Cross-term spatio-temporal coding (xSPEN) techniques for single-sided MRI

Abstract

The present disclosure provides a single-sided magnetic resonance imaging system comprising a plurality of components. The present disclosure also provides a method for generating a multi-dimensional image by using an imaging system. The present disclosure also provides a process of applying multiple xSPEN readout gradients on the single-sided magnetic resonance imaging system disclosed herein. There are several techniques for collecting dynamic magnetic resonance (MR) images. The fast pulse sequence may be used to fast sample the entire k-space, thereby generating images acquired within a small time window. For example, a spiral trajectory gradient echo sequence may be used to sample the entire k-space within tens of milliseconds. Another method is to use small flip angle excitation to quickly collect lines in k-space. These methods may also be combined with undersampling to collect images more quickly.

Description

[0001] Cross Reference to Related Applications

[0002] This application claims the benefit of U.S. Provisional Application Serial No. 63 / 479,162, filed January 9, 2023, which is incorporated by reference herein in its entirety. BACKGROUND

[0003] The present disclosure relates to medical systems, devices, and methods, particularly for biomedical imaging, including single-sided or open magnetic resonance (MRI) scanners.

[0004] Single-sided or open magnetic resonance imaging (MRI) scanners generally have a permanent or intrinsic gradient magnetic field along a longitudinal axis that extends from the single-sided MRI device into a field of view. The permanent gradient magnetic field can be produced by two sets of gradient coils on the face of a rare earth magnet and a permanent magnet. This orientation allows imaging within a field of view above the face of the magnet. By designing a system with this form factor, imaging can be performed without having to enclose the region being imaged. As such, imaging can be performed without having the patient enter a bore, thereby allowing the scanner to be used with other medical devices, such as biopsy robots, for example. Imaging outside of the imaging bore of a traditional enclosed MRI scanner is also more comfortable for patients with claustrophobia. Single-sided MRI can also be portable and can image any object positioned within the field of view.

[0005] Using surface gradient coils with single-sided scanners, while generally required for single-sided scanning, can result in variations in the field of view along the Z-axis, echo shifts, and / or ultimately k-space truncation, which can result in blurring and effectively limit the image quality obtained by single-sided MRI scanners. Accordingly, there is a need for improved sample acquisition for single-sided or open MRI scanners. SUMMARY

[0006] There are several techniques for collecting dynamic magnetic resonance (MR) images. Fast pulse sequences can be used to sample the entire k-space quickly, generating images acquired within a small time window. For example, a spiral trajectory gradient echo sequence can be used to sample the entire k-space within tens of milliseconds. Another approach is to use small flip angle excitation to quickly collect lines in k-space. These approaches can also be combined with undersampling to collect images even more quickly. However, each of these techniques assumes a strong, uniform magnetic field and tends to break down as the uniformity of the main magnetic field decreases.

[0007] There are also methods for fast acquisition of images in inhomogeneous fields, with the most relevant being xSPEN. Different kinds of xSPEN pulse sequences are described in Zhang, Zhiyong, et al. “Single-Scan MRI with Exceptional Resilience to Field Heterogeneities.” Magnetic Resonance in Medicine 77 (2017): 623-634 (hereinafter “Zhang, Zhiyong, et al.”, which is incorporated by reference herein in its entirety). In some cases, a pulse sequence similar to a subset of the xSPEN sequences disclosed in the supplemental information in Section 5.4 of Zhang, Zhiyong, et al. can be used. The supplemental informal document of Zhang, Zhiyong, et al. is also incorporated by reference herein in its entirety. These sequences allow one to define the imaging axis without being perturbed by permanent gradients.

[0008] In low SNR environments, it can be necessary to signal average or sample k-space with short echo times to improve SNR. Making these adjustments to existing pulse sequences can slow down the image acquisition speed.

[0009] Existing methods are also not designed for inhomogeneous magnetic fields. In a system with permanent gradients, it is not feasible to collect all of k-space with a single spiral. Permanent gradients can limit the trajectories that a scanner can achieve.

[0010] To address these issues, xSPEN can be used for fast imaging with permanent gradients, but the existing version is not optimal for single-sided low-field systems. At low magnetic fields with strong permanent gradients, the SNR is much lower, which limits the applicability of xSPEN.

[0011] In one aspect, disclosed herein is a unilateral magnetic resonance imaging system, comprising: a housing comprising: a front surface; a permanent magnet to provide a static magnetic field extending from the permanent magnet into a region of interest with respect to a first axis, wherein the first axis is perpendicular to the permanent magnet; a radio frequency transmit coil; and a unilateral gradient coil set, wherein the radio frequency transmit coil and the unilateral gradient coil set are positioned proximate to the front surface; a radio frequency receive coil; a power supply, wherein the power supply is configured to cause a current to flow through at least one of the radio frequency transmit coil or the unilateral gradient coil set to generate an electromagnetic field in the region of interest, wherein at least a portion of the region of interest is outside of the front surface; and a control circuit configured to: emit an excitation pulse with the radio frequency transmit coil; emit a first refocusing pulse with the radio frequency transmit coil; during the first refocusing pulse, apply a first xSPEN readout gradient along a second axis that is orthogonal to the first axis with the unilateral gradient coil set; perform phase encoding along a third axis that is orthogonal to the first axis and orthogonal to the second axis with the unilateral gradient coil set; emit a second refocusing pulse with the radio frequency transmit coil; during the second refocusing pulse, apply a second xSPEN readout gradient along the second axis with the unilateral gradient coil set; and receive data with the radio frequency receive coil. In some embodiments, the control circuit is further configured to: emit any number of refocusing pulses with the radio frequency transmit coil before, between, or after refocusing pulses that occur concurrently with the xSPEN readout gradients. In some embodiments, the control circuit is further configured to: generate an encoding matrix for generating an image, wherein the encoding matrix accounts for residual linear phase and residual quadratic phase. In some embodiments, the control circuit is further configured to: emit an excitation pulse with the radio frequency transmit coil; emit a first refocusing pulse with the radio frequency transmit coil; during the first refocusing pulse, apply a first xSPEN readout gradient along a second axis that is orthogonal to the first axis with the unilateral gradient coil set; during the first refocusing pulse, apply a second xSPEN readout gradient along a third axis that is orthogonal to the first axis and orthogonal to the second axis with the unilateral gradient coil set; emit a second refocusing pulse with the radio frequency transmit coil; during the second refocusing pulse, apply a third xSPEN readout gradient along the second axis with the unilateral gradient coil set; during the second refocusing pulse, apply a fourth xSPEN readout gradient along the third axis with the unilateral gradient coil set; and receive data with the radio frequency receive coil. In some embodiments, the control circuit is further configured to: emit any number of refocusing pulses with the radio frequency transmit coil before, between, or after refocusing pulses that occur concurrently with the xSPEN readout gradients.In some embodiments, the control circuit is further configured to: during the first refocusing pulse, emit any number of additional xSPEN readout gradients along any number of additional imaging axes with any number of additional gradient coil sets; during the second refocusing pulse, emit any number of additional xSPEN readout gradients along any number of additional imaging axes with any number of additional gradient coil sets; and emit the fourth refocusing pulse with the radio frequency transmit coil. In some embodiments, the amplitude of the xSPEN readout axis is not uniform in time during the refocusing pulse. In some embodiments, the gradient shape of the xSPEN readout axis is not constant in time during the refocusing pulse. In some embodiments, the control circuit is further configured to generate an encoding matrix for generating the image, where the encoding matrix accounts for residual linear phase and residual quadratic phase. In some embodiments, the control circuit is configured to spatially encode along arbitrary axes during acquisition. In some embodiments, the arbitrary axes remain fixed during a single image acquisition. In some embodiments, the arbitrary axes are rotated during a single image acquisition.

