HK1248505B - Radio frequency coil methods and apparatus - Google Patents

Description

Radio frequency coil method and apparatus

Background Art

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. In general, MRI is based on detecting magnetic resonance (MR) signals, which are electromagnetic waves emitted by atoms in response to a change of state caused by an applied electromagnetic field. For example, nuclear magnetic resonance (NMR) technology involves detecting MR signals emitted from the nuclei of excited atoms upon rearrangement or relaxation of the nuclear spins of atoms in the object being imaged (e.g., atoms in human tissue). The detected MR signals can be processed to produce images, which, in the context of medical applications, allows investigation of internal structures and/or biological processes within the body for diagnostic, therapeutic and/or research purposes.

MRI offers an attractive imaging modality for biological imaging due to its ability 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 such as x-rays or introduce radioactive materials into the body). Additionally, MRI is particularly well suited to providing soft tissue contrast, which can be used to image subjects that cannot be satisfactorily imaged by other imaging modalities. Furthermore, MR technology is able to capture information about structures and/or biological processes that cannot be acquired by other modalities. However, there are a number of disadvantages to MRI, which may involve a relatively high cost of the equipment, limited availability and/or difficulty in obtaining access to a clinical MRI scanner and/or the length of the image acquisition process for a given imaging application.

The trend in clinical MRI is 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 increase costs. The vast majority of installed MRI scanners operate at 1.5 or 3 Tesla (T), which refers to the field strength of the main magnetic field, _B0 . A rough cost estimate for a clinical MRI scanner is approximately $1 million per Tesla, which does not take into account the substantial operating costs, service costs, and maintenance costs involved in operating such an MRI scanner.

In addition, conventional high field MRI systems usually require large superconducting magnets and associated electronic devices to generate the strong uniform static magnetic field ( _B0 ) for imaging of an object (e.g., a patient). The size of such a system is quite large, wherein typical MRI equipment includes multiple rooms for magnets, electronic devices, thermal management systems, and console areas. The size and expense of MRI systems usually limit their application to institutions such as hospitals and academic research centers with enough space and resources to purchase and maintain them. The high cost and large space requirements of high field MRI systems cause limited availability of MRI scanners. Therefore, as discussed in further detail below, there is often a clinical situation in which MRI scanning would be useful, but due to one or more of the above-mentioned limitations it is impractical or infeasible.

Summary of the Invention

The inventors have developed a radio frequency component that, in some embodiments, is configured to improve magnetic resonance signal detection, for example to facilitate image acquisition at low field strengths. Some embodiments include a radio frequency coil configured to respond to magnetic resonance signals, the radio frequency coil comprising at least one conductor arranged in a three-dimensional geometry around a region of interest in a configuration optimized to increase sensitivity to magnetic resonance signals emitted within the region of interest.

Some embodiments include a radio frequency component configured to respond to a magnetic resonance signal, the radio frequency component comprising: a first coil; the first coil comprising a first conductor arranged into a plurality of turns, the first coil oriented to respond to a first magnetic resonance signal component; and a second coil comprising a second conductor arranged into a plurality of turns, the second coil oriented to respond to a second magnetic resonance signal component.

Some embodiments include a radio frequency component configured to respond to magnetic resonance signals, the radio frequency component comprising: a first coil; the first coil including a first conductor arranged into a plurality of turns, the first coil oriented to respond to magnetic resonance signal components along a first principal axis; and a second coil including a second conductor arranged into a plurality of turns, the second coil oriented to respond to magnetic resonance signal components along a second principal axis different from the first principal axis.

Some embodiments include a radio frequency component configured to respond to magnetic resonance signals, the radio frequency component comprising: a first coil including a first conductor having a plurality of turns arranged around a region of interest; and a second coil including a second conductor having a plurality of turns arranged around the region of interest and offset from the first coil away from the region of interest.

Some embodiments include a radio frequency component configured to respond to magnetic resonance signals, the radio frequency component comprising at least one conductor arranged in a three-dimensional geometry about a region of interest, wherein a coil configuration of the at least one conductor in the three-dimensional geometry is determined based at least in part on performing at least one optimization using a model of the radio frequency coil.

Some embodiments include a method for determining a configuration of a radio frequency coil, the method comprising: generating a model of the radio frequency coil; and performing optimization to determine a model configuration that satisfies at least one constraint and produces a magnetic field that satisfies predetermined criteria when operation of the model is simulated.

