Patent Publication Number: US-2023136830-A1

Title: Point-of-care magnetic resonance imaging system for lumbar puncture guidance

Description:
CROSS-REFERENCE TO RELATED APPLICATIONS 
     This application claims the benefit of U.S. Provisional Patent Application Ser. No. 63/015,266, filed on Apr. 24, 2020, and entitled “POINT OF CARE MAGNETIC RESONANCE IMAGER FOR LUMBAR PUNCTURE (LP) GUIDANCE,” which is herein incorporated by reference in its entirety. 
    
    
     STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH 
     This invention was made with government support under EB018976 awarded by the National Institutes of Health. The government has certain rights in the invention. 
    
    
     BACKGROUND 
     Lumbar punctures (“LP”) can be used for both diagnostic uses (e.g., sampling CSF) and therapeutic purposes (e.g., such as alleviating intercranial pressure (“ICP”) or delivering drugs or anesthetics). Diagnostically, LP is routine within large hospitals to rule out meningitis and encephalitis, diagnose subarachnoid hemorrhage, examine fevers with central nervous system signs and symptoms, diagnose CNS lymphoma, and provide prognostication in multiple sclerosis. Current therapeutic delivery focuses mainly on pain relief, but also on delivering chemotherapy and antibiotics. However intrathecal administration is poised to have an expanding role in delivering drugs to the CNS with genetic editing therapies (e.g. spinal muscular atrophy) and drugs that have poor blood brain barrier penetration. LP has also emerged as a premiere diagnostic tool for Alzheimer&#39;s Disease (“AD”), where protein CSF biomarkers have been useful in predicting future progression in patients with mild cognitive impairment. 
     Despite its ubiquitous use, the LP is considered difficult to teach because it is purely guided by palpation without visualization of the internal anatomy, leading to repeat attempts and/or avoidance in difficult cases. Image guidance with ultrasound and x-ray is possible, but ultrasound has poor depth resolution and cerebrospinal fluid (“CSF”) contrast, and radiation from x-ray complicates point-of-care (“POC”) use. 
     Thus, there remains a need to provide image guidance of LP procedures that is portable, lightweight, and low-cost in order to enable safe and routine POC use. 
     SUMMARY OF THE DISCLOSURE 
     The present disclosure addresses the aforementioned drawbacks by providing a magnet assembly for a portable magnetic resonance imaging (“MRI”) system. The magnet assembly includes a plurality of magnet blocks configured to create a single-sided permanent magnet. The plurality of magnet blocks are arranged in concentric rings in each of at least two layers to define a central aperture extending through the at least two layers, where the central aperture is sized to receive a medical instrument. 
     It is another aspect of the present disclosure to provide a portable MRI system that includes a magnet assembly, at least one gradient coil, and a radio frequency (“RF”) coil. The magnet assembly extends from an inner, patient-facing surface to an outer surface, and includes a plurality of magnet blocks configured to create a single-sided permanent magnet. The plurality of magnet blocks are arranged in concentric rings in each of at least two layers, where the arrangement of the plurality of magnet blocks is configured to optimize homogeneity over a target field-of-view for spinal imaging. A central aperture extends from the patient-facing surface to the outer surface and is sized to receive a medical instrument. The at least one gradient coil is arranged adjacent the outer surface of the magnet assembly, and the RF coil is arranged adjacent the patient-facing surface of the magnet assembly. 
     The foregoing and other aspects and advantages of the present disclosure will appear from the following description. In the description, reference is made to the accompanying drawings that form a part hereof, and in which there is shown by way of illustration a preferred embodiment. This embodiment does not necessarily represent the full scope of the invention, however, and reference is therefore made to the claims and herein for interpreting the scope of the invention. 
    
    
     
       BRIEF DESCRIPTION OF THE DRAWINGS 
         FIG.  1    is a schematic diagram showing a magnet for a portable MRI system used for spinal imaging for guiding the placement of a needle during a lumbar puncture (“LP”) procedure. 
