strategy may be utilized when acquiring the MRE data. For example, a 25%, 50%, 75% or 100 % sampling may be utilized, although other values may be possible. In addition, the undersampling strategy may be based on variable density Gaussian patterns.

[0059] By way of example, FIG. 5 shows magnitude images acquired from polyvinyl alcohol (PVA) gel phantoms having different iron oxide (IO) concentrations. An acoustic waveguide was placed on top of the gels to generate acoustic waves. Relaxations times of the gels were measured at a very low and high magnetic fields. Specifically, spin-echo (SE), gradient-echo (GE), and b- SSFP scans were performed in a 1.5 T-Siemens scanner with a 32-channel head coil using TE/TR values equivalent to those previously reported for MRE. For the low field images, an MRE-bSSFP pulse sequence was performed using a low field MRI scanner, as described with reference to FIG. 3. Specifically, proton- density weighted images were obtained at 6.5 mT using an NMR frequency of approximately 276 kHz. A 50%-undersampling and fractional encoding (ratio q between vibration and encoding frequencies = 0.5) was used for acquisitions at (1.5×2.5×5.7) mm³ resolution. The total 3D acquisition time was 279 sec for 20 averages.

[0060] The magnitude images acquired using the GE sequence 502, SE sequence 504, b-SSFP sequence 506, and MRE-bSSFP sequence 508 for 0%, 1% and 2% IO concentrations are shown in FIG. 5. As appreciated from the figure, gel samples imaged at low field were homogenous, but were prone to dramatic drop of signal or susceptibility and banding artifacts at high field. A decreased T₁ as well as an increased T₂ was observed with decreasing magnetic field strength. In addition, SNR decreased with increasing iron content at 1.5 T, whereas it remained constant at 6.5 mT.

[0061] In addition to descriptions above, specific examples are provided below, in accordance with the present disclosure. These examples are offered for illustrative purposes only, and are not intended to limit the scope of the present invention in any way. Indeed, various modifications in addition to those shown and described herein will become apparent to those skilled in the art from the foregoing description and the following example and fall within the scope of the appended claims.

EXAMPLE

[0062] Magnetic resonance elastography (MRE) is a powerful technique to assess the mechanical properties of living tissue. However, it suffers from reduced sensitivity in regions with short T2 and T2* such as in tissue with high concentrations of paramagnetic iron, or in regions surrounding implanted devices. In this work, longer T2* are exploited, attainable at ultra-low magnetic fields in combination with Overhauser dynamic nuclear polarization (DNP) to enable rapid MRE at 0.0065 T. A modified 3D balanced steady-state free precession based MRE sequence with undersampling and fractional encoding was implemented on a 0.0065 T MRI scanner. A custom-built RF coil for DNP and a programmable vibration system for elastography were developed. Displacement fields and stiffness maps were reconstructed from data recorded in a polyvinyl alcohol gel phantom loaded with stable nitroxide radicals. A DNP enhancement of 25 was achieved during the MRE sequence, allowing the acquisition of 3D Overhauser-enhanced MRE (OMRE) images with (1.5 × 2.7 × 9) mm³ resolution over eight temporal steps and 11 slices in 6 minutes. Results described herein illustrate that OMRE at ultra-low magnetic field can be used to detect mechanical waves over short acquisition times. This new modality shows promise to broaden the scope of conventional MRE applications, and may extend the utility of low-cost, portable MRI systems to detect elasticity and other mechanical property changes in patients with implanted devices or iron overload. METHODS

Phantom preparation

[0063] A 7%-PVA (polyvinyl alcohol) gel containing 5mM 4-hydroxy- 2,2,6,6-tetramethylpiperidine 1-oxyl (TEMPOL) dissolved in water (Sigma- Aldrich, St. Louis, MO, USA) was obtained by two cycles of freezing–thawing at - 20 °C and room temperature. The gel was placed in a 5.5 cm inner diameter (ID), 6.0 cm outer diameter (OD), and 10.0 cm long 3D-printed cylindrical holder made of polycarbonate. T₁ and T₂ relaxation times of the MRE phantom were obtained with conventional inversion recovery and Carr–Purcell–Meiboom–Gill spectroscopic measurements at 6.5 mT: T1 = 200 ± 2 ms; T2 = 156 ± 2 ms. Ultra-low-field (ULF) MRI scanner and coil design