[0012] In another aspect, disclosed herein is a method for generating a two-dimensional image by collecting a series of projections with different angles, and then generating the two-dimensional image by reconstructing the two-dimensional image.

[0013] In another aspect, disclosed herein is a method for generating a three-dimensional image by adding an additional gradient pulse array to a pulse sequence. In some embodiments, the additional gradient pulse array is parallel to the xSPEN readout axis, and provides spatial information along a permanent gradient axis to be resolved.

[0014] In another aspect, disclosed herein is a method for generating a three-dimensional image with a radial encoding scheme, where the effective axis of the gradient array is rotated together with the axis of the xSPEN readout gradient.

[0015] In another aspect, disclosed herein is a process of applying a plurality of xSPEN readout gradients on a unilateral magnetic resonance imaging system disclosed herein. In some embodiments, applying the plurality of xSPEN readout gradients generates an encoding matrix for generating an image, where the encoding matrix accounts for residual linear phase and residual quadratic phase. In some embodiments, the process further includes applying, with the unilateral gradient coil set, a first xSPEN readout gradient along a second axis that is orthogonal to the first axis during a first refocusing pulse. In some embodiments, the process further includes applying, with the unilateral gradient coil set, a second xSPEN readout gradient along the second axis during a second refocusing pulse. In some embodiments, the process further includes applying, with the unilateral gradient coil set, a third xSPEN readout gradient along a third axis during the first refocusing pulse. In some embodiments, the process further includes applying, with the unilateral gradient coil set, a fourth xSPEN readout gradient along the third axis during the second refocusing pulse.

[0016] Incorporation by Reference

[0017] All publications, patents, and patent applications mentioned in this specification are herein incorporated by reference to the same extent as if each individual publication, patent, or patent application was specifically and individually indicated to be incorporated by reference. To the extent that any publication, patent, or patent application incorporated by reference contradicts any disclosure contained in the specification, the specification hereby takes precedence over any such incorporated publication, patent, or patent application. BRIEF DESCRIPTION OF DRAWINGS

[0018] The novel features of the aspects are set forth with particularity in the appended claims. Nevertheless, the described aspects, both as to organization and method of operation, can be best understood by reference to the following description and the accompanying drawings.

[0019] Figure 1 is a perspective view of a magnetic resonance imaging (MRI) scanner in accordance with various aspects of the present disclosure.

[0020] Figure 2 is a perspective view of a magnetic resonance imaging (MRI) scanner in accordance with various aspects of the present disclosure. Figure 1 is an exploded perspective view of the MRI scanner of

[0021] Figure 3 is a perspective view of a magnetic resonance imaging (MRI) scanner in accordance with various aspects of the present disclosure. Figure 1 is an elevation view of the MRI scanner of

[0022] Figure 4 is an elevation view of the MRI scanner of Figure 1 is an elevation view of the MRI scanner of

[0023] Figure 5is a perspective view of a permanent magnet assembly of an MRI scanner in accordance with various aspects of the present disclosure Figure 1 is a perspective view of a permanent magnet assembly of an MRI scanner in accordance with various aspects of the present disclosure

[0024] Figure 6 is a perspective view of a permanent magnet assembly of an MRI scanner in accordance with various aspects of the present disclosure Figure 1 is an isometric view of a gradient coil set and permanent magnet assembly of an MRI system shown in

[0025] Figure 7 is a control schematic of a single-sided MRI system in accordance with various aspects of the present disclosure

[0026] Figure 8 is a schematic of magnetic gradients along the z-axis in accordance with various aspects of the present disclosure

[0027] Figure 9 is a representative plot of a scan frequency pulse in accordance with various aspects of the present disclosure

[0028] Figure 10 is a plot of a cross-term spatiotemporal encoding pulse sequence for fast imaging with single-sided MRI

[0029] Figure 11 is a plot of a cross-term spatiotemporal encoding pulse sequence for radial imaging with single-sided MRI

[0030] Figure 12 (A) shows a single-sided prostate MRI scanner. The subject sits in the lithotomy position and there is a hole in the magnet for surgery and biopsy of the prostate. Figure 12 (B) shows the xSPEN imaging geometry. Slice selection is performed along the z-dimension of the magnet and both in-plane (x / y) dimensions are typically phase encoded.

[0031] Figure 13 illustrates a phase-encoded xSPEN imaging pulse sequence.

[0032] Figure 14 shows the xSPEN spatiotemporal encoding function moving along the y-direction during a 1.5 ms long readout window in the center slice of an in-vivo scan. The FOV here is 18 x 18 cm and the x-dimension is phase encoded.

[0033] Figure 15 (A) and Figure 15 (B) illustrates the reconstruction of the center slice of a 10 cm diameter ACR phantom. Figure 15 (A) is obtained via the Fourier reconstruction described earlier. Figure 15 (B) is obtained with the described model-based reconstruction that fully accounts for gradient nonlinearities.

[0034] Figure 16 (A) illustrates a single-sided MRI scanner. Figure 16 (B) illustrates a radial xSPEN pulse sequence diagram.

[0035] Figure 17 (A) shows numerically calculated radial xSPEN point spread functions at readout start, in the middle of readout, and at readout end time points. Figure 17 (B) shows middle readout PSFs for 0, 45, and 90 degree radial angles.

[0036] Figure 18 (A) shows simulated (sinc) radial xSPEN PSFs at three different angles in the middle of their readout. Figure 18 (B) shows simulated 45 degree sinc radial xSPEN PSFs at three different time points in their readout. The sinc sweeps from the upper left to the lower right of the FOV. Figure 18 (C) shows phantom image reconstruction with all projections, ¼ projections, and half readout duration.

[0037] Figure 19 (A) shows xSPEN sinograms of a resolution phantom. The data is in the form of an x-ray sinogram and has no k-space dimension. Figure 19 (B) are images reconstructed from the sinogram data with three methods, one is an inverse radon transform that ignores the gradient nonlinearity in its initial Fourier transform step, one is a reconstruction using an analytical sinc function that does not account for Gz nonlinearity, and one is using Figure 17 (a) and Figure 17 (B) show the reconstruction of the numerical PSF method.

[0038] Figure 20 Images of an ACR extremity phantom collected with different fields of view / readout durations are shown.

