Method for measuring hysteresis in MRI system, apparatus for controlling at least one gradient coil of MRI system, and computer-readable storage medium and program product

Abstract

The invention relates to a method for measuring hysteresis in an MRI system, an apparatus for controlling at least one gradient coil of an MRI system, and a computer-readable storage medium and a program product. The apparatus may include at least one computer hardware processor and at least one computer readable storage medium to store processor executable instructions that, when executed by the at least one computer hardware processor, cause the at least one computer hardware processor to perform a method. The method may include: receiving information specifying at least one target pulse sequence; determining a correction pulse sequence to control the at least one gradient coil based on the at least one target pulse sequence and a hysteresis model of induced magnetization in the MRI system caused by operation of the at least one gradient coil; and controlling the at least one gradient coil using the corrected gradient pulse sequence to generate one or more gradient pulses for imaging the patient.

Description

[0001] This application is a divisional application of application number 201980093021.3, filed on 15 November 2019, having the title “Correcting magnetic hysteresis in magnetic resonance imaging”. BACKGROUND

[0002] Magnetic resonance imaging (MRI) provides an important imaging modality for many applications and is widely used in clinical and research settings to produce images of the interior of the human body. Generally, MRI is based on detecting magnetic resonance (MR) signals, which are electromagnetic waves emitted by atoms in response to a change in state induced by an applied electromagnetic field. For example, nuclear magnetic resonance (NMR) techniques involve detecting MR signals emitted from the nuclei of excited atoms as the nuclear spins of atoms in the subject being imaged (e.g., atoms in human tissue) realign or relax. The detected MR signals can be processed to produce images, which enable investigation of internal structures and / or biological processes in the body for diagnostic, therapeutic, and / or research purposes in the context of medical applications.

[0003] MRI provides an attractive imaging modality for biological imaging because it is able to produce non-invasive images with relatively high resolution and contrast without the safety concerns of other modalities (e.g., without the need to expose the subject to ionizing radiation (e.g., x-rays) or to introduce radioactive material into the body). In addition, MRI is particularly well suited to provide soft tissue contrast, which can be useful for imaging subjects that other imaging modalities cannot satisfactorily image. Moreover, MRI techniques are able to capture information about structures and / or biological processes that other modalities cannot obtain. However, MRI has a number of drawbacks, which for a given imaging application can include a relatively high cost of equipment, limited availability, and / or difficulty in obtaining access to a clinical MRI scanner, as well as / or the length of the image acquisition process.

[0004] A trend in clinical MRI has been to increase the field strength of MRI scanners to improve one or more of scan time, image resolution, and image contrast, which in turn continues to drive up costs. The vast majority of installed MRI scanners operate at 1.5 or 3 Tesla (T), where 1.5 or 3 Tesla refers to the field strength of the main magnetic field, B0. A rough cost estimate for a clinical MRI scanner is about one million dollars per Tesla, which does not take into account the substantial operational, service, and maintenance costs involved in operating such a MRI scanner.

[0005] Additionally, conventional high-field MRI systems typically require large superconducting magnets and associated electronics to generate a strong uniform static magnetic field (B0) that images a subject (e.g., a patient). The size of such systems is quite large, with a typical MRI installation including multiple rooms for the magnet, electronics, thermal management system, and control console areas. The size and expense of MRI systems generally limit their use to facilities, such as hospitals and academic research centers, that have sufficient space and resources to purchase and maintain them. The high cost and large space requirements of high-field MRI systems result in limited availability of MRI scanners. As a result, as described in further detail below, there are often clinical situations where an MRI scan would be beneficial, but is not practical or possible due to one or more of the limitations discussed above. SUMMARY

[0006] Some embodiments include an apparatus for controlling at least one gradient coil of a magnetic resonance imaging system (MRI system), the apparatus comprising: at least one computer hardware processor; and at least one computer-readable storage medium storing processor-executable instructions for causing the at least one computer hardware processor, upon execution of the instructions by the at least one computer hardware processor, to perform a method. The method comprises: receiving information specifying at least one target pulse sequence; determining a correction pulse sequence to control the at least one gradient coil based on a hysteresis model of induced magnetization in the MRI system caused by operation of the at least one gradient coil and the at least one target pulse sequence; and controlling the at least one gradient coil using the correction pulse sequence to generate one or more gradient pulses for imaging a patient.

[0007] Some embodiments include a method for controlling at least one gradient coil of a magnetic resonance imaging system (MRI system), the method comprising: with at least one computer hardware processor: receiving information specifying at least one target pulse sequence; determining a correction pulse sequence to control the at least one gradient coil based on a hysteresis model of induced magnetization in the MRI system caused by operation of the at least one gradient coil and the at least one target pulse sequence; and controlling the at least one gradient coil using the correction pulse sequence to generate one or more gradient pulses for imaging a patient.

[0008] Some embodiments include at least one computer-readable storage medium storing processor-executable instructions that, when executed by at least one computer hardware processor, cause the at least one computer hardware processor to perform a method comprising: receiving information specifying at least one target pulse sequence; determining a correction pulse sequence to control the at least one gradient coil based on a hysteresis model of the at least one target pulse sequence and induced magnetization in the MRI system caused by operation of the at least one gradient coil; and using the correction pulse sequence to control the at least one gradient coil to generate one or more gradient pulses for imaging a patient.

[0009] Some embodiments include a method for measuring hysteresis in a magnetic resonance imaging system (MRI system) comprising at least one gradient coil, the method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring a first plurality of magnetic field strengths in an imaging region of the MRI system using a multi-element RF probe placed in the imaging region of the MRI system, each magnetic field strength in the first plurality of magnetic field strengths being at least partially caused by a respective pulse in the first plurality of pulses of the first pulse sequence; estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths; and storing the parameters of the hysteresis model.

[0010] Some embodiments include at least one computer-readable storage medium storing processor-executable instructions that, when executed by at least one computer hardware processor, cause the at least one computer hardware processor to perform a method for measuring hysteresis in a magnetic resonance imaging system (MRI system) comprising at least one gradient coil, the method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring a first plurality of magnetic field strengths in an imaging region of the MRI system using a multi-element RF probe placed in the imaging region of the MRI system, each magnetic field strength in the first plurality of magnetic field strengths being at least partially caused by a respective pulse in the first plurality of pulses of the first pulse sequence; estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths; and storing the parameters of the hysteresis model.

[0011] Some embodiments include an apparatus for controlling at least one gradient coil of a magnetic resonance imaging (MRI) system, the apparatus comprising: at least one computer hardware processor; and at least one computer-readable storage medium storing processor-executable instructions that, when executed by the at least one computer hardware processor, cause the at least one computer hardware processor to perform a method. The method comprises: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring a first plurality of magnetic field strengths in an imaging region of the MRI system using a multi-element RF probe placed in the imaging region of the MRI system, each magnetic field strength of the first plurality of magnetic field strengths being caused at least in part by a respective pulse of the first plurality of pulses of the first pulse sequence; estimating parameters of a magnetic hysteresis model based on the measured first plurality of magnetic field strengths; and storing the parameters of the magnetic hysteresis model.

[0012] Some embodiments include a multi-element probe for measuring magnetic hysteresis in a magnetic resonance imaging (MRI) system, the multi-element probe comprising: an RF transmit coil; a plurality of RF receive elements; and a plurality of liquid samples, each liquid sample being contained within a respective coil of the plurality of RF receive elements.

[0013] Some embodiments include a method for measuring magnetic hysteresis in a magnetic resonance imaging (MRI) system, the MRI system comprising at least one electromagnet, the method comprising: measuring a magnetic field in an imaging region of the MRI system using a multi-element probe. The multi-element probe comprises: an RF transmit coil; a plurality of RF receive elements; and a plurality of liquid samples, each liquid sample being contained within a respective coil of the plurality of RF receive elements. BRIEF DESCRIPTION OF DRAWINGS

[0014] Various aspects and embodiments of the disclosed technology will be described with reference to the following figures. It should be understood that the drawings are not necessarily to scale as the illustrations are for the purpose of clarity in illustrating the various aspects and embodiments.

[0015] Figure 1 Example components of an illustrative magnetic resonance imaging system;

[0016] Figure 2A An illustrative B0 magnet including a plurality of permanent magnets, in accordance with some embodiments;

[0017] Figure 2B An illustrative portion of a B0 magnet, in accordance with some embodiments; Figure 2A A top view of an illustrative structure of a ring of permanent magnets of a B0 magnet;

[0018] Figure 3 An illustrative B0 magnet including a plurality of permanent magnets, in accordance with some embodiments;

[0019] Figure 4 Example partially formed Figure 3 Top view of an example structure of a permanent magnet ring of a B0 magnet shown;

[0020] Figure 5A And 5B Example example ring of permanent magnet segments for a B0 magnet according to some embodiments;

[0021] Figure 5C And 5D Example different views of permanent magnet segments that can be used to form Figure 5E a permanent magnet ring shown;

[0022] Figure 5E Example permanent magnet ring for a B0 magnet according to some embodiments;

[0023] Figure 5F And 5G Example different views of permanent magnet segments that can be used to form Figure 5H a permanent magnet ring shown;

[0024] Figure 5H Example permanent magnet ring for a B0 magnet according to some embodiments;

[0025] Figure 6A And 6B Example portable low-field MRI system according to some embodiments;

[0026] Figure 7 Drive circuitry for driving current through a coil to produce a magnetic field according to some embodiments of the technology described herein is shown;

[0027] Figure 8 Example computer system according to some embodiments;

[0028] Figure 9A Example gradient pulse sequence according to some embodiments;

[0029] Figure 9B Example measured hysteresis curve according to some embodiments;

[0030] Figure 10A Example hysteron model of magnetic particles according to some embodiments;

[0031] Figure 10B Example Preisach model of hysteresis effects based on multiple hysteron models according to some embodiments;

[0032] Figure 11A Example example time-dependent applied magnetic field according to some embodiments;

[0033] Figure 11B Illustrating application of Preisach model according to some embodiments;

[0034] Figure 11C Illustrating example time dependent applied magnetic field according to some embodiments;

[0035] Figure 11D Illustrating application of Preisach model according to some embodiments;

[0036] Figure 12 Illustrating Preisach model based on multiple hysteresis operator models of hysteresis effect according to some embodiments;

[0037] Figure 13A Illustrating schematic of multi-element probe according to some embodiments;

[0038] Figure 13B Illustrating receiving element of multi-element probe according to some embodiments;

[0039] Figure 13C is a top view of a multi-element probe according to some embodiments;

[0040] Figure 13D is a perspective view of a multi-element probe according to some embodiments;

[0041] Figure 13E is a side view of a multi-element probe according to some embodiments;

[0042] Figure 13F is an end view of a multi-element probe according to some embodiments;

[0043] Figure 13G is a perspective view of a multi-element probe according to some embodiments;

[0044] Figure 13H is a perspective view of a multi-element probe according to some embodiments;

[0045] Figure 14 Illustrating method for measuring induced hysteresis effect in MRI system according to some embodiments;

[0046] Figure 15A Illustrating part of pulse sequence for controlling MRI system according to some embodiments;

[0047] Figure 15B Illustrating example gradient crush amplitude sequence according to some embodiments;

[0048] Figure 16 Illustrating measured constant and gradient magnetic field hysteresis effect according to some embodiments;

[0049] Figure 17A Illustrating a Preisach model with hysteresis effects fitted to weights according to some embodiments;

[0050] Figure 17B Illustrating hysteresis vs. gradient curves according to some embodiments;

[0051] Figure 17C Illustrating hysteresis vs. time curves according to some embodiments;

[0052] Figure 18 Illustrating a method for controlling at least one gradient coil of a magnetic resonance imaging (MRI) system according to some embodiments;

[0053] Figure 19A Illustrating example MRI images produced from an MRI system without correction for hysteresis effects; and

[0054] Figure 19B Illustrating example MRI images produced from an MRI system with correction for hysteresis effects according to some embodiments. DETAILED DESCRIPTION

[0055] High-field systems dominate the MRI scanner market, particularly for medical or clinical MRI applications. As discussed above, a general trend in medical imaging is to produce MRI scanners with increasingly large field strengths, with the vast majority of clinical MRI scanners operating at 1.5T or 3T, and higher field strengths of 7T and 9T being used in research settings. As used herein, “high-field” generally refers to MRI systems currently used in clinical settings, and more specifically to MRI systems operating with a main magnetic field (i.e., the B0field) of 1.5T or above, although clinical systems operating between 0.5T and 1.5T are often also characterized as “high-field.” Field strengths between about 0.2T and 0.5T have been characterized as “mid-field,” and as field strengths in the high-field regime continue to increase, field strengths in the range between 0.5T and 1T have also been characterized as mid-field. In contrast, “low-field” generally refers to MRI systems operating with a B0field of less than or equal to about 0.2T, although due to the increase in field strengths at the high end of the high-field regime, systems with a B0field between 0.2T and about 0.3T have sometimes also been characterized as low-field. Within the low-field regime, low-field MRI systems operating with a B0field of less than 0.1T are referred to herein as “very low-field,” and low-field MRI systems operating with a B0field of less than 10mT are referred to herein as “ultra low-field.”

[0056] As discussed above, conventional MRI systems require specialized facilities. MRI systems require an electromagnetically shielded room to operate, and the floor of the room must be structurally reinforced. Additional rooms must be provided for high-power electronics and a control area for the scanning technician. Secure access to the site must also be provided. In addition, a dedicated three-phase electrical connection must be installed to power the electronics, which in turn must be cooled by a chilled water supply, and additional HVAC capacity must be provided. These site requirements are not only expensive, but also significantly limit the locations where MRI systems can be deployed. Conventional clinical MRI scanners also require a great deal of expertise to operate and maintain both. These trained technicians and service engineers add significant ongoing operating costs to operating MRI systems. As a result, conventional MRI is prohibitively expensive and severely limited in accessibility, hindering MRI as a widely available diagnostic tool that can deliver a wide range of clinical imaging solutions wherever and whenever needed. Patients must visit one of a limited number of facilities at a pre-arranged time and place, hindering the use of MRI in many medical applications where it is uniquely effective, such as in ancillary diagnosis, surgery, and patient monitoring.

