Cross-term spatiotemporal encoding (xSPEN) technique for single-plane MRI

Abstract

The present disclosure provides a single-plane magnetic resonance imaging system including multiple components. The present disclosure also provides a method for generating multidimensional images by using the imaging system. The present disclosure also provides a method for applying multiple xSPEN readout gradients to the single-plane magnetic resonance imaging system disclosed herein. Several techniques exist for acquiring dynamic magnetic resonance (MR) images. Images can be acquired and generated in a small time window by rapidly sampling the entire k-space using a fast pulse sequence. For example, a spiral orbital gradient echo sequence can be used to sample the entire k-space in tens of milliseconds. Another approach uses small flip angle excitation to rapidly acquire lines in k-space. These approaches can also be combined with undersampling to further speed up image acquisition.

Description

[Technical Field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims the benefit of U.S. Provisional Patent Application No. 63 / 479,162, filed January 9, 2023, which is incorporated herein by reference in its entirety.

[0002] The present disclosure relates to medical systems, devices, and methods, and in particular to medical systems, devices, and methods for biomedical imaging, including single-sided or open magnetic resonance (MRI) scanners. [Background technology]

[0003] Single-sided or open-type magnetic resonance imaging (MRI) scanners generally have a permanent, or inherent, gradient magnetic field along the longitudinal axis extending from the single-sided MRI machine into the field of view. The permanent gradient magnetic field can be generated on the face of the permanent magnet by a rare-earth magnet and two sets of gradient coils. This orientation allows imaging within the field of view above the face of the magnet. By designing a system with this form factor, imaging can be performed without the need to enclose the area being imaged. This allows imaging without placing the patient inside the bore, enabling the scanner to be used with other medical devices, such as a biopsy robot. It also increases the comfort of claustrophobic patients by allowing them to be imaged outside the imaging bore of a conventional closed-type MRI scanner. Single-sided MRIs can also be portable and can image anything placed within the field of view.

[0004] The use of surface gradient coils in conjunction with single-plane scanners, as is generally required for single-plane scanning, can introduce blurring and substantially limit the image quality obtainable with single-plane MRI scanners by introducing field-of-view changes along the Z axis, echo drift, and / or ultimately k-space truncation. Thus, a need exists for improved sample acquisition for single-plane or open MRI scanners. Summary of the Invention

[0005] Several techniques exist for acquiring dynamic magnetic resonance (MR) images. Rapid pulse sequences are used to rapidly sample the entire k-space, generating images in a small time window. For example, spiral orbital gradient echo sequences can be used to sample the entire k-space in tens of milliseconds. Another approach uses small flip-angle excitation to rapidly acquire lines in k-space. These approaches can also be combined with undersampling to further speed up image acquisition. However, each of these techniques assumes a strong, uniform magnetic field and tends to fail as the uniformity of the main magnetic field decreases.

[0006] Several methods exist for rapidly acquiring images in inhomogeneous fields, the most relevant being xSPEN. Different types of xSPEN pulse sequences are described in Zhang, Zhiyong et al., "Single-Scan MRI with Exceptional Resilience to Field Heterogeneities," Magnetic Resonance in Medicine, vol. 77, 2017, pp. 623-634 (hereinafter "Zhang, Zhiyong et al."), which is incorporated herein by reference in its entirety. In some instances, pulse sequences similar to some of the xSPEN sequences disclosed in the Supplementary Information of Zhang, Zhiyong et al., section 5.4, may be used. The Supplementary Information of Zhang, Zhiyong et al. is also incorporated herein by reference in its entirety. These sequences allow for the definition of the imaging axis without interference from permanent gradients.

[0007] In low SNR environments, it may be necessary to perform signal averaging or k-space sampling at short echo times to improve the SNR. Adding these adjustments to existing pulse sequences can slow down image acquisition.

[0008] Existing techniques are also not designed for inhomogeneous magnetic fields. Acquiring all of k-space with a single spiral is not feasible in systems that use permanent gradients. Permanent gradients can limit the trajectories that can be achieved with the scanner.

[0009] To address these issues, xSPEN can be used to perform high-speed imaging with permanent gradients, but existing versions are not optimal for single-sided low-field systems. At low fields with strong permanent gradients, the SNR is much lower, limiting the applicability of xSPEN.

[0010] In one aspect, a single-sided magnetic resonance imaging system disclosed herein includes a housing having a front surface and a permanent magnet for providing a static magnetic field, the static magnetic field extending from the permanent magnet to a region of interest about a first axis perpendicular to the permanent magnet. The housing further includes a radio frequency transmit coil and a single-sided gradient coil set, the radio frequency transmit coil and the single-sided gradient coil set being located near the front surface. The single-sided magnetic resonance imaging system further includes a radio frequency receive coil and a power source, the power source configured to pass current through at least one of the radio frequency transmit coil or the single-sided gradient coil set to generate an electromagnetic field in the region of interest, at least a portion of which is outside the front surface. The single-plane magnetic resonance imaging system further includes control circuitry configured to transmit an excitation pulse using the radio frequency transmit coil, transmit a first refocusing pulse using the radio frequency transmit coil, apply a first xSPEN read gradient during the first refocusing pulse along a second axis orthogonal to the first axis using the single-plane gradient coil set, perform phase encoding along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-plane gradient coil set, transmit a second refocusing pulse using the radio frequency transmit coil, apply a second xSPEN read gradient during the second refocusing pulse along the second axis using the single-plane gradient coil set, and receive data using the radio frequency receive coil. In some embodiments, the control circuitry is further configured to transmit a number of refocusing pulses using the radio frequency transmit coil before, during, or after multiple refocusing pulses occurring simultaneously with the xSPEN read gradients. In some embodiments, the control circuitry is further configured to generate an encoding matrix used to generate the image, the encoding matrix taking into account the residual primary phase and the residual secondary phase.In some embodiments, the control circuitry is further configured to transmit an excitation pulse using the radio frequency transmit coil, transmit a first refocusing pulse using the radio frequency transmit coil, apply a first xSPEN read gradient during the first refocusing pulse along a second axis orthogonal to the first axis using the single-plane gradient coil set, apply a second xSPEN read gradient during the first refocusing pulse along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-plane gradient coil set, transmit the second refocusing pulse using the radio frequency transmit coil, apply a third xSPEN read gradient during the second refocusing pulse along the second axis using the single-plane gradient coil set, apply a fourth xSPEN read gradient during the second refocusing pulse along the third axis using the single-plane gradient coil set, and receive data using the radio frequency receive coil. In some embodiments, the control circuitry is further configured to transmit any number of refocusing pulses using a radio frequency transmit coil before, during, or after multiple refocusing pulses occurring simultaneously with the xSPEN read gradient. In some embodiments, the control circuitry is further configured to transmit any number of additional xSPEN read gradients during a first refocusing pulse along any number of additional imaging axes using any number of additional gradient coil sets, transmit any number of additional xSPEN read gradients during a second refocusing pulse along any number of additional imaging axes using any number of additional gradient coil sets, and transmit a fourth refocusing pulse using a radio frequency transmit coil. In some embodiments, the amplitude in the xSPEN read axis is non-uniform throughout the refocusing pulse. In some embodiments, the gradient shape in the xSPEN read axis is time-invariant during the refocusing pulse. In some embodiments, the control circuitry is further configured to generate an encoding matrix used to generate the image, the encoding matrix accounting for residual primary and secondary phases. In some embodiments, the control circuitry is configured to spatiotemporally encode along any axis during acquisition.In some embodiments, any axis is fixed during a single image acquisition. In some embodiments, any axis is rotated during a single image acquisition.