Some embodiments include a radio frequency coil configured for a patient body part, the radio frequency coil comprising at least one conductor arranged in a plurality of turns around a region of interest and oriented to respond to magnetic resonance signal components oriented substantially orthogonally to a longitudinal axis of a target body tissue of the patient.

Some embodiments include an apparatus for a magnetic resonance imaging system, the apparatus comprising a first coil and at least one controller configured to operate the coil to generate a radio frequency magnetic field and a gradient field.

Some embodiments include a radio frequency coil configured for use with a human body tissue portion, the radio frequency coil comprising at least one conductor arranged in a three-dimensional geometry around a region of interest, the at least one conductor forming a plurality of turns, wherein spacing between the plurality of turns is uneven.

Some embodiments include a low-field magnetic resonance imaging system comprising: a B0 magnet configured to generate a low-field strength B0 magnetic field to provide a field of view; a first coil configured to respond to a first magnetic resonance signal component emitted from the field of view; and a second coil configured to respond to a second magnetic resonance signal component emitted from the field of view.

BRIEF DESCRIPTION OF THE DRAWINGS

Various aspects and embodiments of the disclosed technology will be described with reference to the following drawings. It should be understood that the drawings are not necessarily drawn to scale.

FIG1 shows a block diagram of an exemplary magnetic resonance imaging system according to some embodiments;

2A and 2B illustrate dual-plane magnet geometries according to some embodiments;

3A and 3B illustrate exemplary head coils according to some embodiments;

4A and 4B illustrate various methods of determining a configuration of a radio frequency coil according to some embodiments;

5 illustrates a method of determining a configuration of a radio frequency coil using a model of the radio frequency coil including a mesh, according to some embodiments;

FIG6A illustrates an exemplary triangular mesh for a model of an exemplary head coil according to some embodiments;

FIG6B illustrates an exemplary triangular mesh for a model of an exemplary leg coil, according to some embodiments;

FIG7A shows an optimized phantom configuration for an exemplary head coil according to some embodiments;

FIG7B shows an optimized model configuration for an exemplary leg coil, according to some embodiments;

8A and 8B illustrate exemplary coil configurations determined from the optimized model configuration shown in FIG. 7A , according to some embodiments;

9A and 9B illustrate exemplary coil configurations determined from the optimized model configuration shown in FIG. 7B , according to some embodiments;

10A and 10B show views of a support surface having slots for accommodating conductors according to the coil configuration shown in FIGs. 8A and 8B ;

FIG11 shows a support surface having slots for accommodating conductors according to the coil configuration shown in FIG9A and 9B;

12 illustrates a method of determining a coil configuration and applying the coil configuration to a support structure according to some embodiments;

FIG13A shows a B0 magnet arranged in a bi-planar geometry;

FIG13B shows B0 magnets arranged in a cylindrical geometry;

FIG13C shows a coil configuration for a head coil depicted using a set of orthogonal axes;

FIG13D shows a coil configuration for a leg coil depicted using a set of orthogonal axes;

14A and 14B illustrate, respectively, a model configuration for a head coil and a coil configuration determined from the model configuration, according to some embodiments;

15A and 15B illustrate, respectively, a model configuration for a leg coil and a coil configuration determined from the model configuration, according to some embodiments;

16A and 16B illustrate coil configurations applied to substrates for head coils and leg coils, respectively, according to some embodiments;

FIG17 illustrates a conductor applied to a substrate according to a coil configuration, according to some embodiments;

18A and 18B illustrate coil configurations for head coils having major axes that are substantially orthogonal to each other, according to some embodiments;

FIG18C illustrates a combination of the coil configurations shown in FIG18A and FIG18B according to some embodiments;

19A and 19B illustrate coil configurations for leg coils having major axes that are substantially orthogonal to one another, according to some embodiments;

FIG19C illustrates a combination of the coil configurations shown in FIG19A and FIG19B according to some embodiments;

FIG20 illustrates a combined coil configuration for a head coil having major axes that are substantially orthogonal to one another, according to some embodiments;

21 illustrates an exemplary coil configuration for a head coil applied to a separate base layer of a support structure, according to some embodiments;