[0057] As discussed above, high-field MRI systems require specially adapted facilities to accommodate the size, weight, power consumption, and shielding requirements of these systems. For example, a 1.5T MRI system typically weighs between 4-10 tons, while a 3T MRI system typically weighs between 8-20 tons. In addition, high-field MRI systems generally require a large amount of heavy and expensive shielding. Many mid-field scanners are even heavier, weighing between 10-20 tons, in part because very large permanent magnets and / or magnetic yokes are used. Commercially available low-field MRI systems (e.g., operating at a Bo magnetic field of 0.2T) are also typically in the range of 10 tons or more due to the large amount of ferromagnetic material used to generate the Bo field, in addition to the tonnage used for shielding. To accommodate such heavy equipment, the room must be built with reinforced flooring (e.g., concrete flooring), typically a minimum size of 30-50 square meters, and the room must be specially shielded to prevent electromagnetic radiation from interfering with the operation of the MRI system. As a result, available clinical MRI systems are immobile and require a significant amount of dedicated space within a hospital or facility, in addition to the considerable cost to prepare the space for operation, and additional ongoing costs in terms of expertise to operate and maintain the system.

[0058] Additionally, currently available MRI systems generally consume a large amount of electrical power. For example, common 1.5T and 3T MRI systems typically consume 20-40 kW of electrical power during operation, while available 0.5T and 0.2T MRI systems typically consume 5-20 kW. When discussing power consumption herein, unless otherwise specified, reference will be made to average power consumed over an interval of interest. For example, the 20-40 kW noted above indicates the average power consumed by a conventional MRI system during the course of an image acquisition, which can include relatively short periods of peak power consumption that significantly exceed the average power consumption (e.g., periods in which the gradient coils and / or radio frequency (RF) coils are pulsed for relatively short periods of the pulse sequence). The intervals of peak (or large) power consumption are generally accommodated via power storage elements (e.g., capacitors) of the MRI system itself. Thus, average power consumption is a more relevant property, as it generally dictates the type of electrical connection required to operate the device. As a result, available clinical MRI systems must also have a dedicated power supply, typically requiring a dedicated three-phase connection to the electrical grid to operate the MRI system. Additional electronics are then required to convert the three-phase power to the single-phase power utilized by the MRI system. The many physical requirements for deploying a conventional clinical MRI system create significant availability issues, and strictly limit the clinical applications in which MRI can be utilized.

[0059] Thus, many of the requirements of high-field MRI make installation costs prohibitive in many cases, limiting deployment to large institutional hospitals or specialized facilities, and generally limiting their use to strictly scheduled appointments, requiring patients to visit a dedicated facility at a pre-scheduled time. Thus, many of the limitations of high-field MRI hinder the full use of MRI as an imaging modality. Despite the drawbacks of high-field MRI noted above, the call for significant increases in SNR at higher fields continues to drive the industry toward ever-higher field strengths for clinical and medical MRI applications, further increasing the cost and complexity of MRI scanners, and further limiting their availability and hindering their use as a general and / or common imaging solution.

[0060] The low SNR of MR signals produced in the low field region, particularly the very low field region, has hindered the development of relatively low cost, low power, and / or portable MRI systems. Conventional "low field" MRI systems operate at the high end of what is typically characterized as the low field range (e.g., clinically useful low field systems have a minimum of about 0.2T) to achieve useful images. While somewhat less expensive than high field MRI systems, conventional low field MRI systems share many of the same drawbacks. In particular, conventional low field MRI systems are large, fixed and immobile installations that consume large amounts of power (requiring dedicated three-phase electrical wiring diagrams), require specially shielded rooms and large dedicated spaces. The challenges of low field MRI have hindered the development of relatively low cost, low power, and / or portable MRI systems capable of producing useful images.

[0061] The inventors have developed techniques that enable portable, low field, low power, and / or low cost MRI systems, which can improve the widespread deployability of MRI technology in a variety of environments beyond the MRI installations currently in hospitals and research facilities. As a result, MRI can be deployed in emergency rooms, small clinics, physician offices, mobile units, on-site, etc., and can be taken to patients (e.g., at the bedside) for a wide variety of imaging procedures or protocols. Some embodiments include very low field MRI systems (e.g., 0.1T, 50mT, 20mT, etc.) that facilitate portable, low cost, low power MRI, thereby significantly increasing the usability of MRI in clinical environments.

[0062] There are many challenges to developing a clinical MRI system in the low field region. As used herein, the term "clinical MRI system" refers to an MRI system that produces clinically useful images, where a clinically useful image is one that has sufficient resolution and appropriate acquisition time to be useful to a physician or clinician for its intended purpose given a particular imaging application. As such, the resolution / acquisition time of a clinically useful image will depend on the purpose for which the image is being obtained. One of the many challenges to obtaining clinically useful images in the low field region is the relatively low SNR. In particular, the relationship between SNR and B0 field strength is approximately B0 5 / 4 at field strengths above 0.2T, and approximately B0 3 / 2As such, the SNR drops substantially with decreasing field strength, and the SNR experiences even more significant drop at very low field strengths. This significant drop in SNR due to decreasing field strength is an important factor that has hindered the development of clinical MRI systems in the very low field region. In particular, the challenge of low SNR at very low field strengths has hindered the development of clinical MRI systems that operate in the very low field region. As a result, clinical MRI systems have been limited to the range of about 0.2T and above, where these systems are large, heavy (and expensive) MRI systems that consume a large amount of electrical power, generally require a fixed dedicated space (or shielded tent) and a dedicated power supply.

[0063] The present inventors have recognized and appreciated that, in order to develop low field and very low field MRI systems that are capable of producing clinically useful images, the SNR should be increased. Noise sources and errors that have traditionally been ignored in high field MRI systems (as these errors are small relative to the strength of the magnetic field in high field systems) can reduce the SNR of low field MRI systems and significantly impact the quality of the resulting images. As such, these noise sources and errors create problems that have not been addressed in the development of high field MRI systems, and are important to address in low field MRI systems to ensure the highest possible SNR.

[0064] The present inventors have recognized that the hysteresis effect due to induced magnetization in ferromagnetic materials reduces the SNR and / or causes imaging errors in low field MRI systems. Many components of an MRI system and the associated environment are formed from, or include, ferromagnetic materials that are susceptible to magnetization induced by one or more electromagnets of the MRI system. When the ferromagnetic materials of the various components are magnetized, the resulting magnetic fields affect the magnetic fields in the imaging region of the MRI system, which includes the B0field, linear gradient fields, and higher order terms. In high field MRI systems, the magnetic fields resulting from the induced magnetization of the ferromagnetic materials of the various components of the system can be ignored as these magnetic fields are small relative to the high field strength of the magnet of the MRI system. However, the hysteresis effect should be addressed in low field MRI systems to increase the SNR.

[0065] For example, a low field MRI system can include a permanent magnet plate connected by a ferromagnetic yoke (see, e.g., FIG. 1 below) Figure 2Amagnetic system. Such low-field MRI systems include one or more gradient coils for producing gradient magnetic fields in an imaging region of the low-field MRI system. A gradient pulse sequence is used to control the magnetic fields generated by the gradient coils such that the magnetic field in the imaging region varies as a function of time. The dynamic magnetic fields generated by the gradient coil(s) induce magnetization in the permanent magnet plates, ferromagnetic yokes, and other components of the MRI system that are made of ferromagnetic material. The induced magnetization is a hysteresis effect that produces a magnetic field in the imaging region in addition to the magnetic field produced by the gradient coil(s). Thus, the actual magnetic field present in the imaging region is no longer precisely controlled by the gradient coil(s), but is the sum of the magnetic field generated by the gradient coil(s) and the magnetic field generated by the induced magnetization (as well as the magnetic field produced by other magnets of the MRI system, such as the B0magnet). Not precisely controlling the magnetic field in the imaging region degrades the SNR and, in turn, the quality of the MRI images.

[0066] One approach to reducing the errors caused by the hysteresis effect in low-field MRI systems is to include a demagnetization pulse during the imaging routine. However, the addition of a demagnetization pulse increases the time required to perform the imaging routine, and / or requires additional gradient correction lobes, which provide an incomplete correction for the hysteresis effect. The present inventors have developed techniques for reducing imaging errors caused by the hysteresis effect that do not increase the duration of the imaging routine by adding additional demagnetization pulses.

[0067] The present inventors have recognized that the hysteresis effects induced in a low-field MRI system can be measured prior to imaging a patient using the low-field MRI system. In turn, these measurements can be used to produce a hysteresis model of the hysteresis effects that can be used to compensate for these effects during subsequent imaging. For example, the hysteresis model can be used to modify a target pulse sequence to determine a correction pulse sequence to control the electromagnets (e.g., gradient coils) of the MRI system to produce a desired magnetic field strength in the imaging region of the MRI system. In some embodiments, the desired magnetic field strength can reflect the contributions of the induced magnetization of the MRI system and the magnetic field produced by the electromagnets controlled by the correction pulse sequence because the hysteresis effects are accounted for when determining the pulse sequence to control the electromagnets.

[0068] Accordingly, in some embodiments, a hysteresis model is used to determine a correction pulse sequence for controlling one or more electromagnets of an MRI system. Note that it is not necessary to determine the exact location of the induced magnetization. For example, a yoke, other magnetic components, a housing, electronic components, and / or any number of other components of the MRI system can be made of ferromagnetic material and can exhibit a hysteresis effect. It is not necessary to determine the amount of induced magnetization associated with each component. By taking measurements of the hysteresis effect prior to imaging route selection, the induced magnetization from any ferromagnetic material present at the time of measurement can be collectively accounted for by the hysteresis model. Using this technique, the inventors have observed an increase in SNR of about 15%.

[0069] In some embodiments, an apparatus for controlling at least one gradient coil of an MRI system includes: at least one computer hardware processor; and at least one computer-readable storage medium storing processor-executable instructions that, when executed by the at least one computer hardware processor, cause the at least one computer hardware processor to perform a method including: receiving at least one target gradient pulse sequence; determining a correction gradient field pulse sequence to drive the at least one gradient coil based on the at least one target gradient pulse sequence and a history of at least one control setting for the at least one gradient coil; and controlling the at least one gradient coil using the correction gradient pulse sequence.

[0070] In some embodiments, the MRI system includes a ferromagnetic yoke, and the hysteresis model represents an effect of hysteresis induced by operation of the at least one gradient coil at least in the ferromagnetic yoke.

[0071] In some embodiments, determining the correction pulse sequence can include adjusting an amplitude of a pulse within a target gradient pulse sequence of the target pulse sequence. Alternatively or additionally, determining the correction pulse sequence can include determining a correction transmit RF pulse sequence for controlling a radio frequency (RF) transmit coil and / or a correction receive RF pulse sequence for controlling a RF receive coil. For example, determining the correction transmit RF pulse sequence can include adjusting a center frequency or a phase of a transmit RF pulse of the correction transmit RF pulse sequence; and determining the correction receive RF pulse sequence can include adjusting a center frequency or a phase of a receive RF pulse of the correction receive RF pulse sequence.

[0072] The inventors have further recognized that a multi-element probe can be placed within an MRI system to measure magnetic hysteresis effects. The multi-element probe includes a plurality of receive elements for measuring magnetic fields within an imaging region of the MRI system. These measurements are then used to determine a model of the magnetic hysteresis effects induced in the MRI system. The magnetic hysteresis effects induced in the MRI system depend on the history of the pulse sequence applied to the electromagnet. Thus, the magnetic fields in the imaging region caused by induced magnetization can be measured for a plurality of different pulse sequences.

[0073] In some embodiments, a multi-element probe for measuring magnetic hysteresis in an MRI system, the multi-element probe comprising: an RF transmit coil; a plurality of RF receive elements; and a plurality of liquid samples, each liquid sample contained within a respective coil of the plurality of RF receive elements.

[0074] Some embodiments include a method for measuring magnetic hysteresis in a magnetic MRI system, the MRI system comprising at least one electromagnet, the method comprising: measuring magnetic fields in an imaging region of the MRI system using a multi-element probe, the multi-element probe comprising: an RF transmit coil; a plurality of RF receive elements; and a plurality of liquid or gel samples, each liquid or gel sample contained within a respective coil of the plurality of RF receive elements.

[0075] Some embodiments include a method for measuring magnetic hysteresis in an MRI system, the MRI system comprising at least one gradient coil, the method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring a first plurality of magnetic field strengths in an imaging region of the MRI system using a multi-element RF probe placed in the imaging region of the MRI system, each magnetic field strength of the first plurality of magnetic field strengths induced at least in part by a respective pulse of the first plurality of pulses of the first pulse sequence; estimating parameters of a magnetic hysteresis model based on the measured first plurality of magnetic field strengths; and storing the parameters of the magnetic hysteresis model.