[0039] Figure 21 (A) illustrates a single-sided MRI scanner. Figure 21 (B) illustrates a three-dimensional xSPEN pulse sequence that is based on a CPMG-RARE scan with WURST excitation and refocusing pulses and uses CHORUS technology to place the first spin echo after the second refocusing pulse.

[0040] Figure 22 (A) shows cross sections of the bilinear xSPEN spatial encoding phase function through the slab (z) and the in-plane dimension of xSPEN encoding (x) for different x-phase encodings that move the saddle point / sensitivity point to different z-locations through the slab. Figure 22(B) shows cross-sections of the phase function at different time points in the readout (without x-phase encoding), where the saddle / sensitive point is shifted to different in-plane (x) positions.

[0041] Figure 23 (A) shows high-resolution xSPEN spatially encoded bilinear phase maps (top; cross-sections through the slices) and PSF for each slice (bottom). Figure 23 (B) illustrates the forward model of associating a reconstructed image with the data by applying the point spread function for each slice and each readout time point to the input image stack, then computing a type III NUFFT to apply the phase encoding.

[0042] Figure 24 Twelve slices from two consecutive slabs of an ACR phantom are shown, reconstructed from 3D xSPEN data. The images show a clear evolution of features between the slices, reflecting the ability of the sequence to resolve details between sub-slices within a slab.

[0043] The drawings are not intended to be to scale. In several views, corresponding reference characters indicate corresponding parts. For purposes of clarity not every component can be labeled in every drawing. Examples listed herein illustrate certain implementations of the disclosure in one form and such examples should not be construed as limiting the scope of the disclosure in any way. DETAILED DESCRIPTION

[0044] The following international patent applications are incorporated by reference herein in their respective entireties:

[0045] • International Application No. PCT / US2020 / 018352, filed February 14, 2020, entitled “SYSTEMS AND METHODS FOR ULTRALOW FIELD RELAXATION DISPERSION,” published as International Publication No. WO 2020 / 168233;

[0046] • International Application No. PCT / US2020 / 019530, filed February 24, 2020, entitled “SYSTEMS AND METHODS FOR PERFORMING MAGNETIC RESONANCE IMAGING,” published as International Publication No. WO 2020 / 172673;

[0047] • International Application No. PCT / US2020 / 019524, filed February 24, 2020, entitled “PSEUDO-BIRDCAGE COIL WITH VARIABLE TUNING AND APPLICATIONS THEREOF,” published as International Publication No. WO 2020 / 172672;

[0048] • International Application No. PCT / US2020 / 024776, filed March 25, 2020, entitled “SINGLE-SIDED FAST MRI GRADIENT FIELD COILS AND APPLICATIONS THEREOF,” published as International Publication No. WO 2020 / 198395;

[0049] • International Application No. PCT / US2020 / 024778, filed March 25, 2020, entitled “SYSTEMS AND METHODS FOR VOLUMETRIC ACQUISITION IN A SINGLE-SIDED MRI SYSTEM,” published as International Publication No. WO 2020 / 198396; • International Application No. PCT / US2020 / 039667, filed June 25, 2020, entitled “SYSTEMS AND METHODS FOR IMAGE RECONSTRUCTIONS IN MAGNETIC RESONANCE IMAGING,” published as International Publication No. WO 2020 / 264194;

[0050] • International Application No. PCT / US2021 / 014628, filed January 22, 2021, entitled “MRI-GUIDED ROBOTIC SYSTEMS AND METHODS FOR BIOPSY”; and

[0051] • International Application No. PCT / US2021 / 018834, filed February 19, 2021, entitled “RADIO FREQUENCY RECEPTION COIL NETWORKS FOR SINGLE-SIDED MAGNETIC RESONANCE IMAGING”;

[0052] • International Patent Application No. PCT / US2021 / 021464, filed March 9, 2021, entitled “PHASE ENCODING WITH FREQUENCY SWEEP PULSES FOR MAGNETIC RESONANCE IMAGING IN INHOMOGENEOUS MAGNETIC FIELDS”;

[0053] • International Patent Application No. PCT / US2021 / 021461, filed March 9, 2021, entitled “PULSE SEQUENCES AND FREQUENCY SWEEP PULSES FOR SINGLE-SIDED MAGNETIC RESONANCE IMAGING”;

[0054] • International Patent Application No. PCT / US2022 / 071924, filed April 26, 2022, entitled “INTERVENTIONAL LOCALIZATION GUIDE AND METHOD FOR MRI GUIDED PELVIC INTERVENTIONS”; and

[0055] • International Patent Application No. PCT / US2022 / 082551, filed December 29, 2022, entitled “RELAXATION-BASED MAGNETIC RESONANCE THERMOMETRY WITH A LOW-FIELD SINGLE-SIDED MRI SCANNER”.

[0056] U.S. Patent Application Publication No. 2018 / 0356480, entitled “UNILATERAL MAGNETIC RESONANCE IMAGING SYSTEM WITH APERTURE FOR INTERVENTIONS AND METHODOLOGIES FOR OPERATING SAME,” published December 13, 2018, is also incorporated by reference herein in its entirety.

[0057] It should be noted that the illustrative examples can be used in other embodiments or circumstances and that the phraseology or terminology employed herein, and not the examples given, is selected and used merely for convenience and illustrative purposes and is not intended to limit the scope of the inventive examples. Additionally, it is intended that the examples described herein can be combined in different ways with each other and / or with other examples described herein without departing from the scope of the inventive examples.

[0058] According to various aspects, an MRI system can be provided that can include a unique imaging region that can be offset from the face of the magnet. Such an offset and unilateral MRI system is less restrictive compared to traditional MRI scanners. Further, the form factor can have a built-in or inherent magnetic field gradient that creates a range of magnetic field values over the region of interest. In other words, the inherent magnetic field can be non-uniform and can be active at all times. The non-uniformity of the magnetic field strength in the region of interest for the unilateral MRI system can exceed 200 parts per million (200 ppm). For example, the non-uniformity of the magnetic field strength in the region of interest for the unilateral MRI system can be between 200 ppm and 200,000 ppm. In various aspects of the present disclosure, the non-uniformity in the region of interest can be greater than 1,000 ppm and can be greater than 10,000 ppm. In one instance, the non-uniformity in the region of interest can be 81,000 ppm.

[0059] The inherent magnetic field gradient can be generated by a permanent magnet within the MRI scanner. For example, the magnetic field strength in the region of interest for the unilateral MRI system can be less than 1 Tesla (T). For example, the magnetic field strength in the region of interest for the unilateral MRI system can be less than 0.5 T. In other instances, for example, the magnetic field strength can be greater than 1 T and can be 1.5 T. Compared to typical MRI systems, the system can operate at lower magnetic field strengths, allowing for relaxed constraints on gradient and / or radio frequency coil design and / or allowing for additional mechanisms (e.g., robots) to be used with the MRI scanner. Exemplary MRI-guided robotic systems are further described in International Application No. PCT / US2021 / 014628, entitled “MRI-GUIDED ROBOTIC SYSTEMS AND METHODS FOR BIOPSY,” filed January 22, 2021.