[0011] In another aspect, a method is disclosed herein for generating a two-dimensional image by collecting a series of projections at varying angles and then reconstructing the two-dimensional image.

[0012] In another aspect, disclosed herein is a method for generating a three-dimensional image by adding an additional gradient pulse array to the pulse sequence, which in some embodiments is parallel to the xSPEN read axis and provides spatial information along the permanent gradient axis to be resolved.

[0013] In another aspect, a method is disclosed herein for generating three-dimensional images using a radial encoding scheme, where the effective axis of the gradient array is rotated with the axis of the xSPEN readout gradient.

[0014] In another aspect, disclosed herein are methods of applying multiple xSPEN read gradients to a single-plane magnetic resonance imaging system as disclosed herein. In some embodiments, applying the multiple xSPEN read gradients generates an encoding matrix used to generate an image, the encoding matrix accounting for residual primary and secondary phases. In some embodiments, the method further includes applying a first xSPEN read gradient during a first refocusing pulse along a second axis orthogonal to the first axis using a single-plane gradient coil set. In some embodiments, the method further includes applying a second xSPEN read gradient during a second refocusing pulse along the second axis using a single-plane gradient coil set. In some embodiments, the method further includes applying a third xSPEN read gradient during the first refocusing pulse along a third axis using a single-plane gradient coil set. In some embodiments, the method further includes applying a fourth xSPEN read gradient during the second refocusing pulse along the third axis using a single-plane gradient coil set.

[0015] Incorporation by Reference All publications, patents, and patent applications mentioned in this specification are incorporated herein 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 publications and patents or patent applications incorporated by reference conflict with the disclosure contained herein, it is intended that the present specification supersede and / or take precedence over any such conflicting document. [Brief explanation of the drawings]

[0016] The novel features of the various aspects are set forth with particularity in the appended claims, but the described aspects, both as to organization and method of operation, can best be understood by reference to the following description when considered in conjunction with the accompanying drawings.

[0017] [Figure 1] FIG. 1 is a perspective view of a magnetic resonance imaging (MRI) scanner according to various aspects of the present disclosure.

[0018] [Figure 2] 2 is an exploded perspective view of the MRI scanner of FIG. 1 exposing the permanent magnet assembly and gradient coil set within the housing, according to various aspects of the present disclosure. FIG.

[0019] [Figure 3] FIG. 2 is an elevational view of the MRI scanner of FIG. 1 in accordance with various aspects of the present disclosure.

[0020] [Figure 4] FIG. 2 is an elevational view of the MRI scanner of FIG. 1 in accordance with various aspects of the present disclosure.

[0021] [Figure 5] FIG. 2 is a perspective view of a permanent magnet assembly of the MRI scanner of FIG. 1 in accordance with various aspects of the present disclosure.

[0022] [Figure 6] 2 is an elevational view of a gradient coil set and permanent magnet assembly of the MRI system shown in FIG. 1 in accordance with various aspects of the present disclosure.

[0023] [Figure 7] FIG. 1 is a control block diagram for a single-plane MRI system according to various aspects of the present disclosure.

[0024] [Figure 8] 1 is a schematic diagram of a magnetic field gradient along the Z axis according to various aspects of the present disclosure.

[0025] [Figure 9] 1 is a representative graph of a swept frequency pulse according to various aspects of the present disclosure.

[0026] [Figure 10]FIG. 1 is a diagram of a pulse sequence for cross-term spatiotemporal encoding for high-speed imaging using single-plane MRI.

[0027] [Figure 11] FIG. 1 is a diagram of a pulse sequence for cross-term spatiotemporal encoding for radial imaging using single-plane MRI.

[0028] [Figure 12] Figure 12(A) shows a single-plane prostate MRI scanner. The subject is positioned in the lithotomy position, with a hole in the magnet for surgical and biopsy access to the prostate. Figure 12(B) shows the configuration for xSPEN imaging. Slice selection is along the z-dimension of the magnet, and the in-plane (x / y) dimensions are both phase-encoded in the conventional manner.

[0029] [Figure 13] FIG. 1 illustrates a phase-encoded xSPEN imaging pulse sequence.

[0030] [Figure 14] Figure 1 shows the xSPEN spatiotemporal encoding function, moving along the y-direction with a 1.5 ms long readout window in the central slice of an in vivo scan. The FOV in this example is 18x18 cm, and the x-dimension was phase encoded.

[0031] [Figure 15] Figures 15(A) and 15(B) show reconstructions of the central slice of a 10 cm diameter ACR phantom. Figure 15(A) was obtained using the Fourier reconstruction described above. Figure 15(B) was obtained using the model-based reconstruction described herein, which fully accounts for gradient nonlinearity.

[0032] [Figure 16] Figure 16(A) shows a single-plane MRI scanner, and Figure 16(B) shows a radial xSPEN pulse sequence.

[0033] [Figure 17] Figure 17(A) shows the numerically calculated radial xSPEN point spread functions near the beginning, middle, and end of the readout, and Figure 17(B) shows the PSFs midway through the readout at radial angles of 0, 45, and 90 degrees.

[0034] [Figure 18] Figure 18(A) shows simulated (sinc) radial xSPEN PSFs during readout at three different angles. Figure 18(B) shows simulated 45-degree sinc radial xSPEN PSFs at three different points in the readout. With sinc, the FOV is swept from top left to bottom right. Figure 18(C) shows phantom image reconstructions for full projection, quarter projection, and half readout duration.

[0035] [Figure 19] Figure 19(A) shows an xSPEN sinogram of a resolution phantom. The data has the shape of an x-ray sinogram and no k-space dimension exists. Figure 19(B) shows images reconstructed from the sinogram data using three methods: an inverse Radon transform that ignores gradient nonlinearities in the initial Fourier transform step; an analytical sinc function reconstruction that does not consider the Gz nonlinearity; and a numerical PSF approach as shown in Figures 17(A) and 17(B).