[0076] Figure 1 is a block diagram of typical components of an MRI system 100. In Figure 1 an illustrative example, the MRI system 100 includes a computing device 104, a controller 106, a pulse sequence storage 108, a power management system 110, and a magnetic assembly 120. It should be understood that the system 100 is illustrative, and that the MRI system can have one or more other components of any suitable type in addition to or instead of the components shown. However, the MRI system will generally include these high-level components, although the implementation of these components can vary greatly for a particular MRI system, as discussed in further detail below. Figure 1 Figure 1 It should be understood that the system 100 is illustrative, and that the MRI system can have one or more other components of any suitable type in addition to or instead of the components shown. However, the MRI system will generally include these high-level components, although the implementation of these components can vary greatly for a particular MRI system, as discussed in further detail below.​

[0077] like Figure 1 As shown, the magnetic assembly 120 includes a BO magnet 122, a shim coil 124, an RF transmit and receive coil 126, and a gradient coil 128. The magnet 122 can be used to generate the main magnetic field BO. The magnet 122 can be any suitable type of magnetic element or a combination thereof that can generate the desired main magnetic field BO. As discussed above, in the high field region, the BO magnet is typically formed using superconducting materials generally provided in a solenoid geometry, thereby requiring a cryogenic cooling system to maintain the BO magnet in a superconducting state. Therefore, high-field BO magnets are expensive, complex, and consume a large amount of electricity (for example, the cryogenic cooling system requires a large amount of electricity to maintain the extremely low temperature required to maintain the BO magnet in a superconducting state), requiring a large dedicated space and a dedicated dedicated power connection (for example, a dedicated three-phase power connection to the power grid). Traditional low-field BO magnets (for example, BO magnets operating at 0.2 T) are also typically implemented using superconducting materials and have these same general requirements. Other conventional low-field B0 magnets are implemented using permanent magnets. To generate the field strengths that limit conventional low-field systems due to their inability to acquire useful images at lower field strengths (e.g., between 0.2 T and 0.3 T), these permanent magnets must be very large, weighing 5-20 tons. Consequently, the B0 magnets of conventional MRI systems inherently hinder both portability and affordability.

[0078] The gradient coils 128 can be arranged to provide gradient fields and, for example, can be arranged to generate gradients in the B0 field in three generally orthogonal directions (X, Y, Z). The gradient coils 128 can be configured to encode the transmitted MR signals by systematically varying the B0 field (generated by the magnet 122 and / or the shim coils 124) to encode the spatial location of the received MR signals as a function of frequency or phase. For example, the gradient coils 128 can be configured to vary the frequency or phase as a linear function of the spatial location along a particular direction, although more complex spatial encoding profiles can also be provided by using nonlinear gradient coils. For example, a first gradient coil can be configured to selectively vary the B0 field in a first (X) direction to provide frequency encoding in that direction, a second gradient coil can be configured to selectively vary the B0 field in a second (Y) direction generally orthogonal to the first direction to provide phase encoding, and a third gradient coil can be configured to selectively vary the B0 field in a third (Z) direction generally orthogonal to the first and second directions to enable slice selection for volumetric imaging applications. As discussed above, conventional gradient coils also consume significant amounts of power and are typically operated by large, expensive gradient power supplies, as discussed in further detail below.

[0079] MRI is performed by separately using transmit and receive coils, often referred to as radio frequency coils, to excite and detect the transmitted MR signals. The transmit / receive coils can include separate coils for transmitting and receiving, multiple coils for transmitting and / or receiving, or the same coil for both transmitting and receiving. Thus, the transmit / receive assembly can include one or more coils for transmitting, one or more coils for receiving, and / or one or more coils for both transmitting and receiving. Transmit / receive coils are often also referred to as Tx / Rx coils to generically refer to the various structures of the transmit and receive magnetic assemblies of the MRI system. These terms are used interchangeably herein. In Figure 1 The RF transmit and receive coils 126 include one or more transmit coils that can be used to generate RF pulses to induce an oscillating magnetic field Bi. The transmit coil(s) can be configured to generate any suitable type of RF pulses.

[0080] The power management system 110 includes electronics to provide operating power to one or more components of the low-field MRI system 100. For example, as discussed in more detail below, the power management system 110 can include one or more power supplies, gradient power assemblies, transmit coil assemblies, and / or any other suitable power electronics needed to provide suitable operating power to energize and operate the components of the MRI system 100. As shown, the power management system 110 includes a power supply 112, power assembly(s) 114, a transmit / receive switch 116, and a thermal management assembly 118 (e.g., a cryogenic cooling device for a superconducting magnet). Figure 1 The power supply 112 includes electronics to provide operating power to the magnetic assemblies 120 of the MRI system 100. For example, the power supply 112 can include electronics to provide operating power to one or more B0 coils (e.g., B0 magnet 122) to produce the main magnetic field of the low-field MRI system. The transmit / receive switch 116 can be used to select whether the RF transmit coils are being operated or the RF receive coils are being operated.

[0081] The power assembly(s) 114 can include one or more RF receive (Rx) preamplifiers for amplifying MR signals detected by one or more RF receive coils (e.g., coils 126), one or more RF transmit (Tx) power assemblies configured to power one or more RF transmit coils (e.g., coils 126), one or more gradient power assemblies configured to power one or more gradient coils (e.g., gradient coils 128), and one or more shim power assemblies configured to power one or more shim coils (e.g., shim coils 124).

[0082] In traditional MRI systems, the power electronics are large, expensive, and consume significant amounts of power. Typically, the power electronics occupy a separate room from the MRI scanner itself. Not only do these power electronics require a large amount of space, but they are also expensive, complex devices that consume significant power and require a wall-mounted rack to support them. Consequently, the power electronics of traditional MRI systems can hinder the portability and affordability of MRI.

[0083] like Figure 1 As shown, the MRI system 100 includes a controller 106 (also referred to as a console) having control electronics for sending instructions to and receiving information from a power management system 110. The controller 106 can be configured to implement one or more pulse sequences for determining instructions sent to the power management system 110 for operating the magnetic assembly 120 in a desired sequence (e.g., parameters for operating the RF transmit and receive coils 126, parameters for operating the gradient coils 128, etc.). Figure 1 As shown, the controller 106 also interacts with a computing device 104 that is programmed to process the received MR data. For example, the computing device 104 can process the received MR data to generate one or more MR images using any suitable (one or more) image reconstruction processes. The controller 106 can provide information related to one or more pulse sequences to the computing device 104 so that the data can be processed by the computing device. For example, the controller 106 can provide information related to one or more pulse sequences to the computing device 104, and the computing device can perform image reconstruction processing based at least in part on the provided information. In conventional MRI systems, the computing device 104 typically includes one or more high-performance workstations configured to perform computationally expensive processing on MR data relatively quickly. Such a computing device itself is a relatively expensive device.

[0084] In some embodiments, the controller 106 may include a computer hardware processor (not shown) and at least one computer-readable storage medium (not shown) storing processor-executable instructions that, when executed by the at least one computer hardware processor, cause the at least one computer hardware processor to perform one or more methods. The methods may include methods for determining a pulse sequence to be used to control the MRI system 100. Additionally or alternatively, the methods may include methods for measuring hysteresis effects induced in the MRI system 100 and / or determining a hysteresis model associated with the MRI system 100.

[0085] In some embodiments, the controller 106 can be configured to implement the pulse sequence by obtaining information relating to the pulse sequence from a pulse sequence storage 108 that stores information for each of one or more pulse sequences. The information stored by the pulse sequence storage 108 for a particular pulse sequence can be any suitable information that allows the controller 106 to implement the particular pulse sequence. For example, the information stored in the pulse sequence storage 108 for a pulse sequence can include one or more parameters for operating the magnetic assembly 120 in accordance with the pulse sequence (e.g., parameters for operating the RF transmit and receive coils 126, parameters for operating the gradient coils 128, etc.), one or more parameters for operating the power system 110 in accordance with the pulse sequence, one or more programs that include instructions that, when executed by the controller 106, cause the controller 106 to control the MRI system 100B to operate in accordance with the pulse sequence, and / or any other suitable information. The information stored in the pulse sequence storage 108 can be stored on one or more non-transitory storage media.

[0086] In some embodiments, the pulse sequence storage 108 can also store information relating to the measured hysteresis induced in the MRI system 100. For example, the measurement data itself can be stored. Additionally or alternatively, a hysteresis model can be stored in the pulse sequence storage 108. Alternatively, in other embodiments, the information relating to hysteresis can be stored in a storage separate from the pulse sequence storage 108.

[0087] As should be appreciated from the foregoing, currently available clinical MRI systems, including high field, mid field, and low field systems, are large, expensive, fixed installations that require a significant amount of dedicated and specially designed space and dedicated power connections. The present inventors have developed lower cost, lower power, and / or portable low field (including very low field) MRI systems, thereby significantly increasing the availability and applicability of MRI. According to some embodiments, a portable MRI system is provided, thereby allowing the MRI system to be taken to the patient and used at the location where the MRI system is needed.

[0088] As discussed above, some embodiments include a portable MRI system, allowing the MRI device to be moved to locations where the MRI system is needed (e.g., emergency rooms and operating rooms, primary care offices, neonatal intensive care units, specialty, and emergency and mobile transport vehicles and field locations). The development of a portable MRI system presents a number of challenges, including size, weight, power consumption, and the ability to operate in a relatively uncontrolled electromagnetic noise environment (e.g., outside of a specially shielded room). As discussed above, currently available clinical MRI systems range from about 4 to 20 tons. Thus, currently available clinical MRI systems are not portable, as the size and weight of the imaging equipment itself is large, not to mention the fact that currently available systems also require a substantial amount of dedicated space, including a specially shielded room to house the MRI scanner and additional rooms to house the power electronics and technician control areas, respectively. The present inventors have developed MRI systems of suitable weight and size to allow the MRI system to be transported to a desired location, some examples of which are discussed in further detail below.

[0089] Another aspect of portability relates to the ability to operate the MRI system in a wide variety of locations and environments. As discussed above, currently available clinical MRI scanners require to be located in a specially shielded room to allow for proper operation of the device, and this is one of the (many) reasons for the cost, lack of availability, and non-portability of currently available clinical MRI scanners. Thus, to operate outside of a specially shielded room, and more particularly to allow for a generally portable, vehicle-mounted, or otherwise transportable MRI to be implemented, the MRI system must be able to operate in a variety of noise environments. The present inventors have developed noise suppression techniques that allow the MRI system to operate outside of a specially shielded room, facilitating both portable / transportable MRI and the installation of fixed MRI without a specially shielded room. While the noise suppression techniques allow for operation outside of a specially shielded room, these techniques can also be used to suppress noise in a shielded environment (e.g., a less expensive, loose, or ad-hoc shielded environment), and thus can be used in conjunction with areas that are equipped with limited shielding, as these aspects are not limited in this regard.

[0090] Another aspect of portability relates to the power consumption of MRI systems. As also discussed above, current clinical MRI systems consume a large amount of power (e.g., average power consumption ranging from 20 kW to 40 kW during operation), and thus require a dedicated power connection (e.g., a dedicated three-phase power connection to the electrical grid capable of delivering the required power). The requirement of a dedicated power connection is another obstacle to operating MRI systems in various locations other than expensive dedicated rooms specially equipped with appropriate power connections. The present inventors have developed low-power MRI systems that can operate using mains power such as standard wall outlets (e.g., 120 V / 20 A connections in the United States) or public large appliance outlets (e.g., 220-240 V / 30 A), allowing the device to work anywhere that a public power outlet is provided. The ability to "plug into the wall" facilitates installation of both portable / transportable MRI as well as fixed MRI systems without the need for special dedicated power sources such as three-phase power connections.

[0091] As discussed above, a significant contributor to the size, cost, and power consumption of conventional MRI systems is the power electronics used to power the magnetic components of the MRI system. The power electronics used for conventional MRI systems often require a separate room, are expensive, and consume a large amount of power to operate the respective magnetic components. In particular, the gradient coils and the thermal management system used to separately cool the gradient coils generally require a dedicated power connection and are prohibited from operation from a standard wall outlet. The present inventors have developed low-power, low-noise gradient power supplies that can power the gradient coils of an MRI system, which according to some embodiments can be housed in the same portable, vehicle-mounted, or otherwise transportable device as the magnetic components of the MRI system. According to some embodiments, the power electronics used to power the gradient coils of an MRI system consume less than 50 W when the system is idle and between 100-200 W when the MRI system is operating (i.e., during image acquisition). The present inventors have developed power electronics (e.g., low-power, low-noise power electronics) to operate portable low-field MRI systems that fit within the footprint of a portable MRI scanner. According to some embodiments, the innovative mechanical design has enabled the development of MRI scanners that are easy to operate in the range of various clinical environments where the system is needed.

[0092] The core of developing low-power, low-cost, and / or portable MRI systems is to reduce the field strength of the B0magnet, which can help reduce the size, weight, expense, and power consumption. However, as discussed above, reducing the field strength has a corresponding and significant reduction in SNR. This significant reduction in SNR has prevented clinical MRI systems from reducing the field strength below the current lower limit of about 0.2T, which are still large, heavy, expensive, fixed installations requiring dedicated and specialized space. Although some systems have been developed to operate between 0.1T and 0.2T, these systems are specialized devices often used to scan extremities such as hands, arms, or knees. The present inventors have developed MRI systems operating in low and very low fields that are capable of acquiring clinically useful images. Some embodiments include highly efficient pulse sequences in generating MR signals, and / or some embodiments include optimized receive coils for detecting the emitted MR signals, examples of which are discussed in further detail below. As discussed in further detail above and below, measuring and modeling the magnetic hysteresis effects induced in the MRI system, and then using the magnetic hysteresis model to determine a correction pulse sequence, helps reduce imaging errors resulting from changes in the magnetic field in the imaging region caused by the induced magnetic field.

[0093] According to some embodiments, the designs developed by the present inventors also reduce the cost and complexity of operating and maintaining the MRI scanner. For example, conventional clinical MRI systems require a significant amount of expertise for both operation and maintenance, which results in significant ongoing costs for these systems. The present inventors have developed MRI systems that are easy to use, which allow minimally trained or untrained personnel to operate and / or maintain the system. According to some embodiments, an automatic setup process allows the MRI scanner to automatically detect and adapt to its environment in preparation for operation. Network connectivity allows the MRI system to be operated from a mobile device (such as a tablet, laptop, or smartphone) with an easy-to-use interface configured to automatically run a desired scan protocol. The acquired images are immediately transferred to a secure cloud server for data sharing, telemedicine, and / or deep learning. Furthermore, in some embodiments, an easy-to-use multi-element probe is used to measure the magnetic hysteresis effects induced in the MRI system. An automatic measurement process allows the MRI system to automatically measure and determine a magnetic hysteresis model for the MRI system. Furthermore, the determination of the correction pulse sequence is automatic, such that no special expertise is required to reduce errors caused by the induced magnetization of the ferromagnetic components of the MRI system.