[0064] MRE was performed in a custom-built ULF MRI scanner consisting of a bi-planar 6.5 mT electromagnet with bi-planar gradients, as described with reference to FIG. 3. A Redstone MRI console (Tecmag, Houston, TX, USA) was used for sequence programming, data acquisition, and hardware (gradient and RF amplifiers) control. MR signal enhancement was obtained using Overhauser DNP at 6.5 mT. Two coils were used: a 10 cm OD 16 cm long solenoid coil for NMR excitation and detection at 276 kHz, and a 7 cm OD 13 cm long Alderman- Grant resonator for electron spin resonance (ESR) irradiation at 140.8 MHz. The ESR coil was placed inside the NMR coil. SNR in the motion-encoded images was calculated using the ratio between the mean signal magnitude over the whole phantom volume and the standard deviation of a user-defined noise region on each dataset. Wave generation

[0065] Acoustic pressure waves were generated by a 30 cm diameter loudspeaker (B&C Speakers, Pompton Plains, NJ, USA) located outside of the Faraday cage enclosing the low-field scanner (see Fig. 1(a)). A 7.4m long 22mm ID tube was used to guide the acoustic waves into the scanner. The extremity of the tube was closed with a flexible rubber membrane positioned on top of the investigated gel (Fig.1(b)). The resonance frequencies fn of the system are given

[0066] where L is the length of the tube, c_a the speed of sound in air, and n is an integer. The computed resonances were experimentally validated with a microphone (error ±1 Hz) by sweeping the frequency from 1 to 250 Hz. Computed resonances were compared with measured frequencies. Longitudinal waves were transmitted to the gel phantom. Ultra-low-field MRE sequence

[0067] To account for reduced NMR sensitivity in the ULF regime, signal averaging was performed. In order to avoid prohibitively long acquisition times, fast acquisition strategies were employed to accelerate imaging. A 3D balanced steady-state free precession (bSSFP)-based sequence, combined with fractional encoding for MRE and undersampling schemes with variable density Gaussian patterns, was used. bSSFP sequences allow the highest SNR per unit time by pulsing rapidly without waiting for T₁ recovery. A steady state was reached after a short number of pulses, and the transverse magnetization was maximal for T₂/_T1 ~ 1, similar to the regime observed at ULF. Fractional encoding allows encoding the motion at a frequency (f_G) higher than the driving frequency for the vibrations (fV), significantly reducing the time per TR needed for motion encoding. The phase accumulation obtained with this technique is described by the following equation

[0068] where q is the ratio between the vibration frequency f_V and the MSG frequency f_G, γ the gyromagnetic ratio, A_tot the displacement amplitude, and G the gradient amplitude. The motion sensitivity of the ULF scanner (with imaging gradient strength of the order of 1 mT/m) was simulated and compared with a conventional high-field clinical scanner with gradients ranging from 20 to 30 mT/m (FIG. 6). This comparison reflects only the motion sensitivity efficiency and not differences in SNR. The addition of SNR considerations to Eqn. 4 allows assessment of the minimum displacement the 6.5 mT system can detect given the error on the phase images (err = 1/SNR).

[0069] Simulations were performed with one MSG alternatively positioned on the three spatial axes: X (read direction), Y (3D phase encode direction), and Z (2D phase encode direction) (see FIG. 4). Transverse magnetization using T₁ and T₂ measured in the MRE phantom was calculated using a MATLAB-based Bloch equation solver for a bSSFP-based sequence with varying flip angle and off- resonance values due to B₀ inhomogeneities. Results are shown in FIG.7A.

[0070] The imaging parameters were the following: matrix size =(65 × 64 × 11), resolution = (1.5 × 2.7 × 9) mm³, TE/TR = 24/39 ms, α = 90°, and NA = 1. The flip angle used was the optimal one as determined in the simulations described above.

[0071] Eight temporal steps evenly distributed over one period of the acoustic excitation were acquired, leading to an acquisition time of 111.5 s per spatial direction. One additional scan was performed with the transducer turned off and served as a reference for B₀ drift correction. A total ESR irradiation time of 28 ms/TR was used to drive Overhauser DNP. As shown in FIG. 7B, ESR irradiation was included within each TR, consisting of a significant source of acceleration compared with conventional OMRI sequences, which usually require long prepolarization pulses for each TR before the MRI acquisition starts. Bloch simulations were performed to model the transverse magnetization approach to steady state in this interleaved hyperpolarization–acquisition scheme, and estimate the maximum signal enhancement due to the Overhauser effect. The input parameters used in the simulation included measured T1 and T2 relaxation times, the measured maximum enhancement obtained with an ESR pulse width of five times the proton T₁ of the sample in a 1D spectroscopy experiment (-30- fold enhancement), TE/TR = 24/39 ms, and α = 90°. The 1D enhancement factor was measured by calculating the SNR ratio between the Overhauser DNP experiment and the same experiment at thermal equilibrium. Image reconstruction and data processing

[0072] k-space in the phase encode directions was randomly undersampled using a variable density Gaussian pattern to reduce acquisition time. The variable density Gaussian sampling was tuned to emphasize the center of k- space while maintaining acquisition of higher spatial frequencies to prevent image blurring. Random undersampling was used to prevent coherent artifacts in the reconstructed images. The missing lines in the undersampled k-space were filled with zeros; no other processing was added before Fourier transformation.