         FIG.  2    is a diagram of a portable magnet in accordance with an embodiment. The magnet reaches 40 mT at the target ROI using 248 N52 NdFeB blocks. The magnetic field is shown with isocontours spaced 0.5 mT apart. 
         FIG.  3    is a side view of the portable magnet shown in  FIG.  2   . 
         FIG.  4    is a schematic block diagram of a portable MRI system in accordance with an embodiment. 
         FIG.  5    is a schematic diagram of a portable MRI system and an articulated arm in accordance with an embodiment. 
         FIG.  6    shows simulated field maps in three planes for the LP-guided magnet design. The dashed box represents the target ROI. In the ROI the mean field is ˜40 mT and the gradient is approximately 50 mT/m. 
         FIG.  7    shows simulated images based on the field maps of  FIG.  6    for the LP guided magnet initial design. The field pattern, supports some aliasing in the image, but the L4/L5 region is clear. Aliasing can be potentially addressed in magnet design and by using an Rx receive array. 
     
    
    
     DETAILED DESCRIPTION 
     Described here are a single-sided magnet and magnetic resonance imaging (“MRI”) system that are portable, lightweight, and low cost and may be used as a point-of care (“POC”) MRI device. The portable single-sided MRI system is low-field and may be used to perform three-dimensional (“3D”) imaging, two-dimensional (“2D”) imaging, or one-dimensional (“1D”) imaging. 
     The portable MRI system may be placed next to a patient during an operation or other medical procedure and, unlike conventional MRI systems, does not require the patient to be transported from a hospital bed to the MRI system and moved into the magnet of the MRI system. The portable MRI system has a lightweight magnet (e.g., less than 25 kg) and dimensions that allow it to easily be moved through doors and into tight spaces. 
     In an embodiment, the portable MRI system is also low cost, for example, by using magnet materials that only cost on the order of hundreds of dollars. The portable MRI system is configured to provide MRI of tissues such as the spinal cord. In an embodiment, imaging of the tissues is to a depth of 8 cm. 
     The disclosed portable, point-of-care MRI system may increase the utility of MRI by extending its reach and enabling applications such as providing MRI guidance of lumbar puncture (“LP”) procedures, whether for diagnostic or therapeutic purposes. 
     Advantageously, the portably MRI system reduces barriers to performing this LP and other important diagnostic and drug delivery procedures outside of central hospital settings, expanding its availability to a broader range of healthcare locations and increasingly inexpert staff. While already routinely used to diagnose several conditions, the need for routine LP is poised to dramatically expand in the diagnosis of Alzheimer&#39;s Disease, where protein CSF biomarkers accessed through LP have emerged as a premiere diagnostic tool for predicting future progression. LP is also expected to see increasing use for CNS delivery of gene editing therapies and drugs that have poor blood brain barrier penetration. 
     Although routinely performed, LP is conventionally guided by palpation without visualization of the internal anatomy. This can lead to clinician anxiety and avoidance and repeated puncture attempts in difficult cases. Ultrasound and x-ray have been used to guide LP, but both have significant shortcomings. Ultrasound imaging cannot see the CSF target and the ionizing nature of x-rays is difficult to manage in routine POC use. In contrast, the portable MRI system described in the present disclosure enables both streamlined LP procedures within a hospital and enables it to be performed outside central hospital settings or by inexpert staff. 
     In some embodiments, the portable MRI system described in the present disclosure integrates a lightweight (e.g., less than 25 kg) device for MRI of the spine and surrounding tissues with a mechanically-constrained 22-gauge needle insertion, thereby guaranteeing registration between the image and needle path without requiring real-time imaging of the needle. During use, the practitioner positions the portable MRI system&#39;s central ˜2 cm diameter needle insertion hole over the standard needle entry mark between L4 and L5 ( FIG.  1   ). Next, a set of anatomical T2 weighted images can be acquired, on which the expected path of the needle can be plotted. Based on the image plan, a mechanical needle guidance track is set to the planned path avoiding vertebrae and any calcifications in the ligamentum flavum. The maximum depth is planned to just enter the T2-bright subarachnoid space. With the angulation and depth-stop set, the practitioner manually pushes the needle in along this track. 