[0094] The various concepts related to lower cost, lower power, and / or portable low-field MRI and embodiments thereof are explained in greater detail below. It should be understood that the embodiments described herein can be implemented in any of numerous ways. Examples of specific implementations are provided for illustrative purposes only. It should be understood that the embodiments and features / capabilities provided can be used alone or in any combination of two or more, since the aspects of the technology described herein are not limited in this regard.

[0095] A significant contributor to the high cost, size, weight, and power consumption of high-field MRI is the B0magnet itself along with the equipment needed to power and thermally manage the B0magnet. In particular, to produce the field strength characteristics of high-field MRI, the B0magnet is typically implemented as an electromagnet configured in a solenoid geometry using superconducting wire that requires a cryogenic cooling system to maintain the wire in a superconducting state. Not only is the superconducting material itself expensive, but the cryogenic equipment used to maintain the superconducting state is also expensive and complex.

[0096] The inventors have recognized that the low-field environment allows for B0magnet designs that are not feasible in the high-field regime. For example, due at least in part to the lower field strength, superconducting material and the corresponding cryogenic cooling system can be eliminated. Due in part to the low field strength, a B0electromagnet constructed using non-superconducting material (e.g., copper) can be employed in the low-field regime. However, such an electromagnet can still consume a relatively large amount of power during operation. For example, operating an electromagnet using copper conductors to generate a magnetic field of 0.2T or greater requires a dedicated or specialized power connection (e.g., a dedicated three-phase power connection). The inventors have developed an MRI system that can be operated using mains power (i.e., standard wall power), thereby allowing the MRI system to be powered at any location having a common power connection, such as a standard wall outlet (e.g., a 120V / 20A connection in the United States) or a common large appliance outlet (e.g., 220-240V / 30A), among others. Thus, the low-power MRI system facilitates portability and availability, allowing the MRI system to be operated at a location where the MRI system is needed (e.g., the MRI system can be brought to the patient rather than bringing the patient to the MRI system), examples of which are discussed in further detail below. Additionally, operating from standard wall power eliminates the electronics that are traditionally needed to convert three-phase power to single-phase power and to smooth the power provided directly from the grid. Instead, the wall power can be converted directly to DC and distributed to power the components of the MRI system.

[0097] Figure 2AAn example permanent B0 magnet according to some embodiments is illustrated. In particular, B0 magnet 200 is formed from permanent magnets 210a and 210b arranged in a bi-planar geometry, and a magnetic yoke 220 that captures the electromagnetic flux produced by the permanent magnets and conveys that flux to the opposing permanent magnet to increase the flux density between permanent magnets 210a and 210b. Each of permanent magnets 210a and 210b is formed from a plurality of concentric permanent magnets. In particular, as shown, permanent magnet 210b includes an outer ring of permanent magnets 214a, a middle ring of permanent magnets 214b, an inner ring of permanent magnets 214c, and a permanent magnet disk 214d at the center. Permanent magnet 210a can include the same set of permanent magnet elements as permanent magnet 210b. Figure 2A

[0098] The permanent magnet material used can be selected according to the design requirements of the system. For example, according to some embodiments, the permanent magnets (or a portion thereof) can be made of NdFeB, which produces a magnetic field having a relatively high magnetic field per unit volume of material once magnetized. According to some embodiments, SmCo material is used to form the permanent magnets or a portion thereof. While NdFeB produces a higher field strength (and is generally less expensive than SmCo), SmCo exhibits less thermal drift and thus provides a more stable magnetic field in the face of temperature fluctuations. Other types of permanent magnet material(s) can also be used as these aspects are not limited in this regard. Generally, the type(s) of permanent magnet material utilized will depend at least in part on the field strength, temperature stability, weight, cost, and / or ease of use requirements of a given B0 magnet implementation.

[0099] The size of the permanent magnet rings is determined and arranged so that a uniform field of a desired strength is produced in the central region (field of view) between permanent magnets 210a and 210b. In the example embodiment shown, the permanent magnet rings are formed from a plurality of segments, each segment being formed using a plurality of blocks stacked in a radial direction, and positioned around a perimeter in an abutting fashion with one another to form the respective ring. The inventors have recognized that by varying the width of each permanent magnet (in a direction tangential to the ring), less waste of useful space can be achieved while using less material. For example, by varying the width of the blocks (e.g., according to the radial position of the blocks), the space between stacks that do not produce a useful magnetic field can be reduced, allowing for a tighter fit to reduce wasted space, and maximizing the amount of magnetic field that can be generated in a given space. As discussed in further detail below, the dimensions of the blocks can also be varied in any desired fashion to facilitate the production of a magnetic field of a desired strength and uniformity. Figure 2A

[0100] ​​The B0magnet 200 also includes a magnetic yoke 220 configured and arranged to capture the magnetic flux generated by the permanent magnets 210a and 210b and direct the magnetic flux to the opposite side of the B0magnet to increase the flux density between the permanent magnets 210a and 210b and thereby increase the field strength within the field of view of the B0magnet. By capturing the magnetic flux and directing it to the region between the permanent magnets 210a and 210b, less permanent magnet material can be used to achieve the desired field strength, thereby reducing the size, weight, and cost of the B0magnet. Alternatively, for a given permanent magnet, the field strength can be increased, thereby improving the SNR of the system without having to use an increased amount of permanent magnet material. For the exemplary B0magnet 200, the magnetic yoke 220 includes a frame 222 and plates 224a and 224b. The plates 224a and 224b capture the magnetic flux generated by the permanent magnets 210a and 210b and direct the magnetic flux to the frame 222 to circulate via the magnetic return of the yoke, thereby increasing the flux density in the field of view of the B0magnet. The magnetic yoke 220 can be constructed of any desired ferromagnetic material (e.g., low carbon steel, CoFe, and / or silicon steel, etc.) to provide the desired magnetism to the yoke. According to some embodiments, the plates 224a and 224b (and / or the frame 222 or portions thereof) can be constructed of silicon steel, etc. in regions where eddy currents can most commonly be induced by the gradient coils.

[0101] The exemplary frame 222 includes arms 223a and 223b attached to the plates 224a and 224b, respectively, and supports 225a and 225b that provide a magnetic return for the flux generated by the permanent magnets. The arms are generally designed to reduce the amount of material needed to support the permanent magnets while providing sufficient cross-section for the return of the magnetic flux generated by the permanent magnets. The arm 223a has two supports within the magnetic return of the B0field generated by the B0magnet. The supports 225a and 225b are created in a manner that creates a gap 227 between the two, thereby providing stability to the frame and / or a measure of light weighting of the structure while providing sufficient cross-section for the magnetic flux generated by the permanent magnets. For example, the cross-section needed for the return of the magnetic flux can be divided between the two support structures, thereby providing sufficient return while increasing the structural integrity of the frame. It should be understood that additional supports can be added to the structure as the technology is not limited to only two supports and any particular number of support structures.

[0102] As discussed above, exemplary permanent magnets 210a and 210b include a plurality of rings of permanent magnetic material arranged concentrically with the permanent magnet disks at the center. Each ring can include a plurality of stacks of ferromagnetic material to form the respective ring, and each stack can include one or more blocks (and / or in some embodiments, a single block in some rings), which can be of any number. The size and arrangement of the blocks forming each ring can be such that a desired magnetic field is produced. The present inventors have recognized that, according to some embodiments, the size of the blocks can be such that cost is reduced, weight is reduced, and / or uniformity of the produced magnetic field is improved in many ways, as discussed in further detail in connection with exemplary rings of permanent magnets that collectively form a B0 magnet.

[0103] Figure 2B A top-down view of an exemplary permanent magnet 210, which can be used, for example, as Figure 2A The design of permanent magnets 210a and 210b of B0 magnet 200 shown. Permanent magnet 210 includes concentric rings 210a, 210b, and 210c, each of which is constructed from a plurality of stacks of ferromagnetic blocks and a ferromagnetic disk 210d at the center. The orientation of the frame of the magnetic yoke to which the permanent magnet is attached is indicated by arrow 22. In embodiments where the magnetic yoke (e.g., magnetic yoke 220) is not symmetric, it will cause the magnetic field produced by the permanent magnet for which it traps and concentrates the magnetic flux to also be asymmetric, negatively impacting the uniformity of the B0 magnetic field.

[0104] According to some embodiments, the block size is varied to compensate for the effect of the magnetic yoke on the magnetic field produced by the permanent magnet. For example, Figure 2B The size of the blocks in the four regions 215a, 215b, 215c, and 215d labeled in the center can vary according to the region in which the respective block is located. In particular, the height of the blocks (e.g., the dimension of the blocks perpendicular to the plane of the circular magnet 210) can be greater in region 215c, which is farthest from the frame, than in the respective blocks in region 215a, which is closest to the frame. The block height can vary in one or more rings or portions thereof, as the techniques to compensate for the effect of the magnetic yoke are not limited to changing any particular blocks, groups of blocks, and / or any particular size. One example of varying the block size to compensate for the yoke effect is discussed in further detail below.

[0105] Figure 3 An exemplary B0 magnet 300 according to some embodiments is illustrated. B0 magnet 300 can be similar to B0 magnet 200, except that the permanent magnets 310a and 310b are constructed from a plurality of rings of permanent magnetic material arranged concentrically with the permanent magnet disks at the center, each of which is constructed from a plurality of stacks of ferromagnetic blocks and a ferromagnetic disk at the center. The orientation of the frame of the magnetic yoke to which the permanent magnet is attached is indicated by arrow 22. In embodiments where the magnetic yoke (e.g., magnetic yoke 220) is not symmetric, it will cause the magnetic field produced by the permanent magnet for which it traps and concentrates the magnetic flux to also be asymmetric, negatively impacting the uniformity of the B0 magnetic field. Figure 2A and Figure 2BThe illustrated B0 magnet 200 shares design components. In particular, the B0 magnet 300 is formed from permanent magnets 310a and 310b arranged in a bi-planar geometry, with a magnetic yoke 320 coupled to the permanent magnets 310a and 310b to capture the magnetic flux generated by the permanent magnets and convey that flux to the opposing permanent magnet, thereby increasing the flux density between the permanent magnets 310a and 310b. Each of the permanent magnets 310a and 310b is formed from a plurality of concentric permanent magnets, as illustrated by permanent magnet 310b, which includes an outer ring permanent magnet 314a, a middle ring permanent magnet 314b, an inner ring permanent magnet 314c, and a permanent magnet disk 314d at the center. The permanent magnet 310a can include the same set of permanent magnet elements as the permanent magnet 310b. The permanent magnet material used can be selected according to the design requirements of the system (e.g., NdFeB, SmCo, etc. according to the desired properties).

[0106] The size of the permanent magnet rings is determined and arranged so that a uniform field of the desired strength is generated in the central region (field of view) between the permanent magnets 310a and 310b. In particular, the permanent magnets 310a and 310b are arranged so that the magnetic flux generated by the permanent magnets 310a and 310b is captured by the magnetic yoke 320 and conveyed to the opposing permanent magnet, thereby increasing the flux density between the permanent magnets 310a and 310b. Figure 3 In the illustrated example embodiment, as discussed in further detail below, each permanent magnet ring includes a plurality of circular arc segments that are sized and positioned to generate the desired B0 magnetic field. The circular arc segments are arranged in a bi-planar geometry, with a magnetic yoke coupled to the permanent magnets to capture the magnetic flux generated by the permanent magnets and convey that flux to the opposing permanent magnet, thereby increasing the flux density between the permanent magnets. Figure 2A and Figure 2B In a similar manner to the magnetic yoke 220 illustrated, the magnetic yoke 320 is configured and arranged to capture the magnetic flux generated by the permanent magnets 310a and 310b and direct that magnetic flux to the opposite side of the B0 magnet to increase the flux density between the permanent magnets 310a and 310b. Thus, the magnetic yoke 320 increases the field strength within the field of view of the B0 magnet with less permanent magnet material, thereby reducing the size, weight, and cost of the B0 magnet. The magnetic yoke 320 also includes a frame 322 and plates 324a and 324b that capture the magnetic flux generated by the permanent magnets 310a and circulate that magnetic flux via the magnetic circuit of the yoke to increase the magnetic flux density in the field of view of the B0 magnet. The structure of the magnetic yoke 320 can be similar to that described above to provide sufficient material to accommodate the magnetic flux generated by the permanent magnets and provide sufficient stability while minimizing the amount of material used, e.g., to reduce the cost and weight of the B0 magnet.

[0107] Figure 4 A top-down view of an example permanent magnet 410, which can be used, for example, as the permanent magnet 310a or 310b of the B0 magnet 300 is illustrated. Figure 3300. The permanent magnets 410a and 410b of the B0 magnet 300 are shown. The permanent magnet 410 includes concentric rings 410a, 410b, and 410c, each composed of a plurality of circular arc segments of ferromagnetic material, and a ferromagnetic disk 410d at the center. In embodiments where the yoke (e.g., yoke 320) is asymmetric, the yoke will cause the magnetic field generated by the permanent magnet to which it captures and converges the magnetic flux to also be asymmetric, thereby negatively affecting the uniformity of the B0 magnetic field. According to some embodiments, one or more dimensions of the circular arc segments are changed to compensate for the effect of the yoke on the magnetic field generated by the permanent magnet. For example, as discussed in further detail below, Figure 4 One or more dimensions of the arc segments in the four quadrants 415a, 415b, 415c, and 415d marked therein may be varied to compensate for the effect of the magnetic yoke on the B0 magnetic field.