[0060] Figures 1-6 An MRI scanner 100 and its components are depicted. As Figure 1 and Figure 2The MRI scanner 100 includes a housing 120 having a recessed and concave front or anterior surface 125. In other aspects, the front of the housing 120 can be flat and planar. The anterior surface 125 can face the object being imaged by the MRI scanner. As shown in Figure 1 and Figure 2 The housing 120 includes a permanent magnet assembly 130, an RF transmit coil (TX) 140, a gradient coil set 150, an electromagnet 160, and an RF receive coil (RX) 170, as shown in

[0061] Referring primarily to Figures 3-5 The permanent magnet assembly 130 includes a magnet array. The magnet array forming the permanent magnet assembly 130 is configured to cover the anterior surface 125 or patient-facing surface of the MRI scanner 100 (see Figure 3 ) and is shown in Figure 4 as horizontal bars. The permanent magnet assembly 130 includes a plurality of cylindrically shaped permanent magnets configured in parallel. Referring primarily to Figure 5 The permanent magnet assembly 130 includes parallel plates 132 held together by a bracket 134. The system can be attached to the housing 120 of the MRI scanner 100 at the bracket 136. There can be a plurality of holes 138 in the parallel plates 132. For example, the permanent magnet assembly 130 can include any suitable magnetic material including, but not limited to, rare earth-based magnetic materials such as exemplified by neodymium-based magnetic materials.

[0062] The permanent magnet assembly 130 defines an entry hole or aperture 135 that can pass through the housing 120 from an opposite side of the housing 120 to the patient. In other aspects of the disclosure, the permanent magnet array forming the permanent magnet assembly in the housing 120 can be holeless and define an uninterrupted or continuous arrangement of permanent magnets without being defined as a hole through the permanent magnets. In other cases, the permanent magnet array in the housing 120 can form more than one hole / entry hole through the permanent magnet array.

[0063] According to various aspects of the disclosure, the permanent magnet assembly 130 provides a magnetic field B0in a region of interest 190 along the Z-axis, as shown in Figure 1The Z-axis is perpendicular to the permanent magnet assembly 130. In other words, the Z-axis extends from the center of the permanent magnet assembly 130 and defines the direction of the magnetic field B0 away from the front face of the permanent magnet assembly 130. The Z-axis can define the main magnetic field B0 direction. The main magnetic field B0 can decrease along the Z-axis away from the front face of the permanent magnet assembly 130 (i.e., intrinsic gradient) and in the direction indicated by the arrow in Figure 1

[0064] In one aspect, the non-uniformity of the magnetic field in the region of interest 190 for the permanent magnet assembly 130 can be approximately 81,000 ppm. In another aspect, the non-uniformity of the magnetic field strength in the region of interest 190 for the permanent magnet assembly 130 can be between 200 ppm to 200,000 ppm, and can be greater than 1,000 ppm in some cases, and greater than 10,000 ppm in various cases.

[0065] In one aspect, the magnetic field strength of the permanent magnet assembly 130 can be less than 1 T. In another aspect, the magnetic field strength of the permanent magnet assembly 130 can be less than 0.5 T. In other cases, the magnetic field strength of the permanent magnet assembly 130 can be greater than 1 T, and for example, can be 1.5 T. Reference is made primarily to Figure 1 The Y-axis extends upward and downward from the Z-axis, and the X-axis extends left and right from the Z-axis. The X-axis, Y-axis, and Z-axis are all orthogonal to each other, and the positive direction of each axis is indicated by the respective arrow in Figure 1

[0066] The RF transmit coil 140 is configured to transmit RF waveforms and associated electromagnetic fields. The RF pulses from the RF transmit coil 140 are configured to rotate the magnetization produced by the permanent magnet 130 by generating an effective magnetic field called B1, which is orthogonal to the direction of the permanent field (e.g., orthogonal plane).

[0067] Reference is made primarily to Figure 3 The gradient coil set 150 includes two sets of gradient coils 152, 154. The sets of gradient coils 152, 154 are positioned on the front or anterior surface 125 of the permanent magnet assembly 130, in the middle of the permanent magnet assembly 130 and the region of interest 190. Each set of gradient coils 152, 154 includes coil portions on opposite sides of the bore 135. Reference is made to the axes in Figure 1

[0068] Reference is now made to Figure 7 ​​​, shows a control schematic diagram for a unilateral MRI system 300. In various aspects of the present disclosure, the unilateral MRI scanner 100 and / or its components ( Figures 1-6 ) can be incorporated into the MRI system 300. For example, the imaging system 300 includes a permanent magnet assembly 308, which can be similar in various cases to the permanent magnet assembly 130 (see Figures 2-5 ). The imaging system 300 further includes an RF transmit coil 310, which may be similar to the RF transmit coil 140 (see Figure 3 ). In addition, the imaging system 300 includes an RF receiving coil 314, which may be similar to the RF receiving coil 170 (see Figure 3 In various aspects, the RF transmit coil 310 and / or the RF receive coil can also be positioned in the housing of the MRI scanner, and in some cases, the RF transmit coil 310 and the RF receive coil 314 can be combined into an integrated Tx / Rx coil. The system 300 also includes a gradient coil 320 configured to generate a gradient field to facilitate imaging of an object in the field of view 312.

[0069] The unilateral MRI system 300 also includes a computer 302 in signal communication with a spectrometer 304 and configured to send and receive signals between the computer 302 and the spectrometer 304 .

[0070] The main magnetic field B0 generated by the permanent magnet 308 extends away from the permanent magnet 308 and away from the RF transmit coil 310 into a field of view 312. The field of view 312 contains the object being imaged by the MRI system 300.

[0071] During the imaging process, the main magnetic field B0 extends into the field of view 312. The direction of the effective magnetic field B1 changes in response to RF pulses and associated electromagnetic fields from the RF transmit coil 310. For example, the RF transmit coil 310 is configured to selectively transmit RF signals or pulses to objects, such as tissue, in the field of view. These RF pulses change the effective magnetic field experienced by spins in the sample (e.g., patient tissue). The RF pulses change the effective magnetic field experienced by magnetization in a rotating coordinate system. When the RF pulse is in a resonant state, the effective magnetic field is only along the axis of the RF pulse. When in a non-resonant state, the axis of the effective field will be located between the applied RF pulse and the static magnetic field. The RF pulses can be, for example, chirped or frequency-swept pulses, as further described herein.

[0072] Further, when an object in the field of view 312 is excited with RF pulses from the RF transmit coil 310, the precession of the object results in an induced or MR current that is detected by the RF receive coil 314. The RF receive coil 314 can send the excitation data to the RF preamplifier 316. The RF preamplifier 316 can boost or amplify the excitation data signals and send them to the spectrometer 304. The spectrometer 304 can send the excitation data to the computer 302 for storage, analysis, and image construction. For example, the computer 302 can combine multiple stored excitation data signals to create an image.