     The portable MRI system uses a built-in gradient for readout encoding (or slice-select encoding), and an external gradient coil (e.g., an approximately planar gradient coil) for phase-encoding. As described above, with the needle track mechanically registered to the portable MRI system, the practitioner can pre-visualize the needle path relative to the targeted subarachnoid space, avoiding vertebral bodies and calcifications of the ligaments. Since the target CSF is readily visible on T2 images, the target depth can be determined, and the user can set a mechanical stop on the insertion device to place the needle tip just inside the arachnoid space. This helps minimize disruption of the cauda equina fibers and avoid over-shooting the target and puncturing a vertebral disk. It may also reduce the frequency of post-lumbar puncture headache by reducing the number of attempts at puncture. Knowledge of the target depth provided when using the portable MRI system described in the present disclosure can also avoid the need for multiple removals of the needle&#39;s stylet to check for CSF as the needle advances. 
     Referring now to  FIG.  1   , in one embodiment, the portable, single-sided MRI system is configured for reduced field-of view spinal imaging and capable of high resolution one-dimensional (1D), for example, depth profiling, to three-dimensional (3D) imaging. The MRI system includes a lightweight single-sided permanent magnet.  FIG.  1    is a schematic diagram showing a magnet for a portable MRI system  100 . In  FIG.  1   , a single-sided magnet  102  (a B 0  magnet) has an inner, patient-facing surface  104 , an out surface  106 , and a central aperture  108  formed therebetween. The patient-facing surface  104  of the magnet  102  can have a generally planar shape to allow the magnet  102  to be positioned adjacent a patient&#39;s back  110 . When so arranged, the central aperture  108  allows access of a medical instrument, such as a needle, to advance to a target location at or near the patient&#39;s spine. For instance, the medical instrument can be advanced to a desired depth  112  (D) measured from the patient-facing surface  104  of the magnet  102  and located within the ROI  114  imaged by the MRI system  100 . The magnet  102  can be designed such that the sensitive ROI  114  is sized and positioned to encompass the spinal cord  116  when the magnet  102  is positioned adjacent the patient&#39;s back  110 . 
     In an embodiment, the magnet  102  is designed to closely fit adjacent the patient&#39;s back  104  in order to maximize the B 0  field strength. In some instances, the magnet  102  can be mounted on an articulating arm such that the magnet  102  can be arranged adjacent the patient&#39;s back  110  with a hands-free operation. Additionally or alternatively, the magnet  102  can be arranged adjacent the patient&#39;s back  110  and then secured in place by strapping the magnet  102  to the patient. 
     As discussed further below, the magnet  102  is designed from a plurality of rare-earth (e.g., NdFeB) permanent magnet blocks arranged in a layered configuration on a former. As a non-limiting example, the magnet  102  can be constructed to have two or more layers of magnet blocks, where the magnet blocks are arranged in concentric rings within each layer, as shown in  FIGS.  2  and  3   . In such configurations, the concentric rings are coaxial with the central aperture  108 . The magnet  102  has a transverse-oriented B 0  field with the imaging ROI  114 , which can be positioned to include part of the subject&#39;s spine and spinal cord  116 . Shown in  FIG.  1    for reference are an x-axis  118  and a z-axis  120 . 