22A and 22B illustrate views of individual base layers of a support structure for an exemplary coil configuration for an application head coil, according to some embodiments;

23A and 23B show views of a head coil having conductors arranged according to respective coil configurations having major axes that are substantially orthogonal to one another;

FIG. 24 illustrates an exemplary coil configuration for a leg coil applied to a separate base layer of a support structure, according to some embodiments;

FIG25 shows a leg coil having conductors arranged according to respective coil configurations having major axes that are substantially orthogonal to one another;

26A and 26B illustrate a controller configured to operate a multi-function coil according to some embodiments; and

27 illustrates a controller configured to operate a multifunction coil using a specific geometry of a gradient coil, in accordance with some embodiments.

DETAILED DESCRIPTION

The MRI scanner market is dominated by high-field systems and is specifically designed for medical or clinical MRI applications. As mentioned above, the general trend in medical imaging is to produce MRI scanners with increasingly larger field strengths, with the vast majority of clinical MRI scanners operating at 1.5T or 3T, while higher field strengths of 7T and 9T are used in research environments. As used herein, "high field" generally refers to MRI systems currently used in clinical settings, and more specifically refers to MRI systems operating at a main magnetic field (i.e., B0 field) of 1.5T or higher, although clinical systems operating between 0.5T and 1.5T are also generally considered to be "high field". In contrast, "low field" generally refers to MRI systems operating at a B0 field of less than or equal to about 0.2T, although systems with a B0 field between 0.2T and about 0.3T are sometimes considered to be low field due to the increased field strength in high-field systems.

Low-field MRI has been adopted in limited environmental conditions for non-imaging research purposes and narrow and specific contrast-enhanced imaging applications, but low-field MRI is generally considered unsuitable for producing clinically useful images, particularly at field strengths substantially below 0.2 T (e.g., 100 mT or lower). For example, resolution, contrast, and/or image acquisition time are generally not considered suitable for clinical purposes such as, but not limited to, tissue differentiation, blood flow or perfusion imaging, diffusion-weighted (DW) or diffusion tensor (DT) imaging, functional MRI (fMRI), etc. The inventors have developed technology for producing a portable and/or low-cost low-field MRI system with improved quality, which can improve the wide-range deployability of MRI technology in a variety of environments beyond the large MRI equipment of hospitals and research institutions.

The challenge of low-field MRI is the relatively low signal-to-noise ratio. In particular, the signal-to-noise ratio of MR signals is related to the strength of the main magnetic field B0 and is one of the factors that drive clinical systems to operate under high-field systems. Therefore, due to the low field strength, the MR signal intensity is relatively weak in a low-field environment, increasing the importance of being able to detect as many signals as possible. Some aspects of the inventor's contribution stem from their following understanding: the performance of a low-field MRI system can be improved by optimizing the configuration of radio frequency (RF) transmitting and/or receiving coils (referred to herein as RF transmitting/receiving coils or simply referred to as RF coils) to improve the ability of the RF transmitting/receiving coils to transmit magnetic fields and detect the transmitted MR signals. As mentioned above, low-field MRI systems produce weaker MR signals than their relatively high-field MRI systems, making it more important in terms of lower signal-to-noise ratio (SNR) that the RF transmitting/receiving coils operate optimally (e.g., by transmitting optimal magnetic pulses and detecting the majority of the transmitted MR signals with as much fidelity as possible).

Briefly, MRI involves placing the subject to be imaged (e.g., the entire body or a portion of a patient) in a static, uniform magnetic field, B0, so that the net magnetization of the subject's atoms (typically represented by a net magnetization vector) is aligned along the direction of the B0 field. One or more transmit coils are then used to generate a pulsed magnetic field, B1, having a frequency related to the precession rate of the atomic spins of the atoms in the magnetic field, B0, so that the net magnetization of the atoms has a component in a direction transverse to the direction of the B0 field. After the B1 field is turned off, the transverse component of the net magnetization vector precesses, its amplitude decaying over time, until the net magnetization is realigned with the direction of the B0 field—if permitted. This process generates an MR signal that can be detected, for example, by an electrical signal induced in one or more receive coils of the MRI system, which are tuned to resonate at the frequency of the MR signal.