[0108] Figure 5A and 5B Inner ring 510 (eg, Figure 4 The exemplary ring 510 includes a plurality of ( Figure 5A and 5B ferromagnetic arc segments (e.g., segments formed of NdFeB, SmCo, etc.) are provided in the exemplary ring 510 (e.g., eight in the illustrated exemplary ring 510). In the exemplary ring 510, the arc segments (e.g., exemplary arc segment 505) are sized to provide a ring having an inner radius R1 and an outer radius R2, and a height or depth z8. According to some embodiments, the dimensions of the inner ring 510 are: R1 is between 45-47 mm (e.g., 46.08 mm), R2 is between 62-64 mm (e.g., 62.91 mm), and z8 is between 22-25 mm (e.g., 23.46 mm). It should be understood that the number of arc segments and their dimensions can be selected as desired to produce a desired B0 magnetic field (e.g., a desired field strength and / or uniformity), as these aspects are not limited in this regard.

[0109] Figure 5C and 5D Examples can be used to form Figure 5E The middle ring 510 shown (e.g., Figure 4 For example, segment 515 may be used to provide a different view of segment 515 of ring 410b). Figure 5E The segments in the illustrated quadrants Q1 to Q4 (again, for example, Figure 4 The exemplary portion 1815' comprises a plurality of ferromagnetic arc segments (eg, segments formed of NdFeB, SmCo, etc.). Figure 5C to 5E, two arc segments (e.g., exemplary arc segment 505'), each spanning 45°, form a quadrant of ring 510'. In exemplary portion 515' of ring 510', the arc segments are sized to provide a ring having an inner radius R1 and an outer radius R2, and a height or depth z9, wherein the dimensions for each quadrant can be selected to achieve a desired magnetic field, non-limiting examples of which are provided below.

[0110] Figure 5F and 5G Examples can be used to form Figure 5H The outer ring 510 shown (e.g., Figure 4 For example, segment 515" may be used to provide a different view of the segment 515 of the ring 410a). Figure 5H The segments in the illustrated quadrants Q1 to Q4 (again, for example, Figure 4 The exemplary portion 515" comprises a plurality of ferromagnetic arc segments (eg, segments formed of NdFeB, SmCo, etc.). Figure 5F to 5H In the example, five arc segments (e.g., exemplary arc segment 505"), each spanning 18° of the ring 510", form a quarter circle of the ring 510". In the exemplary segment 515" of the ring 510", the arc segments are sized to provide a ring having an inner radius R1 and an outer radius R2 and a height or depth z. 10 A ring, wherein the dimensions of each quadrant can be selected to achieve the desired magnetic field.

[0111] As discussed above, the inventors have developed a low-power, portable, low-field MRI system that can be deployed in virtually any environment and brought to the patient undergoing an imaging procedure. This allows patients in emergency rooms, intensive care units, operating rooms, and many other locations to benefit from MRI in situations where MRI is traditionally unavailable. Aspects that facilitate portable MRI are discussed in further detail below.

[0112] Figure 6A and 6B A low-power, portable, low-field MRI system according to some embodiments is illustrated. The portable MRI system 600 includes a B0 magnet 605 comprising at least one first permanent magnet 610a and at least one second permanent magnet 610b magnetically coupled to each other by a ferromagnetic yoke 620 configured to capture and direct magnetic flux to increase the magnetic flux density within an imaging region (field of view) of the MRI system. The permanent magnets 610a and 610b can be formed using any suitable technology, including any of the technologies described herein (e.g., using a combination of Figure 2A The B0 magnet 200 shown and / or Figure 3The yoke 620 may also be constructed using any of the techniques described herein (e.g., using a combination of Figure 2A The magnetic yoke 220 and Figure 3 320 and the accompanying description thereof). It should be understood that in some embodiments, electromagnets can be used to form the BO magnet 605 using any of the electromagnet technologies described herein. The BO magnet 605 can be enclosed or enclosed in a housing 612 along with one or more other magnetic components, such as the system's gradient coils (e.g., x-gradient coils, y-gradient coils, and z-gradient coils) and / or any shim components (e.g., shim coils or permanent magnet shims), BO correction coils, etc.

[0113] The BO magnet 605 may be coupled to or otherwise attached or mounted to the base 650 via a positioning mechanism 690 (such as a goniometer table or the like) such that the BO magnet can be tilted (e.g., rotated about its center of mass) to provide an inclined surface to accommodate the patient's anatomy as needed. Figure 6A In the figure, the B0 magnet is shown as horizontal without any bevel, and Figure 6B , the B0 magnet is shown as tilting the face of the structure supporting the patient being scanned after being rotated. The positioning mechanism 690 may be fixed to one or more load-bearing structures of the base 650, which is arranged to support the weight of the B0 magnet 600.

[0114] In addition to providing a bearing structure for supporting the B0 magnet, the base 650 also includes an internal space configured to accommodate the electronics 670 required to operate the portable MRI system 600. For example, the base 650 can accommodate power components to operate the gradient coils (e.g., X, Y, and Z) and the RF transmit / receive coils. The inventors have developed general low power, low noise, and low cost gradient amplifiers that are configured to appropriately power the gradient coils in the low field region, are designed to be relatively low cost, and are constructed for installation within the base of the portable MRI system (i.e., instead of being statically placed in a separate room in a fixed installation as is conventionally done). Suitable power components for operating the gradient coils (e.g., in combination with the RF transmit / receive coils) are described in further detail below. Figure 1 and Figure 7Examples of the power components described. According to some embodiments, the power electronics used to power the gradient coils of the MRI system consume less than 50 W when the system is idle, and between 100-300 W when the MRI system is in operation (i.e., during image acquisition). The base 650 can also house RF coil amplifiers (i.e., power amplifiers used to operate the transmit / receive coils of the system), power supplies, a control console, a power distribution unit, and other electronics required to operate the MRI system, details of which are further described below.

[0115] According to some embodiments, the electronics 670 required to operate the portable MRI system 600 (e.g., an MRI system using a permanent B0magnet solution) consume less than 1 kW of power, in some embodiments, less than 750 W of power, and in some embodiments, less than 500 W of power. Techniques that facilitate low power operation of MRI devices are discussed in further detail below. However, systems that consume greater power can also be utilized as these aspects are not limited in this regard. Figure 6A and 6B The example portable MRI system 600 shown can be powered via a single power connection 675 that is configured to be connected to a source of mains power, such as a receptacle that provides single phase power (e.g., a standard or large appliance receptacle), etc. Thus, the portable MRI system can be plugged into a single available power receptacle and operate from that power receptacle, thereby eliminating the need for a dedicated power source (e.g., eliminating the need for a dedicated three-phase power source, eliminating the need for further power conversion electronics that convert three-phase power to single-phase power for distribution to respective components of the MRI system), and increasing the availability of the MRI system as well as the situations and locations in which the portable MRI system can be used.

[0116] Figure 6A and 6B The portable MRI system 600 shown also includes a transport mechanism 680 that allows the portable MRI system to be transported to different locations. The transport mechanism can include one or more components configured to facilitate movement of the portable MRI system to a location where MRI is needed, for example. According to some embodiments, the transport mechanism includes a motor 686 coupled to drive wheels 684. In this way, the transport mechanism 680 provides motorized assistance in transporting the MRI system 600 to a desired location. The transport mechanism 680 can also include a plurality of caster wheels 682 to assist with support and stability as well as facilitate transport.

[0117] According to some embodiments, the transport mechanism 680 includes a motorized assist controlled using a controller (e.g., a joystick or other controller manipulable by a person) to guide the portable MRI system during transport to a desired location. According to some embodiments, the transport mechanism includes a power assist component configured to detect when a force is applied to the MRI system and, in response, engage the transport mechanism to provide motorized assistance in the direction of the detected force. For example, Figure 6A and 6B The tracks 655 of the base 650 shown in FIG. 6 can be configured to detect when a force is applied to the tracks (e.g., by a person pushing on the tracks) and engage the transport mechanism to provide motorized assistance to drive the wheels in the direction of the applied force. As a result, a user can guide the portable MRI system with the assistance of the transport mechanism that responds to the direction of the force applied by the user. The power assist mechanism can also provide a safety mechanism against collisions. In particular, contact forces with another object (e.g., a wall, a bed, or other structure) can also be detected, and the transport mechanism will react accordingly with a motorized motion response away from the object. According to some embodiments, the motorized assistance can be disengaged, and the portable MRI system can be transported by having a person move the system to a desired location using manual force.

[0118] The portable MRI system 600 includes a slide 660 that provides electromagnetic shielding to the imaging region of the system. The slide 660 can be transparent or translucent to maintain the open feel of the MRI system to assist patients who can experience claustrophobia during traditional MRI in a closed bore. The slide 660 can also be perforated to allow airflow to increase the open feel and / or dissipate acoustic noise generated by the MRI system during operation. The slide can have a shield 665 incorporated therein to block electromagnetic noise from reaching the imaging region. According to some embodiments, the slide 660 can also be formed of a conductive mesh that provides the shield 665 to the imaging region and facilitates the open feel of the system. Thus, the slide 660 can provide an electromagnetic shield that is movable to allow a patient to be positioned within the system, to permit adjustment by a person once the patient is positioned or during acquisition, and / or to enable a surgeon to gain access to the patient, etc. Thus, the movable shield facilitates flexibility that allows the portable MRI system to be used not only in a non-shielded room, but also to be able to perform procedures that would otherwise not be available. Exemplary slides that provide varying levels of electromagnetic shielding are discussed in further detail below.

[0119] According to some embodiments, the portable MRI system does not include a sliding element, thereby providing a substantially open imaging region, which facilitates easier placement of a patient within the system, thereby reducing a feeling of claustrophobia and / or improving access to a patient positioned within the MRI system (e.g., allowing a physician or surgeon to access the patient prior to, during, or after an imaging procedure without having to remove the patient from the system). The present inventors have developed techniques that facilitate MRI with varying levels of electromagnetic shielding, including no or substantially no shielding of the imaging region, including noise suppression systems adapted to suppress electromagnetic noise in the environment. According to some embodiments, the portable MRI system 600 can be equipped with a noise reduction system that uses one or more of the noise suppression and / or avoidance techniques described herein to dynamically adapt the noise suppression / cancellation response in coordination with the shielding structure of a given shielding arrangement of the portable MRI system 600, for example. Thus, the portable low-field MRI system 600 can be transported to a patient and / or a desired location and operated outside of a specially shielded room (e.g., in an emergency room, operating room, NICU, general practitioner’s office, clinic) and / or taken directly to the bedside of a patient in an arbitrary location, thereby allowing MRI to be performed at a time and place where MRI is needed. As discussed in further detail below, to facilitate a portable MRI that can be operated at almost any location, the present inventors have developed low-power MRI systems that, according to some embodiments, are configured to be powered by mains power (e.g., single-phase power from a standard or industrial wall outlet).

[0120] As discussed above, conventional MRI systems consume a large amount of power, thereby requiring a dedicated three-phase power supply to operate. In particular, conventional MRI systems that use superconducting material to form the B0magnet require a cryogenic cooling system that consumes a large amount of power to maintain the conductor in a superconducting state. Additionally, the power amplifiers used to operate the gradient amplifiers are large power components that absorb a large amount of power and are typically stored in a separate room from the electronic components that house the system. Moreover, the power components configured to operate the transmit / receive coil system of a conventional MRI system also consume a large amount of power. Many conventional high-field MRI systems require an HVAC system that also absorbs a large amount of power.

[0121] Conventional MRI systems are fixed installations that require dedicated and specialized space. As a result, the requirement of a dedicated three-phase power connection to operate an MRI system is not a key limitation of these systems, as this requirement is merely one of many specialized and dedicated features of a conventional MRI facility. However, the requirement of a dedicated three-phase power supply places a significant limitation on the locations where a portable MRI system can operate. Accordingly, the present inventors have developed low-power MRI systems that facilitate the portability of MRI systems. For example, according to some embodiments, a low-power MRI system is configured to operate using mains power (e.g., single-phase power from a standard or industrial outlet). Exemplary aspects of low-power MRI systems are discussed in further detail below.

[0122] According to some embodiments, a low-power MRI system includes a permanent B0 magnet (e.g., any of the permanent magnets discussed herein, such as the permanent magnet shown in FIG. 1). Figure 2A and Figure 3 As a permanent B0 magnet will generate its own persistent magnetic field once magnetized, no power is required to operate the permanent B0 magnet to generate its magnetic field. As a result, as discussed in further detail below in connection with exemplary low-power MRI systems, a significant (often dominant) contributor to the overall power consumption of an MRI system can be eliminated, thereby facilitating the development of MRI systems that can be powered using mains power (e.g., via a standard wall outlet or common large appliance outlet).

[0123] Further, conventional power components adapted to operate gradient coil systems are generally not suitable for use in low-field MRI, at least in part due to expense and noise levels, and are not suitable for low-power and / or portable MRI due to power consumption, size, and weight. For example, while the cost of conventional power components used to operate gradient coils in currently available MRI systems is relatively insignificant in view of the overall cost of a high-field MRI installation, this cost can be unacceptably high in the context of a low-field MRI system designed as a low-cost alternative. Accordingly, the cost of power components conventionally used for high-field MRI can be disproportionately large, and thus unsatisfactory for some lower-cost low-field MRI systems.

[0124] Additionally, the relatively low SNR in low fields (and particularly in the very low and ultra-low field regions) makes traditional gradient coil power assemblies unsuitable. In particular, traditional power assemblies for driving gradient coils are generally unsuitable for low field MRI systems as they are not designed to drive the coils at sufficiently low noise. Although the noise injected by such power assemblies can be acceptable in the high SNR region of high field MRI systems, such assemblies generally do not provide a sufficiently low level of noise to provide acceptable image quality in low field MRI systems. For example, traditional power assemblies can exhibit unsatisfactory variations in output (e.g., ripple) used in low field environments, thereby injecting relatively significant noise into the gradient coil system of a low field MRI system.