[0036] [Figure 20] FIG. 20 shows images of the ACR extremity phantom acquired with varying field of view / read duration.

[0037] [Figure 21]Figure 21(A) shows a single-plane MRI scanner, and Figure 21(B) shows a 3D xSPEN pulse sequence based on a CPMG-RARE scan using WURST excitation and refocusing pulses, with the first spin echo positioned after the second refocusing pulse by using the CHORUS technique.

[0038] [Figure 22] Figure 22(A) shows a cross section of a bilinear xSPEN spatial encoding phase function through the slab (z) and xSPEN encoded in-plane dimension (x) for different x phase encodings where the saddle / sensitive point moves through the slab to different z positions. Figure 22(B) shows a cross section of the phase function at different times during a readout (without x phase encoding) where the saddle / sensitive point shifts to different in-plane (x) positions.

[0039] [Figure 23] Figure 23(A) shows the high-resolution xSPEN spatial encoding bilinear phase map (top; cross-section through the slice) and the PSF for each slice (bottom). Figure 23(B) shows that the forward model relating the reconstructed image to the data is achieved by applying the point spread functions for each slice and each read time to the input image stack, and then computing a type III NUFFT to apply the phase encoding.

[0040] [Figure 24] Twelve slices from two consecutive slabs of the ACR phantom reconstructed from 3D xSPEN data. The images clearly show feature development across slices, reflecting the sequence's ability to resolve intra-slab subslice detail. DETAILED DESCRIPTION OF THE INVENTION

[0041] The accompanying drawings are not intended to be drawn to scale. Corresponding reference characters indicate corresponding parts throughout the several views. For clarity, not every component may be labeled in every figure. The examples described herein illustrate, in one form, particular embodiments of the present disclosure, and such illustrations are not to be construed as limiting the scope of the present disclosure in any manner.

[0042] The following international patent applications are each incorporated by reference herein in their entirety: International Patent Application No. PCT / US2020 / 018352, entitled "SYSTEMS AND METHODS FOR ULTRALOW FIELD RELAXATION DISPERSION," filed on February 14, 2020, and published as International Publication No. WO2020 / 168233 International Patent Application No. PCT / US2020 / 019530, entitled "SYSTEMS AND METHODS FOR PERFORMING MAGNETIC RESONANCE IMAGING," filed February 24, 2020, and published as International Publication No. WO2020 / 172673 International Patent Application No. PCT / US2020 / 019524, entitled "PSEUDO-BIRDCAGE COIL WITH VARIABLE TUNING AND APPLICATIONS THEREOF," filed on February 24, 2020, published as International Publication No. WO2020 / 172672 International Patent Application No. PCT / US2020 / 024776, entitled "SINGLE-SIDED FAST MRI GRADIENT FIELD COILS AND APPLICATIONS THEREOF," filed March 25, 2020, published as International Publication No. WO2020 / 198395 International Patent Application No. PCT / US2020 / 024778, entitled "SYSTEMS AND METHODS FOR VOLUMETRIC ACQUISITION IN A SINGLE-SIDED MRI SYSTEM," filed on March 25, 2020, published as International Publication No. WO2020 / 198396 International Patent Application No. PCT / US2020 / 039667, entitled "SYSTEMS AND METHODS FOR IMAGE RECONSTRUCTIONS IN MAGNETIC RESONANCE IMAGING," filed June 25, 2020, published as International Publication No. WO2020 / 264194 · International Patent Application No. PCT / US2021 / 014628, entitled "MRI-GUIDED ROBOTIC SYSTEMS AND METHODS FOR BIOPSY," filed on January 22, 2021 · International Patent Application No. PCT / US2021 / 018834, filed February 19, 2021, entitled "RADIO FREQUENCY RECEPTION COIL NETWORKS FOR SINGLE-SIDED MAGNETIC RESONANCE IMAGING" · International Patent Application No. PCT / US2021 / 021464, entitled "PHASE ENCODING WITH FREQUENCY SWEEP PULSES FOR MAGNETIC RESONANCE IMAGING IN INHOMOGENEOUS MAGNETIC FIELDS," filed on March 9, 2021 · International Patent Application No. PCT / US2021 / 021461, entitled "PULSE SEQUENCES AND FREQUENCY SWEEP PULSES FOR SINGLE-SIDED MAGNETIC RESONANCE IMAGING," filed March 9, 2021 · International Patent Application No. PCT / US2022 / 071924, entitled "INTERVENTIONAL LOCALIZATION GUIDE AND METHOD FOR MRI GUIDED PELVIC INTERVENTIONS," filed April 26, 2022 · International Patent Application No. PCT / US2022 / 082551, entitled "RELAXATION-BASED MAGNETIC RESONANCE THERMOMETRY WITH A LOW-FIELD SINGLE-SIDED MRI SCANNER," filed on December 29, 2022

[0043] U.S. Patent Application Publication No. 2018 / 0356480, published December 13, 2018, and entitled "UNILATERAL MAGNETIC RESONANCE IMAGING SYSTEM WITH APERTURE FOR INTERVENTIONS AND METHODOLOGIES FOR OPERATING SAME," is also incorporated herein by reference in its entirety.

[0044] It should be noted that the illustrative examples are not limited in application or use to the details of construction and arrangement of parts shown in the accompanying drawings and specification. The illustrative examples may be implemented in or incorporated into other embodiments, variations, and modifications, and may be practiced or carried out in various ways. Furthermore, unless otherwise noted, the terms and phrases used herein have been chosen for the convenience of the reader and for the purpose of describing the illustrative examples, not for the purpose of limitation thereof. It will also be understood that one or more of the following embodiments, embodiment expressions, and / or examples may be combined with any one or more of the other following embodiments, embodiment expressions, and / or examples.

[0045] According to various aspects, an MRI system is provided that can include a unique imaging region that can be offset from the face of the magnet. Such offset and single-plane MRI systems are less restrictive than traditional MRI scanners. Additionally, this form factor can have built-in or inherent magnetic field gradients that generate a range of magnetic field values ​​across the region of interest. In other words, the inherent magnetic field can be non-uniform and always active. The non-uniformity of the magnetic field strength in the region of interest for a single-plane MRI system can be greater than 200 parts per million (ppm). For example, the non-uniformity of the magnetic field strength in the region of interest for a single-plane 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, or greater than 10,000 ppm. In one example, the non-uniformity in the region of interest can be 81,000 ppm.