[0073] The impact of undersampling on the displacement fields was investigated by acquiring MRE data with 25, 50, and 75% sampling was compared with full k-space sampling. Each dataset was corrected for B₀ drift and unwrapped before comparison. The undersampling rate that resulted in the minimum acquisition time with the best motion sensitivity was then used for acquisitions using the fractional encoding approach as described.

[0074] Viscoelastic maps were also calculated. For each dataset, phase images were unwrapped using a Laplacian unwrapping algorithm and corrected for B₀ drift. Phase maps were obtained by subtracting the reference unwrapped phase from the unwrapped phase images acquired for each direction with vibrations turned on. Three-dimensional convolution filtering was applied to the resulting phase maps with a Gaussian kernel using the filter2 function from MATLAB. The curl operator was then applied in order to remove all contribution from compressional waves, leaving pure shear displacements. The amplitude and phase of the wave in each direction were locally assessed after Fourier transformation in the time domain and demodulation at the excitation frequency. The equation of motion was finally inverted to calculate locally the isotropic complex shear modulus G*

[0075] where ρ is the density of the phantom (equal to water density), u the curl of the displacement field, ^² the Laplacian operator, and G* = G_d + iG_l (G_d being the dynamic or storage modulus, and G_l the loss modulus). All data was processed with MATLAB (MathWorks, Natick, MA, USA) scripts written in house. RESULTS

[0076] The calculated mechanical resonances of the system were in excellent agreement with the measured resonances (r² = 0.998). The fourth- order harmonic resonance corresponding to a 103 Hz acoustic wave was chosen as the operational frequency for vibration, f_V. [0077] The simulated approach to steady state of the transverse magnetization showed that the signal rapidly builds up to 25 times that of the thermal equilibrium signal (FIG. 7B). The build-up time constant corresponds to the T₁ relaxation time constant of the sample (200 ms). It was deduced from the Bloch simulations that the signal reached about 90% of its steady-state value after 546 ms, corresponding to 14 echoes.

[0078] The influence of undersampling on unwrapped phase images is shown in FIGs. 4A-B. A line was drawn crossing the entire gel and the corresponding normalized profile was recorded. The periodicity of the propagating wave was the same regardless of the sampling percentage used. However, the peak-to-peak amplitude was about 40% lower for 25% sampling and remained close to what observed for full sampling when sampling 75 or 50% of k-space (a maximum of 10% variation was observed). A 50% undersampling was then chosen for the rest of the study.

[0079] Over the different temporal steps, regions of interest enclosing the entire gel section of each slice were drawn in the phantom images, and the corresponding mean SNR for the entire volume was 39.8 ± 8.7. The error on the phase calculated from mean SNR and Eqn. 4 led to a minimum motion sensitivity of 45 μm. The waves propagated over the entire volume and the total mean peak-to-peak total displacement was 1.47 ± 0.49 mm.

[_{0080] The storage (Gd} ^{) and loss (G} _l ^{) moduli of one slice are shown in FIG.} 5. Mean G_d and G_l and standard deviations across this same slice were calculated to be G_d = 2.3 ± 1.0 kPa and G_l = 1.9 ± 0.9 kPa. DISCUSSION

[0081] Results described herein demonstrate that Overhauser DNP can be used to increase the sensitivity of MRE at ultra-low magnetic field. A total 3D acquisition time for eight temporal steps in 6 min was achieved, providing the opportunity for magnetic susceptibility artifact-free MRE.“Snapshots” of the acoustic wave propagating across a PVA-gel with 5mM TEMPOL were acquired with SNR = 39.8 ± 8.7 and mean total wave amplitude of 1.47 ± 0.49 mm. Displacements were measurable along the entire gel (10 cm), meaning that the transducer system used was reliable in transmitting longitudinal waves through soft materials and in particular tissue with a depth that is commensurate with the size of small animals. Thus, the displacement field map quality is sufficient to consider a transfer to in vivo applications.