     In an embodiment, the sensitive volume of the magnet  102  may extend 8 cm into the patient&#39;s back  110 , such that the imaging ROI  114  contains the patient&#39;s spinal cord  116  to facilitate guidance of a medical instrument, such as a needle. An MRI system utilizing magnet  102  may be used for spinal imaging over a 3D volume and may include gradient coils (not shown) placed external to the magnet  102  on the outer surface  106  and an RF coil (not shown), which in some embodiments may be positioned on the inner, patient-facing surface  104  of the magnet  102 . In an embodiment, the shape of the ROI  114  may be configured to match the Bi sensitivity profile of the RF coil used for imaging. 
     As described above, a guidance track  118  can be coupled to the magnet  102 , for instance by coupling the guidance track  118  to the central aperture  108  of the magnet  102 . The guidance track  118  provides a mechanical guide to advance a medical instrument, such as a needle  120 , along a trajectory whose position and orientation is known relative to the magnet  102  by way of coupling the guidance track  118  to the magnet  102 . By knowing the spatial relationship between the magnet  102  and the guidance track  118 , the portable MRI system  100  is capable of imaging the ROI  114  and visualizing or otherwise overlaying the trajectory defined by the guidance track  118  on the images of the ROI  114 . In this way, the precise placement of the needle  120  can be visualized prior to inserting the needle  120  into the guidance track  118  and advancing the needle  120  into the patient. 
       FIGS.  2  and  3    are diagrams showing a portable magnet  102  in accordance with an embodiment. In an embodiment, the portable magnet  102  may have dimensions so that it may be held by hand. As mentioned, the inner, patient-facing surface  104  is designed to be generally planar with the patient&#39;s back. As mentioned, the portable MRI magnet  102  may be used in an MRI system configured for reduced field-of view spinal imaging. 
       FIG.  4    is a schematic block diagram of a portable MRI system  400  in accordance with some embodiments described in the present disclosure. In  FIG.  4   , a schematic representation of the position of various elements in the MRI system  400  with respect to one another is shown using blocks rather than the specific shape described above. The MRI system  400  includes a magnet assembly  414  having a magnet  402 , gradient coils  404 , and an RF coil  406  disposed within a housing  408  and positioned in close proximity to or on (e.g., close fitting) a subject  410 . The magnet  402  is a single-sided magnet. As one non-limiting example, the single-sided magnet  402  is designed from a plurality of permanent magnet blocks (e.g., NdFeB permanent magnet blocks) arranged in a layered configuration on a former. The layered configuration can include two or more layers of permanent magnet blocks. As one example, the layered configuration can include two or more layers of permanent magnet blocks arranged in concentric rings within each layer. In one specific embodiment, such as the example shown in  FIGS.  2  and  3   , the layered configuration can include two layers of concentric rings. 
     Single-sided magnets typically have large field gradients moving away from the magnet surface. This built-in B 0  gradient may be used for readout and slice select encoding. The magnet  402  may also be designed to avoid very strong (e.g., greater than 1 T/m) gradients. Gradient coils  404  can also be configured to be positioned on the outer surface (e.g., surface  106  shown in  FIG.  1   ) of the magnet  402 . For example, a pair of gradient coils may be used to enable phase encoding on the other two directions that are orthogonal to the readout direction. The RF coil  406  is configured to be positioned on an inner surface (e.g., surface  104  shown in  FIG.  1   ) of the magnet  402 . The RF coil  406  may be used to provide excitation and RF signal detection. In other embodiments, separate RF coils can be provided for excitation and signal detection. As discussed further below, the assembly of the magnet  402 , gradient coils  404 , and RF coil  406  can each include a former (not shown) on which the permanent magnet blocks, gradient coils, and RF coil(s) are mounted. 
     A controller  412  is coupled to the magnet  402 , gradient coils  404 , and RF coil  406  and configured to control the operation of the magnet  402 , gradient coils  404 , and RF coil  406  to acquire MR images of the subject  410 . For example, controller  412  is configured to drive the gradient coils  404  and RF coil  406  for gradient waveform generation and RF waveform generation, respectively, using known hardware and methods. In addition, controller  412  is configured to record MR signals received by the RF coil  406  from the subject  410 . Controller  412  may also be configured to generate images based on the received MR signals using known reconstruction methods. 