[0046] The intrinsic magnetic field gradient may be generated by a permanent magnet within the MRI scanner. The magnetic field strength in the region of interest for a single-plane MRI system may be, for example, less than 1 Tesla (T). For example, the magnetic field strength in the region of interest for a single-plane MRI system may be less than 0.5 T. In other examples, the magnetic field strength may be greater than 1 T, for example, 1.5 T. The system may operate at lower magnetic field strengths compared to typical MRI systems, allowing for relaxation of design constraints on the gradient and / or radio frequency coils and / or allowing additional mechanisms, such as robotics, to be used with the MRI scanner. Exemplary MRI-guided robotic systems are further described, for example, in International Patent Application No. PCT / US2021 / 014628, filed January 22, 2021, and entitled "MRI-GUIDED ROBOTIC SYSTEMS AND METHODS FOR BIOPSY."

[0047] 1-6 illustrate an MRI scanner 100 and its components. As shown in FIGS. 1 and 2, the MRI scanner 100 includes a housing 120 having a concave face or front surface 125. In other embodiments, the face of the housing 120 may be flat and planar. The front surface 125 may face a subject being imaged by the MRI scanner. As shown in FIGS. 1 and 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. In other examples, the housing 120 may not include the electromagnet 160. Furthermore, in certain examples, the RF receive coil 170 and the RF transmit coil 140 may be incorporated into an integrated Tx / Rx coil array. In various examples, the MRI scanner 100 is a single-sided scanner, with various components, such as the permanent magnet assembly 130, the RF transmit coil (TX) 140, the gradient coil set 150, the electromagnet 160, and the RF receive coil (RX) 170, located on the same side of the field of view.

[0048] Referring primarily to FIGS. 3-5, the permanent magnet assembly 130 includes an array of magnets. The array of magnets forming the permanent magnet assembly 130 is configured to cover the front surface 125, or patient-facing surface, of the MRI scanner 100 (see FIG. 3) and is shown as a horizontal bar in FIG. 4. The permanent magnet assembly 130 includes multiple cylindrical permanent magnets in a parallel configuration. Referring primarily to FIG. 5, the permanent magnet assembly 130 includes parallel plates 132 held together by a bracket 134. The system may be mounted to the housing 120 of the MRI scanner 100 at a bracket 136. Multiple bores 138 may be present in the parallel plates 132. The permanent magnet assembly 130 may include any suitable magnetic material, including, but not limited to, a rare earth-based magnetic material, such as, for example, a neodymium-based magnetic material.

[0049] The permanent magnet assembly 130 defines an access opening or bore 135 that may provide access to the patient through the housing 120 from the opposite side of the housing 120. In other aspects of the present disclosure, the array of permanent magnets forming the permanent magnet assembly within the housing 120 may be boreless and may define an uninterrupted or continuous arrangement of permanent magnets without a bore defined therethrough. In yet other examples, the array of permanent magnets in the housing 120 may form more than one bore / access opening through the permanent magnets.

[0050] According to various aspects of the present disclosure, the permanent magnet assembly 130 provides a magnetic field B to the region of interest 190 along the Z-axis shown in FIG. 1 . The 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 B away from the face of the permanent magnet assembly 130. The Z-axis may define the direction of the main magnetic field B. The main magnetic field B may decrease at a specific gradient along the Z-axis, i.e., in the direction indicated by the arrow in FIG. 1 , the further away from the face of the permanent magnet assembly 130.

[0051] In one embodiment, the magnetic field non-uniformity in the region of interest 190 of the permanent magnet assembly 130 can be approximately 81,000 ppm. In another embodiment, the magnetic field strength non-uniformity in the region of interest 190 of the permanent magnet assembly 130 can be between 200 ppm and 200,000 ppm, in certain instances greater than 1,000 ppm, and in various instances greater than 10,000 ppm.

[0052] In one embodiment, the magnetic field strength of the permanent magnet assembly 130 may be less than 1 T. In another embodiment, the magnetic field strength of the permanent magnet assembly 130 may be less than 0.5 T. In other examples, the magnetic field strength of the permanent magnet assembly 130 may be greater than 1 T, such as 1.5 T. Referring primarily to FIG. 1 , the Y-axis extends up and down 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 perpendicular to one another, with the positive direction of each axis indicated by the corresponding arrow in FIG. 1 .

[0053] The RF transmit coil 140 is configured to transmit an RF waveform and associated electromagnetic field. 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, referred to as B1, that is orthogonal to the direction of the permanent magnetic field (e.g., in the orthogonal plane).

[0054] Referring primarily to FIG. 3 , gradient coil set 150 includes two sets of gradient coils 152, 154. The set of gradient coils 152, 154 is located on the face or front surface 125 of permanent magnet assembly 130, intermediate permanent magnet assembly 130 and region of interest 190. Each set of gradient coils 152, 154 includes a coil portion on either side of bore 135. Referring to the axes in FIG. 1 , gradient coil set 154 can be, for example, the gradient coil set corresponding to the X-axis, and gradient coil set 152 can be, for example, the gradient coil set corresponding to the Y-axis. Gradient coils 152, 154 enable encoding along the X-axis and Y-axis, as described further herein.

[0055] Referring now to FIG. 7 , a control block diagram for a single-plane MRI system 300 is shown. The single-plane MRI scanner 100 and / or its components ( FIGS. 1-6 ) can be incorporated into the MRI system 300 in various embodiments of the present disclosure. For example, the imaging system 300 includes a permanent magnet assembly 308, which in various examples can be similar to the permanent magnet assembly 130 (see FIGS. 2-5 ). The imaging system 300 also includes an RF transmit coil 310, which can be similar to, for example, the RF transmit coil 140 (see FIG. 3 ). Furthermore, the imaging system 300 includes an RF receive coil 314, which can be similar to, for example, the RF receive coil 170 (see FIG. 3 ). In various embodiments, the RF transmit coil 310 and / or the RF receive coil can also be disposed within the housing of the MRI scanner; in certain examples, 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 gradient magnetic fields to facilitate imaging of objects within a field of view 312.

[0056] The single-sided MRI system 300 also includes a computer 302 in signal communication with a spectrometer 304 and configured to transmit and receive signals between the computer 302 and the spectrometer 304 .

[0057] The main magnetic field B0 generated by the permanent magnet 308 extends in a field of view 312, away from the permanent magnet 308 and away from the RF transmit coil 310. The field of view 312 includes the object being imaged by the MRI system 300.

[0058] During the imaging process, a main magnetic field B0 extends within the field of view 312. The direction of the effective magnetic field (B1) changes in response to RF pulses and the associated electromagnetic field from the RF transmit coil 310. For example, the RF transmit coil 310 is configured to selectively transmit RF signals or pulses to objects within the field of view, e.g., tissue. These RF pulses change the effective magnetic field experienced by spins within a sample (e.g., a patient's tissue). The RF pulses change the effective magnetic field experienced by magnetization in the rotating frame. When the RF pulse is on resonance, the effective magnetic field is only along the axis of the RF pulse. When off resonance, the axis of the effective magnetic field is between the applied RF pulse and the static magnetic field. The RF pulse may be, for example, a chirp or frequency sweep pulse, as described further herein.