The MR signal is a rotating magnetic field, often referred to as a circularly polarized magnetic field, which can be considered to include linearly polarized components along orthogonal axes. That is, the MR signal is composed of a first sinusoidal component oscillating along a first axis and a second sinusoidal component oscillating along a second axis orthogonal to the first axis. The first and second sinusoidal components oscillate 90° out of phase with each other. An appropriately arranged coil tuned to the resonant frequency of the MR signal can detect the linearly polarized component along one of the orthogonal axes. In particular, an electrical response can be induced in a tuned receiving coil by a linearly polarized component of the MR signal oriented along an axis approximately orthogonal to the coil's current loop, which is referred to herein as the coil's principal axis.

MRI is therefore performed by exciting and detecting emitted MR signals using transmit/receive coils (also interchangeably referred to as radio frequency (RF) coils or Tx/Rx coils). These transmit/receive coils can include separate coils for transmission and reception, multiple coils for transmission and/or reception, or the same coil for both transmission and reception. To transmit the excitation pulse sequence and detect the emitted MR signals, the transmit/receive coils must resonate at a frequency that depends on the strength of the B0 field. Therefore, compared to transmit/receive coils in low-field systems, transmit/receive coils in high-field systems must resonate at significantly higher frequencies (shorter wavelengths). The length of the resonant coil's conductive path is constrained by the frequency at which the resonant coil is intended to resonate. Specifically, the higher the frequency, the shorter the conductive path required for the resonant coil to operate satisfactorily. Therefore, the conductive path of a high-field transmit/receive coil needs to be very short. To meet this requirement, high-field transmit/receive coils are typically single-turn conductive loops formed by etching, cutting, or grinding a conductive sheet (e.g., copper). The conductive path for a typical high-field transmit/receive coil is limited to a length of tens of centimeters.

The low frequencies involved in low-field MRI allow for relatively long conduction paths for transmit/receive coils, thereby allowing for coil designs that would be unsuitable (or unusable) for high-field MRI due to the constraints on conduction path length imposed by the high frequencies involved in high-field MRI. According to some embodiments, a transmit/receive coil can be formed using a single conduction path disposed above a three-dimensional curved surface corresponding to a region of interest. Due in part to the relaxation of the constraints on conductor length, the transmit/receive conduction path can be arranged above the three-dimensional curved surface in multiple turns or loops. As used herein, a "turn" refers to a conductive path disposed 360° or approximately 360° around a reference axis (e.g., the main axis of the coil, as discussed in further detail below). It should be understood that a turn need not form a closed loop, as long as the conductive path forms approximately 360° around the reference axis. For example, a conductor arranged in a spiral geometry can include multiple turns, even though each turn does not form a closed loop. An exemplary coil having a conductor arranged in multiple turns will be discussed in further detail below. By providing a coil with a plurality of turns (eg, 5 turns, 10 turns, 15 turns, 20 turns, 30 turns, 50 turns or more), the sensitivity of the coil in response to MR signals may be improved.

The increase in allowable conductor length also allows coils with a single conductor that can be arranged to cover any geometric shape to facilitate the configuration of the transmit/receive coil for a desired portion of body tissue (anatomy). For example, to image the head, a low-field transmit/receive head coil can be created by winding a conductor around a substrate manufactured to be worn by a person as a helmet. The conductor can be arranged, for example, by arranging (e.g., winding) the conductor in a spiral geometry around a curved surface of a helmet to provide sufficient coverage to deliver transmit pulses to a region of interest (e.g., the brain or a portion thereof) and/or detect MR signals emitted from the region of interest. As another example, to image the torso or appendage (e.g., the leg or a portion thereof, such as the knee), the conductor can similarly be arranged around a curved surface in a spiral geometry configured to accommodate the desired body tissue.

The above-mentioned transmit/receive coil geometry can be achieved through aspects of the low-field system. As mentioned above, the low field strength allows the use of significantly longer conductive paths. In addition, clinical high-field MRI systems typically generate the B0 field via a solenoid coil wrapped around a cylindrical bore, into which the patient being imaged is placed. Therefore, the B0 field is oriented along the longitudinal axis of the bore and the body placed therein. To perform MRI, the transmit/receive coil generates a B1 field perpendicular to the B0 field and detects MR signals emitted in this transverse direction. This limits the geometry of the transmit/receive coil designed for high-field MRI. Low-field MRI facilitates the design of an "open" system in which the B0 field is generated using, for example, a dual-plane magnet with the patient being imaged placed therebetween, so that the B0 field is substantially oriented perpendicular to the longitudinal axis of the body. For example, any of the low-field systems described in U.S. application Ser. No. 14/845,652, filed on Sep. 4, 2015 (the '652 application), entitled “Low-field Magnetic Resonance Imaging Methods and Apparatus,” or U.S. application Ser. No. 14/846,255, filed on Sep. 4, 2015 (the '255 application), entitled “Ferromagnetic Augmentation for Magnetic Resonance Imaging,” the entire contents of each of which are incorporated herein by reference.