[0125] Additionally, traditional power assemblies configured to drive the gradient coil systems of currently available MRI systems are not designed to be power efficient, thereby consuming large amounts of power. Moreover, traditional power assemblies configured to operate the gradient coil systems of currently available MRI systems are large, heavy devices that are typically housed in a separate room adjacent to the MRI device along with other electronic components. Thus, traditional gradient power assemblies are not suitable for use in low power, portable MRI systems.

[0126] Figure 7 An example of a drive circuit for driving current through a coil 705 of an MRI system to produce a magnetic field in accordance with a desired pulse sequence is illustrated in accordance with some embodiments. A power assembly 703 drives current through the coil 705 based on control signals from a controller 701. As discussed above, the controller 701 can generate the control signals to drive the power assembly 703 based on a pulse sequence implemented by the controller 701 (or provided by one or more other controllers). In some embodiments, the coil 705 can be a gradient coil 128. However, the techniques described herein are not limited in this regard, as the coil 705 can be a coil of a magnet 122, a shim coil 124, or an RF transmit and / or receive coil 126. In some embodiments, the controller 701 can correspond to the controller 106 of Figure 1 , and the drive power assembly 703 can correspond to the power assembly 114 of Figure 1 .

[0127] Power components configured to supply power to gradient coils typically provide relatively high power, and often need to provide precise control of the current supplied to the gradient coils so that the desired pulse sequence can be faithfully delivered. Inaccuracies in delivering the commanded current into the gradient coils results in a reduction in signal-to-noise due to differences between the gradient pulse sequence being delivered and the intended (and expected) pulse sequence. The power components configured to drive the gradient coils should also be responsive in delivering the commanded current to the gradient coils, including rapid transitions between commanded current levels, to produce the current waveforms needed to faithfully produce the desired pulse sequence. Accordingly, the inventors have developed power components that can be controlled to operate to faithfully reproduce a desired pulse sequence by accurately and precisely providing current having relatively low noise and relatively high efficiency to one or more gradient coils, some embodiments of which are discussed in further detail below.

[0128] In some embodiments, the power component 703 can be a "current mode" power component that drives a desired current through the coil 705. The desired current can be produced by the power component 703 in response to a current command from the controller 701. In this regard, the power component 703 can operate as a current source controlled by the current command (which can be provided as a voltage level by the controller as an indication of the current to be provided to the coil 705). The controller 701 can vary the current command so that the power component 703 produces current values that vary according to a selected pulse sequence. For example, the controller 701 can command the power component to drive one or more gradient coils according to a pulse sequence that includes a plurality of gradient pulses. For each gradient pulse, the power component can need to ramp up the current provided to the respective gradient coil at the rising edge of the gradient pulse, and ramp down the current provided to the gradient coil at the falling edge of the gradient pulse. Example operations of a power component configured to drive a gradient coil to provide a plurality of such gradient pulses are explained in further detail below.

[0129] Figure 8 is a diagram of an exemplary computer system that can implement embodiments described herein. For example, the computer system 800 can be used to implement the method for measuring magnetic hysteresis effects in an MRI system (see, e.g., the method 1000 of FIG. 10 below) and / or the method for measuring magnetic hysteresis effects in an MRI system (see, e.g., the method 1100 of FIG. 11 below). Figure 16) and methods for controlling gradient coils of an MRI system (e.g., see FIG. 17 below). The computer system 800 can include one or more computer hardware processors 802 and one or more articles of manufacture including a non-transitory computer readable storage medium (e.g., memory 804 and one or more non-volatile storage devices 806). The processor 802 can control writing data to and reading data from the memory 804 and the non-volatile storage 806 in any suitable manner, as aspects of the disclosure provided herein are not limited in this regard. To perform any of the functionality described herein, the processor 802 can execute one or more processor-executable instructions stored in one or more non-transitory computer readable storage media (e.g., the memory 804), which can serve as a non-transitory computer readable storage medium that stores processor-executable instructions for execution by the processor 802.

[0130] Figure 9A An example gradient pulse sequence that can be used by an MRI system to control one or more gradient coils and generate a magnetic field in an imaging region of the MRI system is illustrated. Figure 9A The dashed line in FIG. 1 represents an ideal gradient pulse sequence as a function of time. The pulse sequence causes the gradient coils to produce a magnetic field that oscillates from a maximum value to a minimum value, while the maximum and minimum values of the oscillation decay over time. The constantly changing magnetic field causes an induced magnetization of components of the MRI system. Figure 9A The solid line in FIG. 1 represents the resulting gradient hysteresis (i.e., induced magnetization) as a function of time. The induced magnetization produces its own magnetic field in the imaging region that adds to the magnetic field produced by the gradient coils, resulting in Figure 9B the hysteresis effect shown in FIG. 1, which Figure 9B The hysteresis gradient field (i.e., the gradient field generated by the induced magnetization) is plotted as a function of the gradient field strength generated by the gradient coils during the decaying oscillation pulse sequence. As can be seen from the plot of the hysteresis effect in FIG. 1, Figure 9B As can be seen from the plot of the hysteresis effect in FIG. 1, the magnitude of the hysteresis effect depends on the strength of the magnetic field generated by the gradient coils of the MRI system and whether the magnetic field generated by the gradient coils is constantly increasing or constantly decreasing.

[0131] The Preisach model, which uses a weighting function to characterize the magnetism of a material, can be used to measure and model the hysteresis effects in systems such as MRI systems. Details of the Preisach model and determining the weighting function are described in “Difficult in identification of Preisach hysteresis model weighting function using first order reversal curves method in soft magnetic materials,” Applied Mathematics and Computation Vol. 319, pp. 469-485 (February 15, 2018), http: / / dx.doi.org / 10.1016 / j.amc.2017.05.017, which is incorporated herein by reference in its entirety. In the case of applying the Preisach model to an MRI system, the magnetism of the MRI system is characterized as a whole, and these magnetisms can include the properties of the magnetic yoke, electronics, housing, or any other ferromagnetic material in the vicinity of the MRI system.

[0132] The Preisach model uses an ideal dipole model with a rectangular hysteresis loop, called a “hysteron,” to model the hysteresis effects of an MRI system. Figure 10A An example of a hysteron according to some embodiments is illustrated. The hysteron has three parameters that characterize the hysteresis loop 2300: a lower magnetic field strength value 1001 (H d ), an upper magnetic field strength value 1003 (H u ), and a dipole moment value (m s ) that defines an upper magnetization value 1005 and a lower magnetization value 1007 of the hysteresis loop 1000. When the external magnetic field increases and becomes greater than the upper magnetic field strength value 1003, the dipole magnetic moment switches to the upper magnetization value 1005. Similarly, when the external magnetic field decreases and becomes less than the lower magnetic field strength value 1001, the dipole magnetic moment switches to the lower magnetization value 1007. When the external magnetic field strength is between the lower magnetic field strength value 1001 and the upper magnetic field strength value 1003, the state of the dipole magnetic moment depends on the previous state.

[0133] The Preisach model uses a plurality of hysternons arranged in a two-dimensional array of hysternons, where each hysteron in a particular row of the array has the same upper magnetic field strength value 1003, and each hysteron in a particular column of the array has the same lower magnetic field strength value 1001. Figure 10BSchematic diagram illustrating a simplified Preisach model 1050 including five rows and five columns. It should be understood that the Preisach model used in some embodiments may have any number of rows and columns. For example, some embodiments may include tens, hundreds, or thousands of rows and columns. Figure 10B As shown, the hysteresis operators in the first column all have the same lower magnetic field strength value 1001. The hysteresis operators in the second column also have lower magnetic field strength values ​​1001 that are the same as each other but higher than the lower magnetic field strength value 1001 of the first column. Similarly, when moving rightward to higher-numbered columns in the Preisach model 1050, the lower magnetic field strength value 1001 increases. In a similar manner, the hysteresis operators in the first row all have the same upper magnetic field strength value 1003. The hysteresis operators in the second row also have upper magnetic field strength values ​​1003 that are the same as each other but lower than the upper magnetic field strength value 1003 of row 1. Similarly, when moving downward to higher-numbered rows in the Preisach model 1050, the upper magnetic field strength value 1003 decreases.

[0134] Preisach model 1050 is a two-dimensional array in the shape of a triangle. No hysteresis operator is present in the lower right portion of the model because such a hysteresis operator would correspond to a hysteresis loop that is physically unrealizable because the upper magnetic field strength value 1003 must be greater than the lower magnetic field strength value 1001. The diagonal hysteresis operator along the bottom of the triangle corresponds to a hysteresis operator where the lower magnetic field strength value 1001 is equal to the upper magnetic field strength value 1003.

[0135] Using the Preisach model 1050, the magnetization of an MRI system can be determined using the following equation:

[0136]

[0137] where H(t) is the externally applied magnetic field strength from, for example, a gradient coil, and is the hysteresis operator acting on the external magnetic field H(t), which is calculated according to H u and H d The value of +m is obtained by the value of and the direction of change of H(t) (for example, whether the external magnetic field is increasing or decreasing). s or -m s According to some embodiments, the weighting function w(H u ,H d) specifies the value of the fundamental magnetic momentum of the MRI system and is determined by measuring the magnetic hysteresis of the MRI system. Thus, the Preisach model of the MRI system has a state at any given time that is based on the current value of the externally applied magnetic field and the history of the externally applied magnetic field (e.g., as represented by the previous state of the Preisach model). Thus, when the value of the externally applied magnetic field changes, the state of the Preisach model also changes.

[0138] As Figure 11A and 11C illustrated, the dynamics of the Preisach model can be illustrated using an external magnetic field that is a function of time, Figure 11A and 11C illustrates a magnetic field that changes linearly but oscillates between a high value and a low value, where the high value and the low value decrease in magnitude over time. Figure 11A The asterisk (*) in illustrates the current value of the externally applied magnetic field after being turned on some time ago and is currently increasing in magnitude. Figure 11B illustrates the state of the Preisach model at this particular time. The hysteresis operators that are polarized up are illustrated as solid circles, while the hollow circles represent hysteresis operators that have a downward polarization. The instantaneous external field strength is illustrated by the horizontal line 1101. As time progresses, the horizontal line 1101 moves from the bottom to the top of the array of hysteresis operators. The hysteresis operators below the current level of the horizontal line 1101 are in an upwardly polarized configuration. The hysteresis operators above the current level of the horizontal line 1101 remain unchanged relative to the previous state of the Preisach model. Figure 11B

[0139] The asterisk in illustrates the current value of the externally applied magnetic field after being turned on some time ago and is currently decreasing in magnitude. Figure 11C illustrates the state of the Preisach model at this particular time. As in Figure 11D Figure 11B illustrates the state of the Preisach model at this particular time. As in Figure 11D

[0140] The above explanation of the Preisach model explains the state of the model and how that state changes when an externally applied magnetic field (e.g., from a gradient coil) is increased and / or decreased. Figure 12 ​​Another illustration of the Preisach model 1200 is illustrated, where H d -H u The discrete locations in the plane are each represented by a respective box 1201. Each box is associated with a weight used to adjust the target pulse sequence to obtain a corrected pulse sequence used to control the gradient coils of the MRI system. In some embodiments, the amplitude of a pulse in the gradient pulse sequence is altered based on the state of a Preisach model. For example, if a particular hysteresis operator is determined to be in a downward polarization, then no adjustment to the amplitude of a pulse in the gradient pulse sequence can be made. However, if the same hysteresis operator is determined to be in an upward polarization state, then a weight can be added to the amplitude of the pulse, whereby the amplitude of the pulse is changed based on the particular hysteresis operator only when the state of the hysteresis operator is in one of the two states. Alternatively, the weight can be determined such that when a particular hysteresis operator is determined to be in a downward polarization, the weight is subtracted from the amplitude of a pulse in the gradient coil sequence, and if the same hysteresis operator is determined to be in an upward polarization state, then a weight can be added to the amplitude of the pulse, whereby the amplitude of the pulse is changed based on the particular hysteresis operator but in the opposite direction for the two states of the hysteresis operator.

[0141] The weights associated with the boxes 1201 of the Preisach model 1200 can be determined based on a measurement of the hysteresis effect performed in a calibration phase prior to imaging a patient. To measure the induced magnetization of the MRI system, a multi-element probe is placed within the imaging region of the MRI system, with each receiving element of the multi-element probe located at a different location within the imaging region such that the induced gradient field can be measured. The various electromagnets of the MRI system are then controlled by a controller 701 using a pulse sequence comprising a plurality of pulses to generate a dynamic magnetic field measured by the field probe to determine the hysteresis effect of the applied magnetic field. Figure 7

[0142] Figure 13A A multi-element field probe according to some embodiments is illustrated. The multi-element field probe 1300 includes a plurality of RF receiving elements 1301a-1301h, a housing 1302, an RF transmit coil 1303, a plurality of electrical connectors 1305, control electronics 1307, a base support 1309, fasteners 1311, and a plurality of liquid samples 1321. Although the multi-element field probe 1300 is illustrated with a particular number of components and arrangement of components, it should be understood that the illustrated probe is merely an example and that other arrangements of components can be used. The diameter of the housing 1302 can determine the size of the multi-element field probe 1300, which can be approximately 12 cm. This size is chosen to be large enough to have a large SNR, but small enough to reduce the inhomogeneity of the magnetic field. ​

[0143] The multi-element field detector 1300 is positioned within the MRI system so that it experiences the various magnetic fields generated by the BO magnet, gradient coils, and shim coils. The RF transmit coil 1303 is controlled to generate RF pulses that cause precession of molecules in the liquid sample 1321. The receive element 1301 measures the MRI signal obtained from each corresponding liquid sample 1321 and sends the signal to the control electronics 1307.