[0082] The total acquisition time was 6 minutes at 6.5 mT. By way of comparison, previous MRE work carried out at 1.5 T had a minimum acquisition time of 10 sec for a single temporal step and encoding direction, over a single 10mm thick slice. If one considers that the imaging approach described herein produced an equivalent SNR and spatial resolution, the fastest case described in the previous MRE work with imaging parameters equivalent to those used present (i.e. the same number of temporal steps, slices, phase-encode steps, and 3D motion encoding) would require a 5.5 minutes acquisition even with much higher B₀ (about 230 times higher). The time saving in acquisition comes from the use of fast imaging strategies combined with Overhauser DNP, fractional encoding, and adapted to the ultra-low-field regime. This work reports for the first time MRE imaging at field strengths well below 1.5 T.

[0083] The complex modulus was calculated from MRE data, and the storage modulus was found to be in the same range as described in the literature, although slightly smaller. This may be due to the fact that we are operating at a different vibration frequency (rheometry versus MRE, with storage modulus being frequency dependent), or due to the addition of free radicals to the initial preparation, which might impact the polymerization of PVA in solutions.

[0084] To determine the increase in signal due the Overhauser effect, a single pulse, 1D spectroscopy experiment was performed. The signal magnitude of the gel with ESR was compared with the signal magnitude without ESR. The enhanced signal was 30 times larger than without using the Overhauser effect. Because the ESR irradiation is turned off during acquisition, this represents an upper limit to the signal enhancement during MRE experiments. The simulations showed that the embedded ESR irradiation in the MRE sequence successfully produced a 25-fold enhancement. Here, the steady-state enhancement we attain approaches the maximum obtainable on a 1D single- pulse spectroscopic experiment. This result indicates high polarization efficiency of the embedded ESR scheme in the imaging sequence.

[0085] Operation at more than 20 times lower gradient strength resulted in reduced sensitivity to motion compared to state-of-the-art high-magnetic-field clinical scanners. Nonetheless, the sensitivity obtained was sufficient for hepatic MRE. In order to further increase motion sensitivity, a number of approaches may be taken. For instance, higher amplitude pressure waves could translate to motion sensitivity gains. This can be achieved by improvement in the coupling between the waveguide and the phantom and by using a more rigid waveguide, causing less attenuation of the acoustic wave. In addition, Eqn. 4 shows that the phase sensitivity to displacement is linearly proportional to the gradient amplitude. Thus, increasing the imaging gradient amplitude from its current value of about 1 mT/m to 2 mT/m would double the efficiency. A higher increase in gradient strength is also possible if the main magnetic field is increased as well in order to reduce concomitant field effects, which would result in image distortion. A three to four times higher motion sensitivity could be obtained with B₀ twice as high (13 mT) and gradients of the order of 3–4 mT/m. Such a configuration would permit a threefold increase in SNR due to the increase in field strength, or similarly a 32-fold reduction of the total acquisition time.

[0086] Adaptations to in vivo applications may benefit from a number of improvements. First, coil designs with enhanced filling factor for NMR detection, and more localized ESR transmission to the liver, are envisioned. Second, the coupling between the waveguide and the animal could be optimized according to anatomical constraints. Third, motion compensation for respiration may also be implemented. Finally, issues related to the in vivo reactivity of nitroxide radicals may be alleviated by using the low-toxicity long in vivo half-life triarylmethyl radicals developed for in vivo oximetric imaging. Doses in the literature range from less than 1mM to a few tens of millimolar of free radicals. The intravenous route might be preferable in order to have fast uptake in the liver and reduce the effect of radicals being scavenged by metabolic processes before reaching the targeted organs. The hardware improvements described above are envisioned to improve imaging efficiency, permitting either faster acquisitions or the use of reduced nitroxide concentration. A compromise between acquisition time and free radical concentration would allow a safer transfer to in vivo applications. Overall, low-field MRE could provide a simple and portable system for detection of elasticity changes in subjects with iron overload or implanted devices.

[0087] In summary, it was demonstrated that MRE can be performed in combination with Overhauser DNP at ultra-low magnetic field. Results described may open new perspectives in the diagnosis of chronic liver diseases in subjects with iron overload. At high magnetic field, MRI is currently used to detect and quantify the iron level in these patients, but is not used to perform dynamic functional imaging, as the signal in the liver drops dramatically due to its very short T2*. One-dimensional transient elastography has shown that elasticity is not affected by iron overload to stage fibrosis in patients with hemochromatosis. This result encourages alternative work with MR, which offers, as opposed to traditional ultrasound techniques, three dimensional access to the organ, as well as information linked to wave attenuation, and thus viscosity, without suffering from the presence of water (ascites), fat and/or bones. Because MRE has also shown promising results in screening non-alcoholic steato-hepatitis subjects at risk of developing fibrosis, the present approach has the potential to become a valuable non-invasive alternative to biopsy for patients with chronic liver diseases and iron overload.