[0059] Furthermore, when an object within the field of view 312 is excited by RF pulses from the RF transmit coil 310, the precession of the object produces induced or MR currents, which are detected by the RF receive coil 314. The RF receive coil 314 may transmit excitation data to an RF preamplifier 316. The RF preamplifier 316 may boost or amplify the excitation data signal and transmit the excitation data signal to the spectrometer 304. The spectrometer 304 may transmit the excitation data to the computer 302 for storage, analysis, and image construction. The computer 302 may, for example, combine multiple stored excitation data signals to generate an image.

[0060] From the spectrometer 304, the signal can be relayed via an RF power amplifier 306 to an RF transmit coil 310 and via a gradient power amplifier 318 to a gradient coil 320. The RF power amplifier 306 amplifies the signal and transmits it to the RF transmit coil 310. The gradient power amplifier 318 amplifies the gradient coil signal and transmits it to the gradient coil 320.

[0061] Systems and methods for effectively acquiring nuclear magnetic resonance spectra and images in inhomogeneous magnetic fields, such as using single-plane MRI scanners 100 and systems 300, are described herein.

[0062] Imaging with single-sided or open MRI presents many challenges. Typically, the two sets of gradient coils in a single-sided system (see Figure 6) are placed at the face of the permanent magnet assembly. As a result, the gradient amplitude decreases with distance from the face of the permanent magnet assembly. Therefore, for a given phase-encoding array, the field of view changes as one moves along the axis of the permanent magnetic 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.

[0063] FIG. 8 is a schematic diagram 500 of the magnetic field gradient along the Z axis for the MRI scanner 100. The permanent magnet 130 has an inherent gradient along the Z axis. The strength of the Z gradient decreases with distance from the permanent magnet 130. The schematic diagram shows that the Z gradient curves, decreasing the strength of the gradient with distance from the permanent magnet. The MRI scanner 100 images multiple slices to generate a slab. Each slice is excited for imaging with a different frequency. A lower frequency excites tissue in the slice farther from the permanent magnet, and a higher frequency excites tissue in the slice closer to the magnet. In the schematic diagram, the slab, or axial image, is shown starting from slice 0 to slice 1. n Each slice has a corresponding frequency f0 to f n f0 is up to f n It is important to note that single-sided MRI systems may be constructed such that there are regions where the frequency may rise along the Z axis, opposite to the gradients described.

[0064] FIG. 9 shows a representative graph 900 of a swept frequency pulse or chirp pulse (shown demodulated to baseband frequency) with a low-to-high sweep direction. A chirp excitation pulse with a low-to-high sweep direction is an example of a frequency-swept excitation pulse. A low-to-high chirp pulse starts at a low frequency and increases in frequency over the duration of the pulse. The pulse can begin at the lowest desired frequency and end when it reaches the maximum desired frequency. The pulse frequency in graph 900 can be a negative-to-positive frequency offset from the baseband frequency. In other words, the frequency sweeps from negative to positive in addition to the baseband frequency. For example, for a frequency sweep of + / - 100 KHz, the sweep is from 100 KHz below the baseband frequency to 100 KHz above the baseband frequency. It is important to note that the frequency direction can be reversed to sweep from positive to negative frequencies.

[0065] The frequency of the chirp pulse can vary from the minimum (lowest) desired frequency to the maximum (highest) desired frequency. The sweep rate of the pulse is the difference between the highest and lowest frequencies in the pulse divided by the time required to transition between the highest and lowest frequencies. In one embodiment, the frequency range covered by the sweep frequency pulse can range from -20 KHz to 20 KHz, i.e., 40 KHz, with the center frequency varying from slab to slab. For example, the slabs 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 slab centered at 2.62 MHz, the chirp pulse sweeps from 2.60 MHz to 2.64 MHz, i.e., 40 KHz. In other embodiments of the present disclosure, bandwidths as low as 10 KHz and as high as 200 KHz can be used for the frequency sweep pulse. Furthermore, the sweep range can be less than 40 KHz in various examples.

[0066] Cross-term Spatiotemporal Encoding (xSPEN) Technique for Fast Imaging Using Single-Plane MRI 10, the present application discloses a version of xSPEN that can improve SNR and enable high-speed three-dimensional image acquisition. In some embodiments, imaging with xSPEN can significantly reduce scan time because one of the axes in the image is sampled at a high rate. Adapting xSPEN to single-sided, low-field systems requires some modifications to the pulse sequence design.

[0067] First, fast sequences can include chirped excitation pulses to acquire thicker slabs. Acquiring thicker slabs can enhance SNR because a larger volume, and therefore a larger magnetization, contributes to the signal. Conventional xSPEN sequences use hard pulses to excite magnetization because they were designed for systems with much weaker gradients. Under these conditions, the hard pulses can excite slices thick enough to achieve sufficient SNR. In some embodiments, pulse sequences can be modified using timing to eliminate the secondary phase that chirped excitation imparts to magnetization. See Foroozandeh, Mohammadali et al., "Improved ultra-broadband chirp excitation," Journal of Magnetic Resonance, vol. 302, 2019, pp. 28-33, and Power, J.E. et al., "Increasing the quantitative bandwidth of NMR measurements," The Royal Society of Chemistry, Chem. Commun., vol. 52, 2016, pp. 2916-2919, each of which is incorporated by reference in its entirety.

[0068] In some embodiments, multiple echoes can be collected using the xSPEN sequence to further increase the SNR. As with other pulse sequences using chirped refocusing pulses, every other echo can be discarded because the quadratic phase imparted by the chirped pulse is preserved. This document demonstrates how collecting multiple echoes per acquisition allows for signal averaging without having to wait for the magnetization to return to the long axis.

[0069] In another embodiment, each echo can be used to sample a different line in k-space. By using gradients to move the magnetization to different points in k-space before collecting the echoes, it is possible to collect multiple k-space lines during excitation, speeding up image acquisition.

[0070] Some embodiments of the fast xSPEN sequences described herein may include interleaved slices to gather three-dimensional information. By designing such scans, changes occurring in all three dimensions, such as the movement of a biopsy needle in a patient or the formation of an ice ball during a cryoablation experiment, can be monitored, requiring the generation of volumetric images. In some embodiments, a volume is generated by acquiring a series of two-dimensional images in the x / y plane along the z-axis. Images can then be interpolated along the z-axis to form the complete volume.