[0073] From the spectrometer 304, the signals can also be relayed via the RF power amplifier 306 to the RF transmit coil 310 and via the gradient power amplifier 318 to the gradient coil 320. The RF power amplifier 306 amplifies the signals and sends them to the RF transmit coil 310. The gradient power amplifier 318 amplifies the gradient coil signals and sends them to the gradient coil 320.

[0074] For example, described herein are systems and methods for efficiently collecting nuclear magnetic resonance spectra and magnetic resonance images in a non-uniform field, such as using the single-sided MRI scanner 100 and system 300.

[0075] Imaging using single-sided or open MRI presents many challenges. Typically, the two sets of gradient coils (see FIG. 1) in a single-sided system are placed on the face of the permanent magnet assembly. As a result, the magnitude of the gradient will drop as one moves away from the face of the permanent magnet assembly. Thus, for a given phase encoding array, the field of view will change as one moves along the axis of the permanent field B0. In other words, the pulsed gradient coils in a single-sided scanner have a small component along the direction of the permanent gradient. Figure 6

[0076] Figure 8 is a schematic 500 of the magnetic field gradient along the Z axis of the MRI scanner 100. The permanent magnet 130 has an inherent gradient along the Z axis. The strength of the Z gradient decreases as one moves away from the permanent magnet 130. As can be seen in the schematic, the Z gradient curves as one moves away from the permanent magnet, resulting in a decrease in the strength of the gradient. The MRI scanner 100 images multiple slices to create a slab. Each slice is excited at a different frequency to be imaged. The lower frequency excites the tissue in the slice further away from the permanent magnet, and the higher frequency excites the tissue in the slice closer to the magnet. In the schematic, the slab or axial image is made from multiple slices from slice 0 to slice n n where f0is less than f n ​​It is worth noting that a single-sided MRI system can be constructed in such a way that there is a region in which the gradient is opposite to that described and the frequency can increase along the Z-axis.

[0077] Figure 9 A representative plot 900 of a sweep frequency pulse or a chirp pulse (shown demodulated to baseband frequency) is shown, with the sweep direction set from low to high. A chirp excitation pulse with the sweep direction set from low to high is an example of a frequency sweep excitation pulse. The frequency of the chirp pulse with the sweep direction set from low to high starts at a low frequency and the frequency increases over time for the duration of the pulse. The pulse can start at the lowest desired frequency and end once the maximum desired frequency is reached. The pulse frequency in plot 900 can be a negative to positive frequency offset from the baseband frequency. In other words, the frequency sweeps from negative to positive plus the baseband frequency. For example, for a frequency sweep of + / - 100 KHz, the sweep is from the baseband frequency minus 100 KHz to the baseband frequency plus 100 KHz. It is worth noting that the frequency direction can be reversed and sweep from positive to negative frequencies.

[0078] The frequency of the chirp pulse can vary from a minimum (lowest) desired frequency to a maximum (highest) desired frequency. The sweep rate of the pulse is the difference between the highest frequency and the lowest frequency in the pulse divided by the time required to pass between the highest frequency and the lowest frequency. In one aspect, the frequency range covered by the sweep frequency pulse can be from -20 KHz to 20 KHz, a 40 KHz range, with the center frequency varying from board to board. For example, the board can be centered at 2.62 MHz, 2.75 MHz, 2.65 MHz, 2.72 MHz, 2.79 MHz, 2.69 MHz, etc. For a board centered at 2.62 MHz, the chirp pulse would sweep from 2.60 MHz to 2.64 MHz, a 40 KHz range. In other aspects of the disclosure, bandwidths as low as 10 KHz up to 200 KHz can be used in the frequency sweep pulse. Further, in various cases, the sweep range can be less than 40 KHz.

[0079] Cross-term spatiotemporal encoding (xSPEN) technique for fast imaging with single-sided MRI

[0080] Reference Figure 10 The present disclosure discloses a version of xSPEN that can improve SNR and allow for fast three-dimensional image acquisition. In some embodiments, imaging with xSPEN allows for fast sampling of one of the axes in the image, significantly shortening the scan time. Adapting xSPEN to a single-sided low-field system requires some changes to the design of the pulse sequence.

[0081] First, the fast sequence can include a chirped excitation pulse to collect thicker slabs. By collecting thicker slabs, the SNR can be increased because a larger volume and thus more magnetization contributes to the signal. The traditional xSPEN sequence uses a hard pulse to excite the magnetization because it was designed for systems with much weaker gradients. Under these conditions, the hard pulse is able to excite a thick enough slab to generate sufficient SNR. In some embodiments, the pulse sequence can be modified by using timing to eliminate the quadratic phase imparted on the magnetization by the chirped excitation. See Foroozandeh, Mohammadali, et al. “Improved ultra-broadband chirp excitation.” Journal of Magnetic Resonance 302 (2019): 28-33, and Power, J. E., et al. “Increasing the quantitative bandwidth of NMR measurements.” The Royal Society of Chemistry, Chem. Commun. 52 (2016): 2916-2919, each of which is incorporated by reference herein in its entirety.

[0082] In some embodiments, to further increase the SNR, multiple echoes can be collected with the xSPEN sequence. As with other pulse sequences that use chirped refocusing pulses, all other echoes can be discarded because they retain the quadratic phase imparted on them by the chirped pulse. By collecting multiple echoes per acquisition, the present applicant shows how signal averaging can be performed without having to wait for the magnetization to return to the longitudinal axis.

[0083] In another implementation, each echo can be used to sample a different line in k-space. Before collecting the echoes, the magnetization can be moved to a different point in k-space using gradients, allowing multiple k-space lines to be collected during excitation, speeding up the image acquisition.

[0084] Some embodiments of the fast xSPEN sequence described herein can contain interleaved slabs to collect three-dimensional information. These scans can be designed to monitor changes occurring in all three dimensions of space, such as the movement of a biopsy needle within a patient or the formation of an ice ball during a cryoablation experiment, so they must generate volumetric images. In some embodiments, a volume is generated by collecting a series of two-dimensional images in the x / y plane along the z axis. The images can then be interpolated along the z axis to obtain a complete volume.

[0085] Another embodiment of the fast xSPEN sequence described herein can include a phase encoded third axis. This allows for the acquisition of thick excitation profiles to take advantage of the increased SNR benefit while encoding image information on that axis. The image is shown in Figure 11

[0086] In another aspect, another way in which a three-dimensional volume can be collected with xSPEN is to use a radial encoding scheme paired with a gradient array. By collecting projections along different angles until enough projections are collected to reconstruct a two-dimensional image along two axes. The third dimension is encoded by adding a gradient pulse array to the two-dimensional radial xSPEN procedure. The direction of the gradient pulse array is rotated with the angle of the projection so that a plane along the third axis and the angle of the projection axis is collected. By collecting multiple angles with this method, a three-dimensional image can be reconstructed. This method produces an MRI image that is not Fourier encoded, although Fourier phase encoding can be additionally applied along the rotating xSPEN encoding axis to increase spatial resolution without the traditional xSPEN resolution SNR tradeoff.