[0144] Figure 13B Examples of examples according to some embodiments Figure 13A One of the RF receiving elements 1301. In some embodiments, the RF receiving elements 1301 can each include a coil wrapped around a corresponding liquid sample 1321. The coil 1323 can be a conductive wiring formed by any suitable conductor such as a metal (e.g., copper). In some embodiments, the coil 1323 can include a multi-layer Litz solenoid so that the liquid sample 1321 is inside the solenoid. For example, a double-layer Litz solenoid can be used. The RF receiving elements 1301 can also each include a twisted pair Litz cable 1325 that electrically connects the RF receiving element 1301 to the control electronics 1307. The liquid sample 1321 and the coil 1323 can be further embedded in an epoxy resin housing 1322.

[0145] In some embodiments, RF receiving elements 1301 are positioned within housing 1302. RF receiving elements 1301 may be evenly spaced apart. For example, the positions of the RF receiving elements 1301 may correspond to the corners of an imaginary cube. In some embodiments, RF receiving elements 1301 may be positioned at the intersection of a corner of the imaginary cube and an imaginary sphere inscribed within the cube. The imaginary sphere may have a diameter of, for example, approximately 12 cm.

[0146] In some embodiments, liquid sample 1321 may include water and / or mineral oil. Liquid sample 1321 may also include agar to hold the liquid in place. Alternatively or additionally, liquid sample 1321 may include copper sulfate to reduce the relaxation time of the liquid sample. In some embodiments, a support member 1304 made of a non-ferromagnetic material may be used to hold liquid sample 1321 and the RF receiving element in place. Although liquid sample 1321 is described as a liquid throughout this application, it should be understood that in some embodiments, liquid sample 1321 may be a gel sample.

[0147] In some embodiments, the RF transmit coil 1303 includes electrically conductive wiring wrapped around the exterior or interior of the housing 1302. The RF transmit coil 1303 can include a plurality of circular loops around the RF receive elements 1301. In the field probe 1300 shown in FIG. 13, the RF transmit coil 1303 includes six circular loops. It will be appreciated that the RF transmit coil 1303 can include more loops or fewer loops.

[0148] In some embodiments, the housing 1302 is supported by a base 1309 that is configured to rest on a surface within an imaging region of an MRI system. The base 1309 can include fasteners 1311 that are configured to position the field probe 1300 at the isocenter of the MRI system. For example, the fasteners 1311 can be a particular shape cut into the base 1309 that fits into a correspondingly shaped fastener located within the MRI system. When the fasteners 1311 engage with the fasteners in the MRI system, an operator of the MRI system can be confident that the field probe 1300 is located at the isocenter of the MRI system.

[0149] In Figure 13C to 13G additional figures of the example field probe 1300 are illustrated in FIGS. 14-17. The support members 1304 are more clearly shown in these figures. In some embodiments, the support members 1304 include openings for housing the plurality of RF receive elements 1301.

[0150] Once the field probe 1300 is in place in the MRI system, it can be used to measure the magnetic field in the imaging region of the MRI system. The measurements of the magnetic field in the MRI system can then be used to determine Figure 12 the weights of the Preisach model 1200. Figure 14 is a flowchart of a method 1400 for measuring hysteresis in an MRI system including at least one gradient coil according to some embodiments.

[0151] At act 1402, the method 1400 includes controlling the at least one gradient coil using a first pulse sequence. As discussed above, the controller 106 can use a pulse sequence to control the magnetic fields produced by the gradient coil(s) 128. The first pulse sequence can include a plurality of pulse sequences related to different electromagnets within the MRI system 100. For example, the first pulse sequence can include one or more pulses for controlling the B0 magnet 122, the shim coils 124, the RF transmit and receive coils 126, and the gradient coils 128. The first pulse sequence can be a pulse sequence specifically designed for measuring the effects of hysteresis in an MRI system.

[0152] Figure 15AA portion 1500 of an example first pulse sequence for measuring the hysteresis effect of an MRI system according to some embodiments is illustrated. The illustrated portion 1500 is a pulse sequence for a single repetition time (TR), which is the time elapsed between two different RF transmit pulses. The portion 1500 of the pulse sequence includes an RF transmit pulse 1501, a read window 1503, and a gradient squeeze pulse 1505. The read window 1503 is between the RF pulse 1501 and the gradient squeeze pulse 1505. However, embodiments may include additional pulses as part of the first pulse sequence. For example, the portion 1500 may be repeated multiple times. In some embodiments, the gradient squeeze pulse 1505 may vary in amplitude after each iteration. A controller may access one or more parameters specifying the first pulse sequence in a memory to control the electromagnet. These parameters may include pulse timing, pulse center frequency, pulse phase, and / or pulse amplitude.

[0153] Figure 15B An exemplary gradient squeeze amplitude sequence 1550 is illustrated in accordance with some embodiments. The gradient squeeze amplitude sequence provides the amplitude of the gradient squeeze pulse 1505 for each subsequent iteration of the portion 1500 of the pulse sequence. Note that although Figure 15B Only the amplitude of the gradient squeeze pulses is shown, but each gradient squeeze pulse is associated with an RF transmit pulse 1501 and a read window 1503 .

[0154] In some embodiments, where an x-, y-, and z-gradient coil are used, there are three separate parts of the gradient squeeze amplitude sequence 1550. For example, first only the x-gradient squeeze pulse is used, followed by the y-gradient squeeze pulse, and finally the z-gradient squeeze pulse. It will be appreciated that in some embodiments, the order may be different, but the squeeze pulse is only used to control one gradient coil at a time until the entire gradient pulse sequence for each gradient coil is complete, and then the pulse sequence continues with the next gradient coil. Figure 15B In the example shown, first the x-gradient coil is controlled by a pulse sequence 1511 with a continuously changing amplitude, thereby oscillating between the maximum and minimum amplitudes of the attenuation until the amplitude is negligible. After the x-gradient pulse sequence is completed, the y-gradient pulse sequence starts with a pulse sequence 1513 with a continuously changing amplitude, thereby oscillating between the maximum and minimum amplitudes of the attenuation until the amplitude is negligible. Finally, after the y-gradient pulse sequence is completed, the z-gradient pulse sequence 1515 starts with a pulse sequence with a continuously changing amplitude, thereby oscillating between the maximum and minimum amplitudes of the attenuation until the amplitude is negligible. Thus, as Figure 15B As shown, in each gradient pulse sequence, there is a portion where the amplitude of the gradient squeeze pulse increases monotonically and a portion where the amplitude of the gradient squeeze pulse decreases monotonically.

[0155] Returning to Figure 14 , the method 1400 continues with act 1604, which includes measuring a first plurality of magnetic field strengths in an imaging region of the MRI system using a multi-element RF probe. In some embodiments, act 1604 can be performed concurrently with act 1604. For example, each RF transmit pulse 1501 is associated with a readout window 1503 during which the measurement of the magnetic field strength occurs. Thus, the measurement occurs during the use of the pulse sequence to control the gradient coils. However, in some embodiments, neither the RF transmit coil nor the gradient coils are "on" when the measurement of the magnetic field strength is made. In this way, the induced magnetic field is measured in the absence of the magnetic fields generated by the gradient coils.

[0156] After measuring the magnetic field strengths, the computer hardware processor is used to extract the phase from each element in the field probe and separate the phase into linear spatial terms (e.g., x, y, z gradient terms) and a constant field term (e.g., a B0term). Figure 16 The hysteresis effect of each individual portion of the extracted phase for each axis of the MRI system is shown. The plots in the top row correspond to the hysteresis effect measured when the x gradient coil is controlled by the gradient pulse sequence described above; the plots in the middle row correspond to the hysteresis effect measured when the y gradient coil is controlled by the gradient pulse sequence described above; and the plots in the bottom row correspond to the hysteresis effect measured when the z gradient coil is controlled by the gradient pulse sequence described above. The plots in the first column represent the constant B0term for each axis; the plots in the second column represent the linear term along the x axis; the plots in the third column represent the linear term along the y axis; and the plots in the fourth column represent the linear term along the z axis. As can be seen, for the linear terms, significant hysteresis effects exist only along the axis corresponding to the gradient coil used. For example, in the top row corresponding to the hysteresis produced by driving the x gradient coil, significant hysteresis in the linear term is found only along the x axis. Similarly, in the middle row corresponding to the hysteresis produced by driving the y gradient coil, significant hysteresis in the linear term is found only along the y axis. Similarly, in the bottom row corresponding to the hysteresis produced by driving the z gradient coil, significant hysteresis in the linear term is found only along the z axis.

[0157] From Figure 16As can be seen from the data shown, hysteresis effects in the linear terms only need to be corrected for the directions associated with the gradient coils. In other words, when adjusting the amplitude of the target gradient pulse sequence for a particular direction to determine the corrective gradient pulse sequence for that particular direction, the controller considers the target gradient pulse sequence for that particular direction, and the corrective gradient pulse sequence is not based on the target gradient pulse sequences for other directions. Therefore, in some embodiments, a separate independent hysteresis model (e.g., a Preisach model) exists for each of the three directions.

[0158] from Figure 16 As can be seen in the graph of , the constant B0 term exhibits hysteresis only along the y-axis. This is because, in the example MRI system used for this data, the B0 field is oriented along the y-axis. Therefore, since the y-gradient coil induces magnetization along the y-axis, this induced magnetization affects not only the linear y term but also the constant B0 term. The x- and z-gradient coils do not induce hysteresis effects in the B0 term. However, it should be understood that other MRI systems may have B0 fields oriented in different directions, in which case other directions may experience hysteresis in the constant B0 term.

[0159] At act 1406 , method 1400 includes estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths. Act 1406 is performed after act 1402 and act 1404 . Figure 17A An example of weights for the Preisach model is shown, which is obtained by iteratively fitting the measured power measurements of the following hysteresis: Figure 17B The hysteresis of the gradient squeeze pulse amplitude as a function of Figure 17C The hysteresis of the Preisach model is a function of time. The weights of the Preisach model are iteratively determined by initializing the weights to initial values. As the gradient squeezing pulse amplitude increases, the columns of the Preisach model are set, for example, by taking a simple difference between the measured magnetic field strength and the initialized weights. Then, as the gradient squeezing pulse amplitude decreases, the rows of the Preisach model are set in a similar manner, for example, by taking a simple difference between the measured magnetic field strength and the initialized weights. After the initial cycle through the Preisach model, the weights have now been updated to be different from the initialization values. The same process can be performed again for a second cycle to further update the weights. This process can be iterated until the weights of the Preisach model stop changing by a threshold amount. Optionally, the algorithm for estimating the weights can run a set number of cycles.

[0160] As mentioned above, there may be a Preisach model for each of the x, y, and z directions. Therefore, the same iterative fitting process can be performed for each of these three directions.

[0161] At act 1408, the method 1400 includes storing the parameters of the hysteresis model in a memory. The parameters can include the weights of the Preisach model associated with the x-direction, the weights of the Preisach model associated with the y-direction, and the weights of the Preisach model associated with the z-direction. In some embodiments, the lower and upper magnetic field strength values associated with each weight (see, e.g., the hysteresis operator of Figure 10A-10B ) can be stored in the memory along with the weights.

[0162] After measuring the hysteresis effects and storing the parameters of the hysteresis model in a memory, the hysteresis model can be used to determine a correction pulse sequence from a target pulse sequence to reduce errors produced by induced magnetization in the MRI system. Figure 18 is a flowchart of a method 1800 for controlling at least one gradient coil of an MRI system.

[0163] At act 1802, the method 1800 includes receiving at least one target pulse sequence. In some embodiments, the controller 106 can receive the target pulse sequence from the pulse sequence repository 108 or any other suitable storage device. The target pulse sequence can include a plurality of pulses for controlling various electromagnets of the MRI system. For example, the target pulse sequence can include pulses for controlling the RF transmit and receive coils 126 and / or the gradient coils 128. Thus, the target pulse sequence can include a plurality of sub-sequences such as an x-gradient pulse sequence, which is a pulse sequence for controlling only the x-gradient coils; a y-gradient pulse sequence, which is a pulse sequence for controlling only the y-gradient coils; a z-gradient pulse sequence, which is a pulse sequence for controlling only the z-gradient coils; an RF transmit pulse sequence, which is a pulse sequence for controlling the RF transmit coils; and an RF receive pulse sequence, which is a pulse sequence for controlling the RF receive coils.

[0164] At act 1804, the method 1800 includes determining a correction pulse sequence to control the at least one gradient coil based on the at least one target pulse sequence and a hysteresis model of induced magnetization in the MRI system caused by operation of the at least one gradient coil. In some embodiments, the hysteresis model has a “state” that is based on a history of gradient pulse sequences, and the correction pulse sequence is based on the state. For example, as described above, the hysteresis model can include one or more Preisach models. In some embodiments, the hysteresis model includes a plurality of different parameters, at least a subset of which are determined from previously obtained hysteresis measurements obtained with a multi-element probe.

[0165] In some embodiments, determining the correction pulse sequence is at the beginning of the pulse sequence and iterated through the plurality of pulses. For example, the gradient pulse sequence can include a plurality of target gradient pulses. The amplitude of each of the target gradient pulses can be adjusted to determine the correction gradient pulse sequence. The adjustment to any given gradient pulse can be based at least on the amplitude of the previous gradient pulses within the gradient pulse sequence.