Thus, the transmit/receive coils are arranged to generate and/or detect magnetic fields transverse to the BO field, allowing for geometries not possible in conventional high-field MRI systems. Consequently, arrangements in which the BO magnet is configured to generate a BO field transverse to the body axis (e.g., a biplane BO magnet) allow for the design of transmit/receive coils that generate/detect magnetic fields oriented along the body's axis, some examples of which are described in further detail below. Transmit/receive coils configured to respond to MR signal components oriented substantially along the longitudinal axis of the body or specific target body tissue (i.e., coils configured with their principal axis substantially aligned with the body's longitudinal axis) generally cannot be used with BO coils that generate magnetic fields aligned with the body axis, such as those typically used in high-field MRI. However, it should be understood that the transmit/receive coils can also be configured to perform MR signal detection with MRI systems having BO magnets that generate BO magnetic fields aligned with the body's longitudinal axis (e.g., BO magnets having a solenoid geometry). In particular, according to some embodiments, an RF coil is provided that includes a conductor having a plurality of turns and is configured to respond to MR signal components oriented perpendicular to the longitudinal axis of the body, some examples of which are described in further detail below.

The inventors have recognized that one or more of the different factors associated with transmit/receive coils in high-field and low-field environments can contribute to optimizing the design of transmit/receive coils for low-field MRI. To this end, the inventors have developed techniques for optimizing the configuration of RF coils to improve the performance of coils for low-field MRI systems.

The inventors have recognized that factors arising from low-field environments favor the use of magnetic field synthesis techniques to generate generally optimal coil designs for RF coils. Magnetic field synthesis is a technique used to model a coil and simulate the magnetic field generated by the modeled coil when energized. The coil model's parameters can then be adjusted based on specific criteria that define one or more constraints on the coil model and/or its parameters to find a set of parameters that generates a desired magnetic field. Due to several factors, magnetic field synthesis techniques have generally not been applicable to the design of RF coils for high-field MRI systems. In particular, such magnetic field synthesis techniques are ineffective in designing RF coils for high-field MRI systems, in part because such coils require relatively high frequencies for resonance when used in high-field systems. Specifically, the higher the operating frequency, the shorter the current paths required for transmission and reception. Therefore, known magnetic field synthesis techniques are not useful for designing receive coils with short current paths (e.g., those required for detecting MR signals in high-field environments). For example, magnetic field synthesis techniques may not be useful or desirable for configuring single-turn conductors with short current paths, which are commonly used in high-field MRI.

As described above, the significantly lower operating frequencies used for transmission and reception in low-field MRI (i.e., the significantly lower frequencies of the transmit pulses and emitted MR signals) allow for significantly longer current paths compared to high-field MRI. This has led to innovative new designs for RF coils used in low-field MRI systems. For example, a general rule of thumb is that the length of the conductor in a resonant coil should not exceed one-tenth of the wavelength at the resonant frequency. Thus, a high-field MRI system with a 3T B0 magnetic field operates at approximately 128 MHz and therefore has a wavelength of approximately 2.3 meters. Therefore, the length of the conductors in the transmit/receive coils used in such a high-field system should not exceed 23 centimeters. In contrast, a low-field MRI system with a 0.1T B0 field operates at approximately 4.3 MHz and therefore has a wavelength of approximately 70 meters, so the transmit/receive coils can include conductors up to approximately 7 meters in length. A low-field MRI system with a 0.5T B0 field operates at approximately 2.15 MHz (approximately 140 meters in wavelength), and the corresponding transmit/receive coils can use conductors up to 14 meters in length, and so on. The inventors have recognized that the significantly longer conductor lengths allowed in the low field regime allow transmit/receive coil configurations that are not possible in the high field regime. Furthermore, the increased conductor length facilitates the use of magnetic field synthesis to determine the optimal transmit/receive coil configuration.