     In an embodiment, such as the one shown in  FIG.  5   , the magnet assembly  414  containing the magnet  402 , gradient coils  404 , and RF coil  406  may be attached or otherwise coupled to an articulated arm  416 , that allows the magnet assembly  414  to be positioned and moved into place adjacent a patient&#39;s back  418 . Using the articulated arm, the magnet assembly  414  may be arbitrarily positioned allowing movement of the sensitive volume of the single-sided magnet  402  to the appropriate region to allow for image-guided placement of a medical instrument (e.g., a needle  120 ) for an LP procedure or other medical procedure. This embodiment may allow the magnet assembly  414  to be, for example, positioned on a moveable cart  420  that can be moved into a patient&#39;s room or bedside to facilitate imaging and performance of the LP procedure or other medical procedure without having to move the patient to a dedicated imaging room. In such an example, the controller  412  may be housed within the moveable cart  420 . Advantageously, the articulated arm  416  also allows the magnet assembly  414  to be moved into position whether the patient is in a sitting or supine position. 
     In use, the portable MRI system  400  allows for both imaging and mechanically constraining the needle insertion. Building these together ensures registration between the image and needle path without requiring real-time imaging of the needle. After marking the entry point (e.g., based on standard L4/L5 landmarks), the practitioner wheels the 20-30 kg MRI assembly  414  (on its stand) up to the patient, who can be in the standard left-lateral position placing the approximately 2 cm diameter central aperture  108  for the needle  120  over the mark. The magnet assembly  414  is supported by the stand (e.g., articulated arm  416 ) but can be additionally secured to the patient&#39;s back  418  with surgical tape. A 3-5 minute set of RARE T2 images of the L4/L5 area is acquired and the expected path of the needle is plotted on these images. Based on the image plan the mechanical needle guidance track  118  is translated and rotated into the proper position to achieve the planned path, including a depth stop to prevent overshooting the subarachnoid space. With the angulation and stop set, the practitioner manually pushes the needle  120  in along this mechanically set track. Since the insertion is manual (but mechanically confined to the chosen path) the practitioner feels the needle  120  passing the expected landmarks similar to an unguided procedure. 
     As mentioned above, the portable, single-sided magnet  102  (shown in  FIG.  1   ) is designed from a plurality of permanent magnet blocks arranged in a layered configuration on a former and may be used to obtain images of a spinal region of a subject that it is positioned adjacent. In an embodiment, the arrangement of the plurality of rare-earth permanent magnet blocks is optimized for the layered configuration. For example, the magnet may be designed with a genetic algorithm optimizing homogeneity over a field-of-view (“FOV”) and the built-in gradient for slice-selection or readout encoding. For example, the placement of the rare-earth magnet material (e.g., NdFeB) may be chosen using the genetic optimization framework. 
     In an embodiment, a magnet array containing N=248 N52-NdFeB magnets arranged in two layers in concentric rings and arcs was optimized, as shown in  FIGS.  2  and  3   . Individual magnet blocks were all 25.4×[y]×25.4 mm 3 , where the [y]-dimensions were numerically optimized in a continuous range of 0-25.4 mm. The magnet was designed with symmetry about the XY and XZ planes to produce an x-oriented B 0  field gradient (i.e., directed into the patient&#39;s back). 
     The sizes and angular orientations of the magnet blocks were optimized to produce a 4×6×8 cm 3  homogeneous magnetic field along the spine at a depth of 8 cm from the magnet surface (e.g., the patient-facing surface  104  in  FIG.  1   ) and overlaying the adult L3-L5 spine (as shown in  FIG.  1   ). As an example, the optimization can be performed by finding the minimum of a constrained nonlinear multivariable function. As a non-limiting example, the “fmincon” tool in MATLAB (Mathworks) can be used with: cost function−range of B 0  magnitude (fcost=range{|B−0|}), computed at points on a 1 cm 3  grid within the ROI; constraints−mean |B0|&gt;=40 mT, 0&lt;=y&lt;=25.4 mm. Magnet blocks can be modeled as L=5 multipole field sources for field computation during optimization. Simulated B 0  maps for the optimized design can be computed using Biot-Savart simulation software (e.g., Ripplon). 