[0088] The present invention has been described in terms of one or more embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.

Claims

1. A magnetic resonance imaging (MRI) system configured to perform an imaging process of a subject, the system comprising:

a magnet system configured to generate a static magnetic field about at least a region of interest (ROI) of the subject arranged in the MRI system;

a plurality of gradient coils configured to establish at least one magnetic gradient field with respect to the static magnetic field;

a radio frequency (RF) system configured to deliver excitation pulses to the subject;

a magnetic resonance elastography (MRE) driver configured to deliver an oscillatory stress to the subject; and

a computer system programmed to:

control the plurality of gradient coils, the RF system, and the driver system to perform a three-dimensional (3D) balanced steady-state free precession (b-SSFP) sequence to acquire MRE data from the subject;

perform, during the 3D b-SSFP pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data;

reconstruct, using the MRE data, at least one image of the subject.

2. The system of claim 1, wherein the computer system is further configured to direct an electron spin resonance (ESR) system to apply the ESR pulses during the pulse sequence.

3. The system of claim 1 wherein the computer system is further programmed to apply an ESR pulse during at least one motion encoding gradient (MEG) applied using the plurality of gradient coils.

4. The system of claim 1, wherein the static magnetic field includes a low-field static magnetic field.

5. The system of claim 1, wherein the static magnetic field is less than 10 mT.

6. The system of claim 1, wherein the computer system is further configured to acquire the MRE data using a sampling less than 100%.

7. The system of claim 1, wherein the computer system is further configured to acquire the MRE data using an undersampling scheme based on a variable density Gaussian pattern.

8. The system of claim 1, wherein a frequency of at least one motion encoding gradient applied using the plurality of gradient coils is higher than the frequency of a drive field applied using the MRE driver.

9. The system of claim 1, wherein at least a portion of the ROI comprises one of an iron overload or an implanted device including a metal.

10. The system of claim 1, wherein the MRI system is low-field magnetic resonance imaging (LFMRI) system and the magnetic field is less than 10 mT.

11. The system of claim 1, wherein the at least one image of the subject includes an elastogram.

12. A method for performing a medical imaging process, the method comprising:

arranging a subject in a low-field magnetic resonance imaging (LFMRI) system configured to generate a static magnetic field about at least a region of interest (ROI) of the subject comprising materials capable of producing magnetic susceptibility artifacts;

controlling the LFMRI system to perform a pulse sequence to acquire magnetic resonance elastography (MRE) data from the ROI; performing, during the pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data; and

reconstructing, using the MRE data, at least one image indicative of mechanical properties of tissues in the subject.

13. The method of claim 12, wherein the method further comprises directing an electron spin resonance (ESR) system to apply the ESR pulses during the pulse sequence.

14. The method of claim 12, wherein the method further comprises applying an ESR pulse during at least one motion encoding gradient (MEG) applied using a plurality of gradient coils of the LFMRI system.

15. The method of claim 12, wherein the method further comprises establishing a steady-state wave in the subject with a frequency in a range between 50 to 500 Hz using an MRE driver of the LFMRI system.

16. The method of claim 12, wherein the static magnetic field includes a low-field static magnetic field.

17. The method of claim 16, wherein the static magnetic field is less than 10 mT.

18. The method of claim 12, wherein the method further comprises performing a three-dimensional (3D) balanced steady-state free precession (b- SSFP) sequence.

19. The method of claim 12, wherein the method further comprises acquiring the MRE data using an undersampling scheme based on a variable density Gaussian pattern.

20. The method of claim 12, wherein a frequency of at least one motion encoding gradient applied during the pulse sequence is higher than the frequency of a drive field applied during the pulse sequence.

21. The method of claim 12, wherein at least a portion of the ROI comprises an iron overload or implanted device.

22. A method for performing a medical imaging process, the method comprising:

arranging a subject in a magnetic resonance imaging (MRI) system for imaging at least a region of interest (ROI) comprising materials capable of producing magnetic susceptibility artifacts;

generating, using the magnetic resonance imaging (MRI) system, a static magnetic field configured to minimize the magnetic susceptibility artifacts;

controlling the MRI system to perform a pulse sequence to acquire magnetic resonance elastography (MRE) data from the ROI;

performing, during the pulse sequence, electron spin resonance (ESR) pulses to enhance signals associated with the MRE data; and

reconstructing, using the MRE data, at least one image indicative of mechanical properties of tissues in the subject.

23. The method of claim 21, wherein the static magnetic field is less than 10 mT.

24. The method of claim 21, wherein the ROI includes a liver.