[0087] Parameter Range

[0088] The potential range for field inhomogeneity is:

[0089] 10 to 300 kHz over a 160 mm depth.

[0090] The potential range for excitation bandwidth is: 5 to 200 kHz.

[0091] The potential range for refocusing bandwidth is: 5 to 200 kHz.

[0092] The potential range for time to acquire an image is: 1 second to 30 minutes.

[0093] ​In another aspect, the present application provides Cartesian xSPEN sequences. In Cartesian xSPEN sequences, one axis is spatiotemporally encoded while the other axis is Fourier encoded. Because of how the spatial information is encoded with xSPEN, it is possible to encode a field of view that is smaller than the size of the object being imaged along the spatiotemporal axis. In a traditional MRI scan, where both axes are Fourier encoded, choosing a field of view that is smaller than the object results in aliasing. Because xSPEN uses a different encoding method, it is possible to encode a smaller field of view, thus allowing the user to zoom in on one axis.

[0094] In some embodiments, the present application discloses a method that allows zooming in along both axes, thus cropping the field of view by using only the readout gradient and not other tools such as saturation bands. Typically, only the spatiotemporal axis can have a field of view that is smaller than the size of the object. In applications where only a portion of the field of view has useful anatomical information, it would be ideal if the field of view could be cropped on both axes so that the image is zoomed in on the relevant anatomy.

[0095] Reference is made to Figure 11 A radial xSPEN pulse sequence generates an image by collecting a series of projections with a rotating axis. These different projections are then combined into a single two-dimensional image using an iterative conjugate gradient reconstruction or filtered back-reconstruction method. Unlike the xSPEN sequence described in the literature, the radial xSPEN sequence has gradients applied during the first two chirped refocusing pulses of both axes so that neither axis is Fourier encoded. By adjusting the relative strength of the gradient pulses applied on each axis, the axis of the projection can be changed. When the gradient is applied to only one axis, the projection will be along that axis only. If some power is applied to both axes, the axis of the projection will lie between the two axes. By imaging in this way, a two-dimensional image that is fully spatiotemporally encoded can be collected. By excluding Fourier encoding from the 2D image acquisition, an image with a field of view that is smaller than the object can be collected without aliasing. In some embodiments, this reduction in field of view is achieved by truncating the readout axis, which can result in a shorter echo time and higher signal strength due to the reduction in transverse relaxation.

[0096] In another embodiment, the geometric center of this reduced field of view can be moved by adjusting the schedule of readout angles between the two axes to be non-uniform sampling. The schedule of angles and readout truncation can also be adjusted to create an elliptical field of view.

[0097] In another embodiment, a small amount of Fourier encoding perpendicular to the direction of xSPEN band travel can be added to resolve the reduced field of view on the xSPEN projections more precisely. One example of this encoding scheme is to apply a rotating Fourier phase encoding along each rotated xSPEN projection to increase the spatial resolution.

[0098] The third dimension for volumetric encoding can be implemented in a number of ways. A first method would collect an array of 2D images sliced along the third dimension, these 2D images comprising the volume. These slices would be acquired by changing the excitation frequency and band or by adding slice-selective electromagnetic gradient pulses during excitation that shift the Larmor frequency of the desired volume subset into the desired range. In some embodiments, a second method can add a third xSPEN readout axis using additional electromagnetic gradients. In this way, the image projection can be tilted in the third dimension, allowing for encoding in the third imaging axis. In some embodiments, a third method can excite the entire desired volume and implement a phase-encoding axis in the third dimension.

[0099] The instantaneous amplitude of the xSPEN readout axis adjusts how fast the point spread function moves across the image. By adjusting the amplitude and thus the speed, one can use it to preferentially increase the SNR of segments along the readout projection. Additionally, if the xSPEN readout axis gradient shape is dynamically changed (e.g., by adding another controllable readout axis), one can dynamically adjust the point spread function size and one can change the pixel resolution across the image.

[0100] Range of parameters

[0101] The potential range for field inhomogeneity is:

[0102] 10 to 300 kHz over a 160 mm depth.

[0103] The potential range for excitation bandwidth is: 5 to 200 kHz.

[0104] The potential range for refocusing bandwidth is: 5 to 200 kHz.

[0105] The potential range for time to acquire an image is: 1 second to 30 minutes.

[0106] Conjugate gradient least squares reconstruction for cross-term spatiotemporal encoding (xSPEN) with unilateral MRI

[0107] For Cartesian xSPEN reconstruction, the gradient fields used for image encoding are linear. When the gradient fields are linear, converting xSPEN data to an image can be as simple as taking the absolute value of the resulting image after Fourier transforming along the Fourier encoding direction.

[0108] For radial xSPEN, it is not yet known how to reconstruct data collected using a radial xSPEN sequence. The data should be able to be converted to an image using similar techniques to those used to reconstruct x-ray images; however, this has not yet been reported.

[0109] In this regard, none of the existing reconstruction methods are able to handle non-linear gradient fields. If the fields are non-linear, the images generated with existing reconstruction can be distorted. Distortions must not be present in the images if they are to be used in clinical diagnosis.

[0110] In one aspect, the present application describes a conjugate gradient least squares reconstruction that uses a map of the gradient fields to correct for distortions caused by the non-linearity of the gradient coils. The method is based on the signal equation for the magnetization generated with an xSPEN sequence, for example, as described with respect to Figure 10 Using this equation, the different contributions to the phase can be calculated and used to solve for an image without distortions. The method can also be used to reconstruct radial xSPEN images, for example, as described with respect to Figure 11

[0111] In some aspects, the reconstruction needs to consider two parts: an ideal non-linear xSPEN reconstruction and a residual phase term. The ideal non-linear reconstruction can be based on equation one. The time domain signal collected is a one-dimensional distribution along the time-space direction spatially encoded with the xSPEN readout gradient applied during the first two refocusing pulses. In this example, the time-space axis is along y, but it can be along any direction. In the equations below, it is assumed that the other axis is Fourier encoded. The encoding matrix is calculated using these equations. The matrix can then be inverted to recover an undistorted image.

[0112] (1) S(t): Time domain signal acquired with the scannerp(x,y): Spatial distribution of the signal. This is the image recovered with the reconstruction

[0113] G x (x,y): Magnetic field generated by the x gradient coil

[0114] PE: Integral of the phase encoding gradient for this particular line

[0115] BW ex: : Bandwidth of the excitation pulse used in the pulse sequence

[0116] G z (x,y): Magnetic field generated by the permanent magnet array

[0117] B: Integral of the xSPEN readout gradient

[0118] Y: Gyromagnetic ratio

[0119] G y (x,y): Magnetic field generated by the y gradient coil

[0120] T: Time in the acquisition window

[0121] T aDuration of acquisition window

[0122] Equation (1) is accurate if the length of the first two chirped refocusing pulses is equal and if the gradients applied during these pulses are identical. However, if the pulses used for excitation are also chirped pulses, the duration of the subsequent two pulses will be different. Their duration difference can depend on the parameters of the pulse sequence. The presence of the difference leads to a residual quadratic and linear phase, which can affect the image quality, especially in radial reconstruction, which can not generate an image at all if the residual phase terms are not taken into account.