[0071] Another embodiment of the fast xSPEN sequence described herein may include a third axis that is phase-encoded. This allows for the acquisition of a thick excitation profile, thereby taking advantage of the increased SNR while simultaneously encoding image information across that axis. This image is shown in FIG. 11. In this manner, volumetric images can be acquired using xSPEN sequences. In one version of this technique, two gradient pulse arrays are used to encode spatial information along two axes while a third is read during the acquisition period. One gradient pulse array is along an axis perpendicular to the read axis, and the other is along the same axis. The gradient pulse array perpendicular to the read axis is used to phase-encode the spatial information. The other array encodes information spatiotemporally, but not along the gradient axis, but instead along the axis of the other gradient that is on during the two chirp pulses used to impart a hyperbolic phase to the magnetization. Thus, with one phase-encoding array, another array encoding spatial information along one axis, and a readout along another axis, all three axes can be encoded, enabling the acquisition of volumetric images using xSPEN.

[0072] In another embodiment, another approach by which three-dimensional volumes can be acquired using xSPEN is to use a radial encoding scheme paired with a gradient array. Encoding along two axes can be achieved by collecting projections along various angles until enough projections along those angles have been collected to reconstruct a two-dimensional image. Encoding the third dimension is achieved by adding a gradient pulse array to the two-dimensional radial xSPEN procedure. The orientation of the gradient pulse array is rotated with the angle of the projections, so that a plane along the third axis and the angle of the projection axis is acquired. By collecting multiple angles in this manner, three-dimensional images can be reconstructed. While this method generates MRI images without using Fourier encoding, the additional application of Fourier phase encoding along the xSPEN-encoded rotation axis can enhance spatial resolution without the traditional xSPEN resolution / SNR tradeoff.

[0073] Parameter Range The possible range of field inhomogeneity is 10kHz to 300kHz over a depth of 160mm.

[0074] The possible range of excitation bandwidth is It is between 5 kHz and 200 kHz.

[0075] The possible range of refocusing bandwidth is It is between 5 kHz and 200 kHz.

[0076] The possible range of times for acquiring images is 1 second to 30 minutes.

[0077] In another aspect, the present application provides a Cartesian xSPEN sequence, in which one axis is spatiotemporally encoded while the other is Fourier encoded. Due to the manner in which spatial information is encoded using xSPEN, it is possible to encode a field of view along the spatiotemporal axis that is smaller than the size of the object being imaged. In conventional MRI scans, in which both axes are Fourier encoded, selecting a field of view smaller than the object results in aliasing. Because xSPEN uses a different encoding method, it is possible to encode a smaller field of view, allowing the user to zoom in on one axis.

[0078] In some embodiments, the present application discloses a method that allows for cropping the field of view using only the readout gradient and without other tools such as saturation bands by zooming in along both axes. Typically, only the spatiotemporal axis may have a field of view that is smaller than the object size. In applications where only a portion of the field of view contains useful anatomical information, it is ideal to be able to crop the field of view in both axes so that the image is zoomed in on important anatomical structures.

[0079] Referring to Figure 11, a radial xSPEN pulse sequence generates an image by rotating an axis and collecting a series of projections. These various projections are then combined into a single two-dimensional image using interactive conjugate gradient reconstruction or a filtered back-reconstruction approach. Unlike xSPEN sequences described in the literature, the radial xSPEN sequence applies gradients to both axes during the first two chirp refocusing pulses, leaving neither of these two axes Fourier-encoded. The projection axis can be changed by adjusting the relative strength of the gradient pulses applied in each axis. If gradients are applied to only one axis, the projection will be along that axis only. If some power is applied to both axes, the projection axis will be between the two axes. By imaging in this manner, it is possible to collect fully spatiotemporally encoded two-dimensional images. By eliminating Fourier encoding from the 2D image acquisition, it is possible to collect images with a field of view smaller than the object without aliasing. In some embodiments, this reduced field of view is obtained by truncating the read axis, resulting in reduced transverse relaxation, thus shortening echo times and increasing signal intensity.

[0080] In another embodiment, the geometric center of this reduced field of view can be moved by adjusting the scheduling of the read angles made between the two axes to accommodate non-uniform sampling. Both the angle scheduling and the read truncation can also be adjusted to generate an elliptical field of view.

[0081] In another embodiment, a small amount of Fourier encoding perpendicular to the direction of xSPEN band movement can be added to more precisely resolve the reduced field of view across the xSPEN projections. One example of this encoding scheme applies rotational Fourier phase encoding along each of the xSPEN rotational projections to enhance spatial resolution.

[0082] The third dimension of volumetric encoding can be achieved by various techniques. In the first, an array of 2D images sliced ​​along the third dimension encompassing the volume is collected. These slices are acquired by varying the excitation frequency and bandwidth or by adding slice-selective electromagnetic gradient pulses during excitation to shift the Larmor frequency of the desired volume subset to the desired range. In some embodiments, a second method can add a third xSPEN read axis using additional electromagnetic gradients. Thus, the image projection is tilted in the third dimension, allowing a third imaging axis to be encoded. In some embodiments, a third method can excite the entire desired volume and implement a phase encoding axis in the third dimension.

[0083] The instantaneous amplitude of the xSPEN read axis controls the speed at which the point spread function encoding moves through the image. By adjusting the amplitude, and therefore the speed, the SNR of a segment along the read projection can be advantageously increased. Furthermore, if the xSPEN read axis gradient shape is dynamically changed (e.g., by adding another controllable read axis), the point spread function size can be dynamically adjusted and pixel resolution can be changed across the image.

[0084] Parameter Range The possible range of field inhomogeneity is 10kHz to 300kHz over a depth of 160mm.

[0085] The possible range of excitation bandwidth is It is between 5 kHz and 200 kHz.

[0086] The possible range of refocusing bandwidth is It is between 5 kHz and 200 kHz.

[0087] The possible range of times for acquiring images is 1 second to 30 minutes.

[0088] Conjugate Gradient Least Square Reconstruction for Cross-Term Spatiotemporal Encoding (xSPEN) Using Single-Plane MRI In Cartesian xSPEN reconstruction, the gradient fields used for image encoding are linear. When the gradient fields are linear, converting the xSPEN data into an image can be as simple as performing a Fourier transform along the Fourier-encoded direction and taking the absolute value of the resulting image.

[0089] For radial xSPEN, there is no known method for reconstructing data acquired using a radial xSPEN sequence. The data should be convertible into images using techniques similar to those used to reconstruct x-ray images, but this has not yet been reported.

[0090] In this regard, none of the existing reconstruction methods can handle nonlinear gradient magnetic fields. When the fields are nonlinear, images produced using existing reconstruction methods may be distorted. Nevertheless, if the images are to be used in clinical diagnosis, they must be free of distortion.