     The magnet can be designed to acquire a ˜3 minute duration 3D RARE spin-echo image with the G x ˜50 mT/m built-in read-out gradient, G y  phase encoding along the echo-train, and G z  phase encoding shot-to-shot. A 2D imaging simulation was performed using the XZ field-map shown in  FIG.  6    for G x  and a linear planar gradient coil producing G z =8 mT/m peak for the Z phase encode. The simulation FOV is 4 cm×8 cm, with  256  readout points, BW=50 KHz, and 97 phase-encodes in Z.  FIG.  7    shows simulated images based on the above field maps for the LP guided magnet initial design. The simulation object is a T2-weighted MR image of the target ROI (around the subarachnoid space between L4 and L5). The simulated image exhibits aliasing due to the non-bijective pattern in the B 0  field-map. However, the L4/L5 region is clear and enables mechanical guidance via the needle track. The aliasing could potentially be addressed using multiple surface coils and/or by optimizing the magnet with a cost function that includes encoding performance. 
     In an embodiment, the continuous magnet material section of a Halbach magnet approximates the desired magnet shape (i.e., layers of concentric rings) and B 0  direction. The continuous magnetization pattern can then be discretized into a plurality of blocks. The discretized Halbach section approximates the continuous magnetization pattern as an assembly of magnet blocks. The discretized Halbach section is practical to construct and has the desired field orientation but is not optimized for in-plane homogeneity or gradient strength. In an embodiment, the optimization may be performed by allowing the genetic algorithm, or other optimization algorithm, to alter the sizes, block magnetization grade, compositions, and translational position of each magnet block. 
     The optimal magnet design may then be converted into a physically-realizable assembly of permanent magnet blocks, as shown in  FIGS.  2  and  3   . The example optimized magnet  102  shown in  FIGS.  2  and  3    is shown as an assembly of standard size and standard material NdFeB blocks  150 . The magnet assembly can include magnet blocks of various size and material combinations including N52 blocks of size 1×1×1⅛ in 3 , N52 blocks of size 1×1×1⅜ in 3 , N42 blocks of size 1×1×1 in 3 , N45 blocks of size 1×1×⅜ in 3 , or other such sized blocks as may be determined through the optimized magnet design. In an embodiment, some blocks may be constructed by sticking multiple smaller blocks together (e.g., an N52 1″×1″×1⅛″ block contained an N52 1″×1″×1″ block and an N52 1″×1″×⅛″ bock). 
     In the example shown in  FIGS.  2  and  3   , the magnet assembly is shown with 0.5 mT magnetic field iso-contours. The magnet measures 8×42×50 cm 3 , contains 24.2 kg of rare-earth material, and produces a mean field of 40.9 mT in the target ROI.  FIG.  6    shows the simulated magnetic field-maps in the three Cartesian planes with the dashed boxes indicating the target ROI. 
     A former is used to hold the magnet blocks prescribed by the optimized design. For example, the former can be constructed to include slots that are sized, positioned, and shaped to receive the magnetic blocks in an optimized magnet design. The magnet former may be constructed of a material such as acrylic. In an embodiment, the magnet former is constructed using 3D printing. The magnet former includes a plurality of slots. The final assembled magnet bocks are inserted into the slots of the magnet former and may be secured to the former using, for example an epoxy. 