The inventors have recognized that magnetic field synthesis techniques can be used to design RF coils for low-field MRI and have developed techniques for optimizing the configuration of RF coils to improve transmission efficiency and/or improve the efficacy of detecting MR signals transmitted in a low-field MRI environment. The inventors have developed RF coil configurations that increase the sensitivity of MR signal detection, thereby improving the signal-to-noise ratio (SNR) of the system.

As described above, MR signals are either rotationally polarized magnetic fields or circularly polarized magnetic fields. The inventors have developed an RF coil design configured for use in low-field systems, comprising multiple coils each having a different principal axis to respond to differently oriented magnetic field components of the MR signal (referred to herein as MR signal components), thereby improving the signal-to-noise ratio (SNR) of MR signal detection. For example, a first coil and a second coil can be arranged with respective principal axes that are orthogonal or substantially orthogonal to one another (i.e., orthogonal coils) to respond to orthogonal components of the transmitted MR signal (e.g., to detect orthogonal linearly polarized components of a circularly polarized MR signal). In this manner, the pair of coils obtains dual measurements of MR signals that are phase-shifted by 90°, which can be used to improve the signal-to-noise ratio (SNR) of MR signal detection by, for example, combining the dual measurements as discussed in further detail below.

It should be understood that the respective principal axes of the plurality of coils may be oriented relative to one another in other relationships (e.g., non-orthogonal relationships). For example, it may be difficult to achieve orthogonality of the principal axes of a pair of coils for a given curved surface. Generally, the closer the principal axes of a pair of coils are to orthogonality, the greater the improvement in SNR, up to an improvement of the square root of two. Furthermore, coils whose principal axes are not orthogonal may exhibit mutual inductance, and it may be necessary to configure the respective coils in a manner that mitigates this mutual inductance, some of which techniques are described in further detail below.

According to some embodiments, an RF transmit/receive component configured to respond to MR signals includes: a first coil formed from at least one conductor arranged into a plurality of turns or loops according to a first coil configuration having a first principal axis; and a second coil formed from at least one conductor arranged into a plurality of turns or loops according to a second coil configuration having a second principal axis different from the first principal axis. For example, the first and second coil configurations can be such that the first and second principal axes are substantially orthogonal to each other, although other relationships between the respective principal axes can also be used. In this manner, the first and second coils can detect different components of the MR signal (e.g., orthogonal linearly polarized components of a circularly polarized MR signal) to improve the signal-to-noise ratio (SNR) of the MR signal detection. According to some embodiments, the first and second coil configurations for the first and second coils, respectively, are determined using magnetic synthesis techniques, although other techniques (e.g., human intuition, experience, etc.) can also be used to determine the coil configurations, as the present invention is not limited in this respect. According to some embodiments, the first and second coils are arranged on separate layers of a support structure to provide an RF transmit/receive component with an improved signal-to-noise ratio (SNR), some examples of which are described in further detail below.

The inventors also realized that the optimal configuration of the receiving coil may vary from individual to individual. For example, the size and shape of an individual's head may affect the optimal configuration of the RF coil for that individual. To address this variability, the inventors have developed a technique for optimizing one or more receiving coils for a specific individual. According to some embodiments, measurements of a target body tissue of a specific individual (e.g., head measurements, torso measurements, appendage measurements, etc.) are obtained, and the obtained measurements are used to perform the optimization techniques described herein. Thus, an optimal configuration for the receiving coil can be obtained for a specific individual. According to some embodiments, a support (e.g., a helmet) for the receiving coil for the target body tissue is manufactured (e.g., via three-dimensional (3D) printing) based on the determined optimal configuration. Thus, an optimal RF coil can be produced quickly and economically and can be customized for a specific individual and/or a portion of body tissue.