[0166] After iterating through all of the pulses of the pulse sequence, the process can be iterated at least one more time for the entire pulse sequence. The results of the previous iteration are used as input for each subsequent iteration. For example, the algorithm can be executed a first time using the target pulse sequence as input. The results of the first iteration is a first correction pulse sequence. The first correction pulse sequence can then be used as a starting point for a second iteration in which the first correction pulse sequence and the hysteresis model are used to determine a second correction pulse sequence. Embodiments can be iterative because the correction to the amplitude of a particular pulse in the pulse sequence in the first iteration can affect the amplitude of a previous pulse in the pulse sequence in a way that could not be adjusted in the first iteration. Thus, at least a second iteration can help determine a more accurate correction pulse sequence. Any number of iterations can be used. However, because the correction due to hysteresis is relatively small, some embodiments include only two iterations.

[0167] In some embodiments, determining the correction pulse sequence can include adding one or more weights from the Preisach model to the amplitude of one or more gradient pulses of the gradient pulse sequence. As discussed above, which weights are added to the amplitude of each pulse of the target gradient pulse sequence depends on the state of the Preisach model.

[0168] In some embodiments, determining the correction pulse sequence can include adjusting the center frequency or phase of the RF transmit pulses and / or the center frequency or phase of the RF receive pulses. This is because, as discussed above, the hysteresis caused by certain gradient coils can affect the strength of the B0 field. When the strength of the B0 field changes, the precession frequency of the atoms in the patient changes. Thus, the center frequency or phase of the RF transmit pulses and / or the center frequency and / or phase of the RF receive pulses are adjusted to match the altered precession frequency due to the change in the B0 field.

[0169] At act 1806, the method 1800 includes using the correction pulse sequence to control at least one gradient coil to generate one or more gradient pulses for imaging a patient.

[0170] Figure 19A to 19B Example of the improvement in SNR resulting from correcting for the hysteresis effect as described in the present invention. Figure 19Ais an MRI image of the patient generated from imaging without correction of the magnetic hysteresis effect using a correction pulse sequence. Figure 19B is an MRI image of the same patient generated from imaging with correction of the magnetic hysteresis effect using a correction pulse sequence. An increase in SNR of 15% is observed, resulting in a better MRI image that can be used for clinical diagnosis.

[0171] Thus, having described several aspects and embodiments of the technology set forth herein, it is to be appreciated various alterations, modifications, and improvements will readily occur to those skilled in the art. Such alterations, modifications, and improvements are intended to be within the spirit and scope of the technology described herein. For example, it will be appreciated that various other components and / or structural arrangements can be used for performing the functions and / or obtaining the results and / or one or more of the advantages described herein, and that such variations and / or modifications are to be considered within the scope of the embodiments described herein. Those skilled in the art will recognize, or be able to ascertain with no more than a practical pertussis of experimentation, many equivalents to the specific embodiments described herein. It is, therefore, to be understood that the foregoing embodiments are presented by way of example only and that within the scope of the appended claims and equivalents thereto, embodiments can be practiced otherwise than as specifically described and claimed. In addition, any combination of two or more features, systems, articles, materials, kits, and / or methods described herein, if such features, systems, articles, materials, kits, and / or methods are not mutually inconsistent, is included within the scope of the present application.

[0172] The above-described embodiments can be implemented in any of numerous ways. One or more aspects and embodiments of the application involving the performance of processes or methods can be implemented using either software working in combination with hardware, software working alone, hardware working alone, or combinations thereof. In this context, a "process" is generally considered a self-consistent application that performs predetermined tasks and / or generates predetermined results. A "method" is generally considered a procedure or sequence of steps that are performed to achieve a particular result. In this context, the term "process" can be used interchangeably with the term "method". In this context, various inventive concepts can be embodied as a computer-readable storage medium (or multiple computer-readable storage media) encoded with one or more programs (i.e., one or more sets of instructions) that, when executed on one or more computers or other processors, perform a method (or implement various aspects thereof) as described above. The computer-readable medium or media can be transportable, such that the one or more programs stored thereon can be loaded onto one or more different computers or other processors to implement various aspects as described above. In some embodiments, the computer-readable medium can be non-transitory (e.g., tangible physical quantities). In some embodiments, the computer-readable medium can be transitory (e.g., signals).

[0173] The terms "program" or "software" are used herein in a generic sense to refer to any type of computer code or set of computer-executable instructions that can be employed to program a computer or other processor to implement various aspects as described above. Additionally, it should be appreciated that according to one aspect, one or more computer programs that when executed perform some or all of the methods of the present application need not reside on a single computer or processor, but can be distributed in a modular fashion amongst a number of different computers or processors to implement various aspects of the present application.

[0174] Computer-executable instructions can be in many forms, such as program modules, executed by one or more computers or other machines. Generally, program modules include routines, programs, objects, components, data structures, etc. that perform particular tasks or implement particular abstract data types. Typically the functionality of the program modules can be combined or distributed as desired in various embodiments.

[0175] Also, data structures can be stored in computer-readable media in any suitable form. For simplicity of illustration, data structures can be shown to have fields that are related through location in the data structure. Such relationships can likewise be achieved by assigning storage for the fields with locations in a computer-readable medium allowing for the

[0176] The above-described embodiments of the application can be implemented in any of various ways. For example, the embodiments can be implemented using hardware, software or a combination thereof. When implemented in software, the software code can be executed on any suitable processor or collection of processors, whether provided in a single computer or distributed among multiple computers. It should be appreciated that any component or collection of components that perform the functions described above can be generically considered as controllers that control the above-discussed functions. The controllers can be implemented in numerous ways, such as with dedicated hardware, or with general purpose hardware programmed to perform the above-discussed functions using microcode or software, and in combination with software or firmware stored on machine-readable media such as computer-readable media, whether provided local to the controllers or distributed among multiple computers.

[0177] Also, it should be appreciated that a computer can be embodied in any of a number of forms, such as a rack-mounted computer, a desktop computer, a laptop computer, or a tablet computer. Additionally, a computer can be embedded in a device not generally regarded as a computer but which has some processing capacity, such as a PDA, a smart phone or any other suitable portable or fixed electronic device.

[0178] Also, a computer can have one or more input and output devices. These devices can be used, among other things, to present a user interface. Examples of output devices that can be used to provide a user interface include printers or display screens for visual presentation of output and speakers or other sound generating devices for audible presentation of output. Examples of input devices that can be used for a user interface include keyboards, and pointing devices, such as mice, touch pads, and digitizing tablets. As another example, a computer can receive input information through speech recognition or in other audible format.

[0179] Such computers can be interconnected by one or more networks in any suitable form, including a local area network or a wide area network, such as an enterprise network, and intelligent network (IN) or the Internet. Such networks can be based on any suitable technology and can operate according to any suitable protocol and can include wireless networks, wired networks or fiber optic networks.

[0180] Also, some embodiments can be implemented as one or more methods, as described. The acts performed as part of these methods can be ordered in any suitable way. Accordingly, embodiments can be constructed in which acts are performed in an order different than illustrated, which can include performing some acts simultaneously, even though shown as being performed sequentially in illustrative embodiments.

[0181] All definitions, as defined and used herein, should be understood to control over dictionary definitions, definitions in documents incorporated by reference, and / or ordinary meanings of the defined terms.

[0182] In the description and claims, the indefinite articles "a" and "an" should be understood to mean "at least one" unless explicitly stated otherwise.

[0183] In the description and claims, the indefinite articles "a" and "an" should be understood to mean "at least one" unless explicitly stated otherwise.

[0184] In the specification and claims, the phrase “at least one” is understood to mean one or more, i.e., at least one, of the elements listed, but does not necessarily include every one of the elements specifically listed. The definition also allows for the optional presence of additional elements not specifically listed, whether related or unrelated to the elements specifically listed. Thus, as a non-limiting example, “at least one of A and B” (or, equivalently, “at least one of A or B”, or, equivalently, “at least one of A and / or B”) can refer, in one embodiment, to the optional presence of at least one A without presence of B (and, optionally, presence of elements other than B); in another embodiment, to the optional presence of at least one B without presence of A (and, optionally, presence of elements other than A); in yet another embodiment, to the optional presence of at least one A and at least one B (and, optionally, presence of elements other than A or B); and so forth.

[0185] In addition, the phraseology and terminology used herein is for the purpose of description and should not be regarded as limiting. The use of “including,” “comprising,” “having,” “containing”, “involving”, “characterized by”, “characterized into” and variations thereof herein, is meant to encompass the item listed thereafter and equivalents thereof as well as additional items.

[0186] In the claims, as well as in the specification above, all transitional phrases such as “comprising”, “including”, “carrying”, “having”, “containing”, “involving”, “holding”, and “comprised of”, and the like are to be understood to be open-ended, i.e., to mean including but not limited to. Only the transitional phrases “consisting of” and “consisting essentially of” shall be closed or semi-closed transitional phrases, respectively.

Claims

1. A method for measuring magnetic hysteresis in a magnetic resonance imaging system (MRI system), the MRI system comprising at least one gradient coil, the method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring, using a multi-element RF detector positioned in an imaging region of the MRI system, a first plurality of magnetic field intensities in the imaging region of the MRI system, each magnetic field intensity in the first plurality of magnetic field intensities being caused at least in part by a corresponding pulse in the first plurality of pulses of the first pulse sequence; estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths; as well as Parameters of the hysteresis model are stored.

2. The method according to claim 1, further comprising: At least one parameter value specifying the first pulse sequence is accessed.

3. The method according to claim 1, wherein Measuring the first plurality of magnetic field strengths is performed after controlling the at least one gradient coil using the first pulse sequence such that the at least one gradient coil is not generating a magnetic field during measuring the first plurality of magnetic field strengths.

4. The method according to claim 1, wherein The amplitude of at least a portion of the first plurality of pulses decreases over time.

5. The method according to claim 4, wherein The amplitude of at least a portion of the first plurality of pulses increases over time.

6. The method according to claim 1, wherein Storing the parameters of the hysteresis model includes: storing multiple weights based on the first multiple magnetic field strengths, multiple lower magnetic field strength values ​​and multiple upper magnetic field strength values, wherein each weight in the multiple weights is associated with one of the multiple lower magnetic field strength values ​​and one of the multiple upper magnetic field strength values.

7. The method according to claim 6, wherein: Estimating the weights of the hysteresis model includes estimating each of the plurality of weights based on a difference between a first magnetic field strength and a target magnetic field strength.

8. The method according to claim 7, further comprising: After measuring the first plurality of magnetic field strengths, controlling the at least one gradient coil using a second pulse sequence comprising a second plurality of pulses; measuring, using the multi-element RF detector, a second plurality of magnetic field intensities in the imaging region, each magnetic field intensity in the second plurality of magnetic field intensities being caused at least in part by a corresponding pulse in the second plurality of pulses of the second pulse sequence; as well as The plurality of weights are updated based on the second magnetic field strength.

9. The method according to claim 6, further comprising: iteratively controlling the at least one gradient coil using subsequent pulse sequences comprising a plurality of pulses, wherein a first iteration comprises controlling the at least one gradient coil using the first pulse sequence; iteratively measuring, using the multi-element RF detector, a subsequent plurality of magnetic field strengths, each magnetic field strength in the subsequent plurality of magnetic field strengths resulting from a corresponding pulse in the plurality of pulses of the subsequent pulse sequence; as well as The plurality of weights are iteratively updated based on the subsequent plurality of magnetic field strengths.

10. The method according to claim 1, further comprising: An RF coil of the MRI system is controlled using a radio frequency drive pulse (RF drive pulse) preceding each of the first plurality of pulses of the first pulse sequence.

11. The method according to claim 10, further comprising: Each RF drive pulse is followed by a readout window and a corresponding pulse in the first plurality of drive pulses of the first pulse sequence.

12. The method according to claim 1, wherein The at least one gradient coil includes an x-gradient coil, a y-gradient coil, and a z-gradient coil.

13. The method according to claim 12, wherein: The first pulse sequence includes a subsequence of x-gradient drive pulses, a subsequence of y-gradient drive pulses, and a subsequence of z-gradient drive pulses.

14. The method according to claim 1, further comprising: The multi-element RF detector is placed at the isocenter of the MRI system.

15. The method according to claim 1, wherein The multi-element RF detector includes a plurality of RF receiving elements.

16. The method according to claim 15, wherein Each of the plurality of RF receiving elements includes a coil.

17. The method according to claim 16, wherein The multi-element RF probe includes a plurality of liquid samples, wherein each liquid sample of the plurality of liquid samples is contained within a corresponding coil of the plurality of RF receiving elements.

18. The method according to claim 17, wherein At least one of the liquid samples includes agar.

19. At least one computer-readable storage medium storing processor-executable instructions that, when executed by at least one computer hardware processor, cause the at least one computer hardware processor to perform a method for measuring hysteresis in a magnetic resonance imaging (MRI) system, the MRI system including at least one gradient coil, the method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring, using a multi-element RF detector positioned in an imaging region of the MRI system, a first plurality of magnetic field intensities in the imaging region of the MRI system, each magnetic field intensity in the first plurality of magnetic field intensities being caused at least in part by a corresponding pulse in the first plurality of pulses of the first pulse sequence; estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths; as well as Parameters of the hysteresis model are stored.

20. An apparatus for controlling at least one gradient coil of a magnetic resonance imaging system (MRI system), the apparatus comprising: at least one computer hardware processor; as well as At least one computer-readable storage medium storing processor-executable instructions that, when executed by the at least one computer hardware processor, cause the at least one computer hardware processor to perform a method comprising: controlling the at least one gradient coil using a first pulse sequence comprising a first plurality of pulses; measuring, using a multi-element RF detector positioned in an imaging region of the MRI system, a first plurality of magnetic field intensities in the imaging region of the MRI system, each magnetic field intensity in the first plurality of magnetic field intensities being caused at least in part by a corresponding pulse in the first plurality of pulses of the first pulse sequence; estimating parameters of a hysteresis model based on the measured first plurality of magnetic field strengths; and Parameters of the hysteresis model are stored.

21. A computer program product comprising a computer program which, when executed by a processor, implements the method according to any one of claims 1 to 18.

Images