[0091] In one aspect, the present application describes a conjugate gradient least-squares reconstruction that uses a map of the gradient field to correct for distortions produced by gradient coil nonlinearities. This method is based on the magnetization signal equation generated by an xSPEN sequence, as described, for example, in connection with FIG. 10. This equation allows for distortion-free resolution by calculating and using different contributions to phase. This method can also be used to reconstruct radial xSPEN images, as described, for example, in connection with FIG. 11.

[0092] In some embodiments, there are two parts to consider during reconstruction. The ideal nonlinear xSPEN reconstruction and the residual phase term can be based on Equation 1. The collected time-domain signal is a one-dimensional profile along the spatiotemporal direction and is spatially encoded using the xSPEN readout gradient applied during the first two refocusing pulses. In this example, the spatiotemporal axis is along y, but it can be along any direction. The equations below assume that the other axis is Fourier encoded. An encoding matrix is ​​calculated using these equations. The undistorted image can then be restored by inverting the matrix.

[0093]

number

[0094] Equation (1) is exact if the first two chirped refocusing pulses are equal in length and the gradients applied during those pulses are identical. However, if the pulse used for excitation is also chirped, the duration of the two subsequent pulses will not be the same. The difference in their durations may depend on the parameters of the pulse sequence. The presence of a difference results in second-order and first-order (linear) residual phases that can affect image quality, especially in radial reconstruction. If the residual phase terms are not taken into account, the method may fail to produce an image completely.

[0095] The two residual phase terms are the quadratic term shown in equation (2) and the linear term shown in equation (3). Equation (1) multiplied by the residual phase terms is used to generate an encoding matrix that can be inverted to enable resolution.

[0096]

number

[0097] (3)φ linear =2T π (CF-1)YG z +2T π (P1CF+P2)YG y

[0098] Referring primarily to FIG. 12(B), xSPEN was applied to exchange a small matrix dimension (z), which is typically frequency encoded, for a large matrix dimension (y), which is typically phase encoded, to speed up imaging.

[0099] Referring primarily to Figure 13, chirped WURST pulses were used throughout to excite and refocus a broad bandwidth of the slice. The first refocusing pulse was 1.1 times longer than the subsequent refocusing pulse to compensate for both the quadratic phase of the subsequent pulse and the quadratic phase of the excitation pulse, resulting in a spin echo signal after the second refocusing pulse. This necessitates increasing the amplitude of the second Gy pulse to maintain the same amplitude / pulse duration product that sets the xSPEN bilinear curvature.

[0100] Referring primarily to FIG. 15(B), model-based reconstruction significantly reduces the geometric distortion inside the phantom and increases the signal uniformity.

[0101] Referring primarily to Figure 16(B), the sequence is a CPMG acquisition using the CHORUS technique, generating a spin echo after the second WURST refocusing pulse, with the bracketed section repeated 12 times. The xSPEN encoding gradient is applied with opposite polarity between the first two refocusing pulses, with different colors representing different in-plane radial angles.

[0102] Referring primarily to Figure 17(A), the numerical PSF is constructed by summing Gz-shifted bilinear phase profiles through the slice dimension. A central section of these phase profiles is shown to the left of each of the calculated in-plane PSFs, showing how the bilinear saddle point shifts during readout, causing a corresponding shift in the main lobe of the PSF.

[0103] Referring primarily to Figure 18(C), undersampling the number of projections results in the same streaking artifacts seen with conventional Fourier radial sampling, but by truncating the readouts, a well-resolved image is left in the center of the FOV without reducing spatial resolution.

[0104] Referring primarily to Figure 20, a phantom was surrounded by a material intended to mimic tissue. As the field of view (FOV) decreased, so did the signal present in the image, and the image was cropped. Finally, the field of view was smaller than the phantom. This also allowed for a shorter echo interval to be used, in this case from 5.4 ms to 3.0 ms, thereby reducing the read train duration of the full 24 echoes from 130 ms to 72 ms.

[0105] Referring primarily to Figure 21(B), the y-axis is Fourier encoded while the x- and z-axes are spatiotemporally encoded. A profile along x is collected during readout, while a z-profile is collected indirectly.

[0106] Referring primarily to FIG. 23(A), to calculate the point-spread function (PSF) for each subslice in the volume, high-resolution xSPEN spatial encoding bilinear phase maps are calculated and summed across each subslice width to obtain the PSF for each slice.

[0107] composition In some embodiments, this reconstruction can be used for any xSPEN sequence, whether Cartesian or radial. This reconstruction can be used with xSPEN sequences with the same or different refocusing pulses, and residual phases generated by various pulses are accounted for in Equations (2) and (3). This reconstruction can also accommodate RARE-style image acquisition. Reconstruction can be performed as follows: using only the ideal signal from Equation (1); using the ideal signal and residual first-order phase; using the ideal signal and second-order phase; and finally using the ideal signal and both second-order and first-order residual phases. In the above equation, the space-time axis is y and the Fourier axis is x, but they may be swapped. Alternatively, Equation (1) can be calculated with the x- and y-axes as space-time axes and no Fourier axis, as described above in connection with FIG. 11. Alternatively, reconstruction of a three-dimensional radial xSPEN sequence can be achieved by using Equation (1) in all three axes.

[0108] In some aspects, many implementations of this reconstruction are possible that balance accuracy with computational speed and memory requirements. For 2D reconstruction, a more accurate reconstruction that does not assume linearity of permanent gradients can be obtained by directly constructing a high-resolution three-dimensional phase matrix for each readout time and then summing the matrices in the slice dimension to obtain a numerical point spread function for each readout time. The point spread function can then be reduced to a matrix that can be resolved by normalized pseudoinverse, by conjugate gradient, or using any other iterative optimization algorithm.

[0109] In some aspects, this same approach is also applied to radial reconstruction scenarios, where the PSF is rotated between iterations. Any Fourier phase encoding applied in any dimension can be realized as a separate explicit matrix multiplication or by a non-uniform fast Fourier transform (NUFFT) before or after applying the point spread function matrix. Instead of summing bilinear phases across slices, the NUFFT also allows for the application of a permanent gradient-induced phase shift at each readout time and then a summation operation across the slices. This significantly reduces the memory size of the PSF matrix.

[0110] In some embodiments, 3D reconstruction is performed using a similar approach by directly computing the bilinear phase at high spatial resolution and then summing to obtain a PSF for each subslice of the 3D volume and for each readout time point. The NUFFT can then be applied to the Fourier phase encoding as in the 2D case. Also, as in the 2D case, the NUFFT can be used to apply a readout phase shift and then sum the subslices, significantly reducing memory requirements.

[0111] process In another aspect, provided herein is a process for applying multiple xSPEN read gradients to the single-plane magnetic resonance imaging system disclosed herein.