     As mentioned, the portable magnet assembly may include a pair of gradient coils (e.g., gradient coils  404  shown in  FIG.  4   ). In an embodiment, one or more planar-shaped gradient coils are provided that are configured for phase encoding (e.g., phase encoding along the y- and z-axes). The gradient coils are constructed on a gradient coil former that may be positioned on an outer surface of the magnet. This design saves valuable space within the magnet to enable a stronger B 0  and allows for improved gradient linearity, at the cost of reduced gradient efficiency. In addition, weak unshielded gradient coils do not produce significant eddy current effects if placed either inside or outside an NdFeB magnet. 
     In an embodiment, the gradient winding patterns for the G y  and G z  gradient coils can be designed using a modified Boundary Element Method (“BEM”) stream function with L1-regularization. The target fields for the G y  and G z  coils can include both the desired linear terms (Y and Z, respectively) and an additional 2nd-order term (XY and XZ, respectively). The efficiency of a single-sided gradient coil decreases as one moves away from it (in this case, along the x-direction), and this decrease is manifested as undesired XY and XZ terms for the G y  and G z  coils, respectively. The addition of the 2nd-order terms in the target field of the BEM stream function design helps compensate for the spurious XY and XZ terms improving linearity over the target ROI. 
     The optimized stream functions can then be converted into wire winding paths. To construct the gradient coils, the optimized stream functions (winding paths) can be projected onto a piecewise-linear surface of a gradient coil former. In an embodiment, a gradient coil former may be constructed by 3D printing a polycarbonate disc, slab, or other appropriately shaped former (e.g., ˜2 mm thick). The former contains wire grooves that correspond to the numerical winding paths computed from the stream function. The wire grooves are configured to receive magnet wire, for example, the wire grooves may be configured for press-fitting two layers of magnet wire into the polycarbonate former. The completed G y  and/or G z  gradient coils and former assembly can then be positioned around a magnet, such as at the outer surface  106  of the magnet  102  shown in  FIG.  1   . 
     As mentioned, the portable magnet assembly may also include an RF coil (e.g., RF coil  136  shown in  FIG.  4   ). An RF coil assembly may be constructed by designing an RF coil (or winding) on a surface (or RF coil former) configured to fit inside the B 0  magnet. In an embodiment, the RF coil may be designed using the same BEM stream function approach described above with respect to the gradient coils. The same static-field approach used for the gradient design may be used because the RF coil dimension (˜0.1 m) may be much less than the wavelength at the Larmor frequency. An RF coil winding may be designed to optimize spatial Bi uniformity within the target ROI (e.g., ROI  114  shown in  FIG.  1   ). The RF coil may be constructed by press-fitting wire (e.g., Litz wire). 
     The MRI system described herein may be used as a point-of-care system to acquire 1D and 3D images, for example of the spine and spinal cord, over a reduced FOV sensitive region in order to facilitate the guidance of a medical instrument (e.g., a needle) during a procedure (e.g., an LP procedure). In an embodiment, a shimming capability may be applied to the B 0  magnet or the optimization may more explicitly penalize peak (“min-max”) inhomogeneities to mitigate any reduction of the slice thickness or signal level. In another embodiment, gradient non-linearity effects may be mitigated by refining the gradient coil design or compensated for in the pulse sequence by adjusting the encoded FOV for each slice. Other options include post-processing approaches which apply a gradient nonlinearity correction or generalized image reconstruction approach. 
     In an embodiment, control of the built-in B 0  gradient is used to provide equal amounts of signal per bandwidth at different positions in the ROI. An improved magnet design with higher linearity (but limited gradient strength) may be used to mitigate artifacts. Additionally, an RF coil with increased spatial uniformity and coverage may improve images. In an embodiment, this may be achieved by either a physically larger RF coil or an RF coil with more windings. Adding winding of increasing size boosts the inductance of the coil more quickly than the resistance. However, this in turn would increase the Q of the coil and decrease its bandwidth, exacerbating any coil BW issue. Resolving issues stemming from narrow coil bandwidth may be approached by shaping the spectral resonance response of the coil. On approach for creating an RF coil with a more uniform frequency response is using a series resistor. Several approaches for creating a coil with a more uniform frequency response without a series resistor include quasi-transmission line coils, coupled resonant structures, used of a low-impedance preamplifier, and inductively coupled negative feedback mechanisms. 