[0123] The two residual phase terms are the quadratic term shown in equation (2) and the linear term shown in equation (3). The residual phase terms are multiplied into equation (1), which is then used to generate the encoding matrix, which can be inverted to solve for the image.

[0124] (2)

[0125] CF: CHORUS factor modifying the duration of the first refocusing pulse and the excitation pulse

[0126] T π : Duration of refocusing pulse

[0127] BW: Bandwidth of refocusing pulse

[0128] G z : Magnetic field produced by permanent magnet array

[0129] P1: Intensity of xSPEN readout pulse applied during the first refocusing pulse

[0130] P2: Intensity of xSPEN readout pulse applied during the second refocusing pulse

[0131] Y: Gyromagnetic ratio

[0132] G y : Magnetic field produced by y gradient coil

[0133] (3) φ 线性 = 2T π (CF-1) YG z + 2T π (P1CF+P2) YG y

[0134] Main reference Figure 12 (B), xSPEN is applied to exchange the matrix dimension that is usually frequency encoded but smaller (z) with the matrix dimension that is usually phase encoded but larger (y) to achieve faster imaging.

[0135] Referring primarily to Figure 13 , a chirped WURST pulse is used throughout to excite and refocus the slice's wide bandwidth. The length of the first refocusing pulse is 1.1 times the length of the subsequent refocusing pulses to compensate for the quadratic phase of the subsequent pulses and the quadratic phase of the excitation pulse to produce a spin echo signal after the second refocusing pulse. This requires increasing the amplitude of the second Gy pulse to maintain the same amplitude-pulse duration product that sets the xSPEN bi-linear curvature.

[0136] Referring primarily to Figure 15 (B), the model-based reconstruction has significantly lower geometric distortion and more uniform signal within the phantom.

[0137] Referring primarily to Figure 16 (B), this sequence is a CPMG acquisition with CHORUS method to generate a spin echo after the second WURST refocusing pulse, and the part in parentheses is repeated 12 times. The xSPEN encoding gradients are applied with opposite polarity during the first two refocusing pulses; different colors represent different in-plane radial angles.

[0138] Referring primarily to Figure 17 (A), the numerical PSF is constructed by summing the bilinear phase profiles shifted by Gz over the slice dimension; the central part of these phase profiles is shown on the left side of each computed in-plane PSF to illustrate how the bilinear saddle point is shifted during readout to produce a corresponding shift in the main lobe of the PSF.

[0139] Referring primarily to Figure 18 (C), under-sampling the number of projections leads to the same striation artifacts seen in traditional Fourier radial sampling, but the truncated readout leaves a fully resolved image in the center of the FOV instead of reducing the spatial resolution.

[0140] Referring primarily to Figure 20 , a phantom is used to surround the material of the tissue. As the field of view is reduced, the signal present in the image is also reduced, thereby cropping the image. Finally, the field of view is smaller than the phantom. This also enables the use of shorter echo intervals, in this case, shortened from 5.4 ms to 3.0 ms, thereby reducing the total duration of the 24 echo readout sequence from 130 ms to 72 ms.

[0141] Referring primarily to Figure 21 (B), the y-axis is Fourier encoded, while the x- and z-axes are spatiotemporally encoded. The distribution along x is collected during readout, while the z-distribution is collected indirectly.

[0142] Referring primarily to Figure 23(A), to compute the point spread function (PSF) for each sub-slice in the volume, compute the high-resolution xSPEN spatially encoded bilinear phase map and sum over each sub-slice width, resulting in the PSF for each slice.

[0143] Configurations

[0144] In some aspects, this reconstruction can be used for any xSPEN sequence, whether Cartesian or radial. It can be used with xSPEN sequences that have the same refocusing pulses or with different refocusing pulses that produce residual phases explained by equations (2) and (3). This reconstruction can also accommodate RARE-style image acquisition. The reconstruction can run in the following configurations: with only the ideal signal from equation (1), with the ideal signal and residual linear phase, with the ideal signal and quadratic phase, and finally with the ideal signal and both residual quadratic and linear phases. In the equations shown above, the spatio-temporal axis is y and the Fourier axis is x, but they can be switched. Alternatively, as described above with respect to Figure 11 equations (1) can be computed so that the x and y axes are the spatio-temporal axes, rather than the Fourier axes. Or, all three axes can be used with equation (1) to implement a three-dimensional radial xSPEN sequence reconstruction.

[0145] In some aspects, many implementations of this reconstruction are possible, which balance accuracy with computational speed and memory requirements. For 2D reconstruction, a more accurate reconstruction without the assumption of linear permanent gradients is obtained by directly constructing a high-resolution three-dimensional phase matrix for each time point in the readout, then summing the matrix over the slice dimension to obtain a numerical point spread function for each time point in the readout. This point spread function can then be converted to a matrix and the image solved by regularized matrix pseudo-inverse, conjugate gradient, or using any other iterative optimization algorithm.

[0146] In some aspects, the same method also applies to radial reconstruction scenarios, in which case the PSF rotates between repetitions. Any Fourier phase encoding applied in any dimension can be implemented as a separate explicit matrix multiplication, or can be implemented via non-uniform fast Fourier transform (NUFFT) before or after applying the point spread function matrix. NUFFT can also be used to apply permanent gradient-induced phase shifts at each time point in the readout, then apply the sum operator over the slices, rather than summing the bilinear phase over the slices. This will greatly reduce the memory size of the PSF matrix.

[0147] In some aspects, 3D reconstruction is implemented using a similar approach in which the bilinear phase is computed directly at high spatial resolution and then summed to obtain the PSF for each time point in the readout and each sub-slice of the 3D volume. Then, the NUFFT can be applied as in the 2D case for the Fourier phase encoding. Furthermore, as in the 2D case, the memory requirements can be greatly reduced by applying the readout phase shift using the NUFFT and then by the sub-slice summation operator.

[0148] Process

[0149] In another aspect, provided herein is a process of applying a plurality of xSPEN readout gradients on a unilateral magnetic resonance imaging system disclosed herein.

[0150] In some embodiments, applying the plurality of xSPEN readout gradients generates an encoding matrix for generating an image, wherein the encoding matrix accounts for a residual linear phase and a residual quadratic phase. In some embodiments, applying the plurality of xSPEN readout gradients generates an encoding matrix for generating an image. In some embodiments, the encoding matrix accounts for a residual linear phase and a residual quadratic phase. In some embodiments, the encoding matrix accounts for a residual linear phase. In some embodiments, the encoding matrix accounts for a residual quadratic phase.

[0151] In some embodiments, the plurality of xSPEN readout gradients disclosed herein comprises a first xSPEN readout gradient, a second xSPEN readout gradient, a third xSPEN readout gradient, and a fourth xSPEN readout gradient. In some embodiments, the plurality of xSPEN readout gradients disclosed herein comprises a first xSPEN readout gradient, a second xSPEN readout gradient, and a third xSPEN readout gradient. In some embodiments, the plurality of xSPEN readout gradients disclosed herein comprises a first xSPEN readout gradient and a second xSPEN readout gradient. In some embodiments, the plurality of xSPEN readout gradients disclosed herein comprises a first xSPEN readout gradient.