[0112] In some embodiments, an encoding matrix used to generate an image is generated by applying multiple xSPEN readout gradients, and the encoding matrix accounts for residual primary and secondary phases. In some embodiments, an encoding matrix used to generate an image is generated by applying multiple xSPEN readout gradients. In some embodiments, the encoding matrix accounts for residual primary and secondary phases. In some embodiments, the encoding matrix accounts for residual primary phase. In some embodiments, the encoding matrix accounts for residual secondary phase.

[0113] In some embodiments, the multiple xSPEN read gradients disclosed herein include a first xSPEN read gradient, a second xSPEN read gradient, a third xSPEN read gradient, and a fourth xSPEN read gradient. In some embodiments, the multiple xSPEN read gradients disclosed herein include a first xSPEN read gradient, a second xSPEN read gradient, and a third xSPEN read gradient. In some embodiments, the multiple xSPEN read gradients disclosed herein include a first xSPEN read gradient and a second xSPEN read gradient. In some embodiments, the multiple xSPEN read gradients disclosed herein include a first xSPEN read gradient.

[0114] In some embodiments, the processes disclosed herein further include applying a first xSPEN read gradient during the first refocusing pulse along a second axis orthogonal to the first axis using a single-sided gradient coil set.

[0115] In some embodiments, the processes disclosed herein further include applying a second xSPEN readout gradient along a second axis during a second refocusing pulse using a single-sided gradient coil set.

[0116] In some embodiments, the processes disclosed herein further include applying a third xSPEN readout gradient during the first refocusing pulse along a third axis using a single-sided gradient coil set.

[0117] In some embodiments, the processes disclosed herein further include applying a fourth xSPEN readout gradient during the second refocusing pulse along a third axis using a single-sided gradient coil set.

[0118] Various modifications to the embodiments described in this disclosure will be readily apparent to those skilled in the art, and the generic principles defined herein may be applied to other embodiments without departing from the spirit or scope of the disclosure. Thus, the claims are not intended to be limited to the embodiments shown herein, but are to be accorded the widest scope consistent with the disclosure, the principles, and novel features disclosed herein.

Claims

1. 1. A single-sided magnetic resonance imaging system comprising a housing, The housing includes: The front and a permanent magnet for providing a static magnetic field; the static magnetic field extends from the permanent magnet to a region of interest about a first axis that is perpendicular to the permanent magnet; further comprising a radio frequency transmitting coil; A single-sided gradient coil set is provided. the radio frequency transmit coil and the single-sided gradient coil set are located adjacent the front surface; The single-plane magnetic resonance imaging system further comprises: a radio frequency receiving coil; Equipped with a power supply, the power source is configured to pass current through at least one of the radio frequency transmit coil or the single-sided gradient coil set to generate an electromagnetic field in the region of interest, at least a portion of the region of interest being outside the front surface; The single-plane magnetic resonance imaging system further comprises a control circuit, the control circuit comprising: transmitting an excitation pulse using the radio frequency transmit coil; transmitting a first refocusing pulse using the radio frequency transmit coil; applying a first xSPEN read gradient during the first refocusing pulse along a second axis orthogonal to the first axis using the single-plane gradient coil set; performing phase encoding along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-sided gradient coil set; transmitting a second refocusing pulse using the radio frequency transmit coil; applying a second xSPEN readout gradient along the second axis during the second refocusing pulse using the single-sided gradient coil set; receiving data using the radio frequency receive coil; It is configured to Single-sided magnetic resonance imaging system.

2. 2. The single-plane magnetic resonance imaging system of claim 1, wherein the control circuitry is further configured to transmit any number of refocusing pulses using the radio frequency transmit coil before, during, or after the multiple refocusing pulses occurring simultaneously with the xSPEN read gradient.

3. the control circuitry is further configured to generate an encoding matrix for use in generating an image; The single-plane magnetic resonance imaging system of claim 1 , wherein the encoding matrix accounts for residual first-order and second-order phases.

4. The control circuit further comprises: transmitting an excitation pulse using the radio frequency transmit coil; transmitting a first refocusing pulse using the radio frequency transmit coil; applying a first xSPEN read gradient during the first refocusing pulse along a second axis orthogonal to the first axis using the single-plane gradient coil set; applying a second xSPEN readout gradient during the first refocusing pulse along a third axis orthogonal to the first axis and orthogonal to the second axis using the single-sided gradient coil set; transmitting a second refocusing pulse using the radio frequency transmit coil; applying a third xSPEN readout gradient during the second refocusing pulse along the second axis using the single-plane gradient coil set; applying a fourth xSPEN readout gradient during the second refocusing pulse along the third axis using the single-sided gradient coil set; receiving data using said radio frequency receiving coil; 10. The single-plane magnetic resonance imaging system of claim 1 configured to:

5. 5. The single-plane magnetic resonance imaging system of claim 4, wherein the control circuitry is further configured to transmit any number of refocusing pulses using the radio frequency transmit coil before, during, or after the multiple refocusing pulses occurring simultaneously with the xSPEN read gradient.

6. The control circuit further comprises: transmitting any number of additional xSPEN readout gradients during the first refocusing pulse along any number of additional imaging axes using any number of additional gradient coil sets; transmitting any number of additional xSPEN readout gradients during said second refocusing pulse along any number of additional imaging axes using any number of additional gradient coil sets; transmitting a fourth refocusing pulse using the radio frequency transmit coil; and 5. The single-plane magnetic resonance imaging system of claim 4 configured to:

7. 10. The single-plane magnetic resonance imaging system of claim 1, wherein the amplitude at the xSPEN read axis is non-uniform over time during the refocusing pulse.

8. 10. The single-plane magnetic resonance imaging system of claim 1, wherein the gradient shape at the xSPEN read axis is time-invariant during the refocusing pulse.

9. the control circuitry is further configured to generate an encoding matrix for use in generating an image; The single-plane magnetic resonance imaging system of claim 4 , wherein the encoding matrix accounts for residual first-order and second-order phases.

10. The single-plane magnetic resonance imaging system of claim 4 , wherein the control circuitry is configured to spatiotemporally encode along any axis during acquisition.

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

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

13. A method of generating a two-dimensional image by collecting a series of projections at 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 the pulse sequence.

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

16. A method for generating a three-dimensional image using a radial encoding scheme, in which the effective axis of the gradient array is rotated with the axis of the xSPEN readout gradient.

17. 10. The method of applying multiple xSPEN read gradients to a single-plane magnetic resonance imaging system according to claim 1.

18. applying the plurality of xSPEN read gradients to generate an encoding matrix used to generate an image; 18. The method of claim 17, wherein the encoding matrix accounts for residual first-order and second-order phases.

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

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

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

22. 22. The method of claim 21, further comprising applying a fourth xSPEN readout gradient along the third axis during the second refocusing pulse with the single-plane gradient coil set.

Images