     In an embodiment, image signal-to-noise ratio (“SNR”) may be improved either with improved system hardware or an improved acquisition. For example, either a stronger B 0  magnet or more uniform B 0  magnet (enabling reduced-bandwidth acquisitions) would improve SNR. A stronger B 0  magnet in the same form factor may be achievable by allowing for a higher density of magnetic material or by adding an additional layer of magnet blocks or otherwise increasing the thickness of the magnet. A more uniform magnet may be realized by the use of B 0  shim coils or shim material. An improved RF coil may also be used to increase SNR. 
     In an embodiment, weighting the sampling density to the center of k-space or utilizing sparsity priors such as compressed-sensing type acquisitions or denoising approaches may be used to boost SNR. In another embodiment, for an acquisition using a RARE pulse sequence a flipback pulse after each RARE train may assist with longitudinal magnetization recovery and increase available signal. 
     In an embodiment, the portable MRI system may also include shielding, for example, either a passive shielding approach such as draped conductive cloth, or an active interference cancellation system. In another embodiment, to address temperature induced drift in B 0 , various approaches may be used including a feedback system controlling a heater to stabilize the temperature, use of a combination of rare-earth materials with differing temperature coefficients, or the use of a field probe to measure B 0  drift for incorporation into a model-based image reconstruction algorithm. 
     As mentioned, the portable B 0  magnet MRI system may be used for reduced-FOV imaging of the spine and spinal cord to guide the placement of a medical instrument, such as a needle, during a procedure. 
     The B 0  magnet may be designed by optimizing the distribution of rare-earth magnets needed to maximize homogeneity over a target ROI. In an embodiment, an interior point method may be used to optimize magnet block size (and thus magnetic dipole size) for a layered Halbach geometry. The three components of a magnetic dipole moment vector are optimized at points on a planar or other suitably shaped surface that can be arranged adjacent a patient&#39;s back to design a magnet assembly that minimizes the absolute range of B 0  magnitude over a target ROI. In a non-limiting example, the optimization can implement a minimum mean B 0 , and constrain all magnetic dipole moment vector magnitudes be less than that of a 1″×1″×1″ block of N53 magnet material. The optimization used an initial guess solution. For example, the optimization may use a “test-tube magnet” as an initial guess solution. In this optimization, each magnet block in the assembly can be modeled as an ideal point dipole source. Next, each dipole moment vector in the optimized solution can be uniformly scaled up until the dipole moment with the largest magnitude matched that of a 1″×1″×1″ block of N52-grade NdFeB material. A design was then generated containing the prescribed number of non-intersecting N52 magnet blocks of differing volume, such that each block&#39;s magnetic dipole moment matched that generated by the numerical optimization. 
     Computer-executable instructions for optimizing the design of a portable magnet and MRI system and for operating a portable MRI system according to the above-described methods may be stored on a form of computer readable media. Computer readable media includes volatile and nonvolatile, removable, and non-removable media implemented in any method or technology for storage of information such as computer readable instructions, data structures, program modules or other data. Computer readable media includes, but is not limited to, random access memory (RAM), read-only memory (ROM), electrically erasable programmable ROM (EEPROM), flash memory or other memory technology, compact disk ROM (CD-ROM), digital volatile disks (DVD) or other optical storage, magnetic cassettes, magnetic tape, magnetic disk storage or other magnetic storage devices, or any other medium which can be used to store the desired instructions and which may be accessed by a system (e.g., a computer), including by internet or other computer network form of access. 
     The present disclosure has described one or more preferred embodiments, and it should be appreciated that many equivalents, alternatives, variations, and modifications, aside from those expressly stated, are possible and within the scope of the invention.