[0152] In some embodiments, the process disclosed herein further comprises applying, during the first refocusing pulse, the first xSPEN readout gradient along a second axis orthogonal to the first axis with the unilateral gradient coil set.

[0153] In some embodiments, the process disclosed herein further comprises applying, during the second refocusing pulse, the second xSPEN readout gradient along the second axis with the unilateral gradient coil set.

[0154] In some embodiments, the process disclosed herein further comprises applying, during the first refocusing pulse, the third xSPEN readout gradient along a third axis with the unilateral gradient coil set. In some embodiments, the process disclosed herein further comprises applying, during the first refocusing pulse, the third xSPEN readout gradient along a third axis with the unilateral gradient coil set.

[0155] In some embodiments, the processes disclosed herein further comprise applying a fourth xSPEN readout gradient along the third axis with the unilateral gradient coil set during the second refocusing pulse.

[0156] Various modifications to the implementations described in this disclosure can be readily apparent to those skilled in the art, and the generic principles defined herein can be applied to other implementations without departing from the spirit or scope of this disclosure. Thus, the claims are not intended to be limited to the implementations shown herein, but are to be accorded the widest scope consistent with the principles and novel features disclosed herein.

Claims

1. A unilateral magnetic resonance imaging system comprising: A housing, comprising: front surface; a permanent magnet for providing a static magnetic field extending from the permanent magnet into the region of interest about a first axis, wherein the first axis is perpendicular to the permanent magnet; radio frequency transmitting coil; and a single-sided gradient coil assembly, wherein the radio frequency transmit coil and the single-sided gradient coil assembly are positioned adjacent the front surface; RF receiving coil; a power supply, wherein the power supply is configured to cause current to flow through at least one of the radio frequency transmit coil or the single-sided gradient coil assembly to generate an electromagnetic field in the region of interest, wherein at least a portion of the region of interest is located outside the front surface; and A control circuit, the control circuit being configured to: transmitting an excitation pulse using the radio frequency transmitting coil; transmitting a first refocusing pulse using the radio frequency transmitting coil; applying a first xSPEN readout gradient along a second axis orthogonal to the first axis using the single-sided gradient coil set during the first refocusing pulse; performing phase encoding along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-sided gradient coil assembly; transmitting a second refocusing pulse using the radio frequency transmit coil; applying a second xSPEN readout gradient along the second axis using the single-sided gradient coil set during the second refocusing pulse; and Data is received using the radio frequency receiving coil.

2. The unilateral magnetic resonance imaging system according to claim 1 , wherein the control circuit is further configured to: Any number of refocusing pulses are transmitted using the radio frequency transmit coils before, during, or after the refocusing pulses occurring simultaneously with the xSPEN readout gradients.

3. The unilateral magnetic resonance imaging system according to claim 1 , wherein the control circuit is further configured to: An encoding matrix for generating an image is generated, wherein the encoding matrix takes into account the residual linear phase and the residual quadratic phase.

4. The unilateral magnetic resonance imaging system according to claim 1 , wherein the control circuit is further configured to: transmitting an excitation pulse using the radio frequency transmitting coil; transmitting a first refocusing pulse using the radio frequency transmitting coil; applying a first xSPEN readout gradient along a second axis orthogonal to the first axis using the single-sided gradient coil set during the first refocusing pulse; applying a second xSPEN readout gradient along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-sided gradient coil set during the first refocusing pulse; transmitting a second refocusing pulse using the radio frequency transmit coil; applying a third xSPEN readout gradient along the second axis using the single-sided gradient coil set during the second refocusing pulse; applying a fourth xSPEN readout gradient along the third axis using the single-sided gradient coil set during the second refocusing pulse; and Data is received using the radio frequency receiving coil.

5. The unilateral magnetic resonance imaging system according to claim 4, wherein the control circuit is further configured to: Any number of refocusing pulses are transmitted using the radio frequency transmit coils before, during, or after the refocusing pulses occurring simultaneously with the xSPEN readout gradients.

6. The unilateral magnetic resonance imaging system according to claim 4, wherein the control circuit is further configured to: During the first refocusing pulse, transmitting any number of additional xSPEN readout gradients along any number of additional imaging axes using any number of additional gradient coil sets; During the second refocusing pulse, transmitting any number of additional xSPEN readout gradients along any number of additional imaging axes using any number of additional gradient coil sets; and A fourth refocusing pulse is transmitted using the radio frequency transmit coil.

7. A single-sided magnetic resonance imaging system according to claim 1 or 4, wherein the amplitude of the xSPEN readout axis is temporally non-uniform during the refocusing pulse.

8. A single-sided magnetic resonance imaging system according to claim 1 or 4, wherein the gradient shape of the xSPEN readout axis is not constant in time during the refocusing pulse.

9. The unilateral magnetic resonance imaging system of claim 4, wherein the control circuit is further configured to generate an encoding matrix for generating an image, wherein the encoding matrix takes into account a residual linear phase and a residual quadratic phase.

10. The unilateral magnetic resonance imaging system of claim 4, wherein the control circuit is configured to perform spatiotemporal encoding along an arbitrary axis during acquisition.

11. The unilateral magnetic resonance imaging system of claim 10, wherein the arbitrary axis remains fixed during a single image acquisition.

12. The unilateral magnetic resonance imaging system of claim 10, wherein the arbitrary axis is rotated during a single image acquisition.

13. A method for generating a two-dimensional image by collecting a series of projections with different angles and then reconstructing the two-dimensional image.

14. A method for generating three-dimensional images by adding an additional gradient pulse array to a pulse sequence.

15. The method for generating a three-dimensional image according to claim 14, wherein the additional gradient pulse array is parallel to the xSPEN readout axis and provides spatial information along the permanent gradient axis to be resolved.

16. A method for generating three-dimensional images using a radial encoding scheme, wherein the effective axis of the gradient array rotates with the axis of the xSPEN readout gradient.

17. A process for applying a plurality of xSPEN readout gradients on the unilateral magnetic resonance imaging system of claim 1.

18. The process of claim 17, wherein applying the plurality of xSPEN readout gradients generates an encoding matrix for generating an image, wherein the encoding matrix accounts for residual linear phase and residual quadratic phase.

19. The process of claim 17, further comprising applying a first xSPEN readout gradient along a second axis orthogonal to the first axis using a single-sided gradient coil set during the first refocusing pulse.

20. The process of claim 19, further comprising applying a second xSPEN readout gradient along the second axis using the single-sided gradient coil set during a second refocusing pulse.

21. The process of claim 20, further comprising applying a third xSPEN readout gradient along the third axis using the single-sided gradient coil set during the first refocusing pulse.

22. The process of claim 21 further comprising applying a fourth xSPEN readout gradient along a third axis using the single-sided gradient coil set during the second refocusing pulse.

Images