The technology described in this article enables the radio frequency component to have improved sensitivity to MR signals, thereby increasing the signal-to-noise ratio of MR signal detection. As mentioned above, relatively weak MR signals are a challenge for low-field MRI. Therefore, the transmit/receive components produced using one or more of the technologies described herein are beneficial for low-field MRI systems that can obtain clinically useful images (e.g., images with a resolution suitable for clinical purposes such as diagnosis, treatment and/or research purposes). In this regard, some embodiments include a low-field MRI system comprising a radio frequency coil having at least one conductor arranged in a three-dimensional geometry around an area of interest in a configuration optimized to increase the sensitivity to MR signals emitted within the area of interest. For example, a low-field MRI system can include a B0 magnet configured to generate a low-field strength (e.g., between 0.2T and 0.1T, between 0.1T and 50mT, between 50mT and 20mT, between 20mT and 10mT, etc.) B0 magnetic field having a field of view, wherein the RF coil is optimized to provide RF pulses to the field of view to induce an MR response and/or to detect MR signals emitted therefrom with improved efficacy.

Some embodiments include a dual-coil radio frequency component having a pair of coils configured for use in a low-field regime and oriented to respond to different MR signal components to improve the signal-to-noise ratio of MR signal detection. For example, some embodiments include a low-field magnetic resonance system comprising: a B0 magnet configured to generate a B0 magnetic field having a low field strength suitable for imaging a field of view (e.g., between 0.2T and 0.1T, between 0.1T and 50mT, between 50mT and 20mT, between 20mT and 10mT, etc.); a first coil configured to respond to a first MR signal component emitted from the field of view; and a second coil configured to respond to a second MR signal component emitted from the field of view. In this regard, in response to MR signals emitted from the field of view of the low-field B0 magnetic field, the first coil and the second coil are configured to detect MR signals at a frequency corresponding to the B0 magnetic field (i.e., in the low-field regime).

According to some embodiments, the first coil and the second coil are arranged to respond to orthogonal MR signal components (e.g., the principal axes of the first coil and the second coil are substantially orthogonal to each other) to maximize the improvement in SNR, although other arrangements may also be used. According to some embodiments, the first coil and the second coil are offset from each other relative to the field of view. According to some embodiments, the respective configurations of the first coil and the second coil are optimized, for example, using magnetic synthesis techniques, although other techniques (e.g., intuition, experience, etc.) may be used to determine the respective configurations.

In some embodiments, the B0 magnet of the low-field MRI system is arranged in a planar geometry (e.g., a single-planar or dual-planar geometry), and in other embodiments, the B0 magnet is arranged in a cylindrical geometry (e.g., a solenoid geometry), and one or more RF coils are configured to transmit RF pulses and/or detect MR signals according to the geometry of the B0 magnet.

The following is a more detailed description of various concepts related to methods and apparatus for generating RF coils, such as for use in low-field MRI, and embodiments of methods and apparatus for generating RF coils, such as for use in low-field MRI. It should be understood that the embodiments described herein can be implemented in any of a number of ways. Examples of specific implementations are provided below for illustrative purposes only. It should be understood that the embodiments and features/capabilities provided can be used individually, together, or in any combination of two or more, as the aspects of the technology described herein are not limited in this respect.

FIG1 is a block diagram of exemplary components of an MRI system 100 (e.g., a low-field MRI system). In the illustrative example of FIG1 , the MRI system 100 includes a computing device 104, a controller 106, a pulse sequence memory 108, a power management system 110, and a magnetic component 120. It should be understood that the system 100 is illustrative and that the MRI system may have one or more other components of any suitable type in addition to or in place of the components shown in FIG1 .

As shown in FIG1 , the magnetic element 120 includes a BO magnet 122, shim coils 124, RF transmit and receive coils 126, and gradient coils 128. The magnet 122 can be used to generate a main magnetic field BO. The magnet 122 can be any suitable type of magnetic component or combination thereof that can generate a desired main magnetic field BO (e.g., any one or combination of an electromagnet, a printed magnet, a permanent magnet, etc.). Therefore, in this document, a BO magnet refers to any one or combination of any type of magnetic component configured to generate a BO field. According to some embodiments, the BO magnet 122 can generate or contribute a BO field greater than or equal to about 20 mT and less than or equal to about 50 mT, greater than or equal to about 50 mT and less than or equal to about 0.1 T, greater than or equal to about 0.1 T and less than or equal to about 0.2 T, greater than or equal to about 0.2 T and less than or equal to about 0.3 T, greater than 0.3 T and less than or equal to about 0.5 T, etc. Shim coils 124 may be used to contribute magnetic fields to improve the uniformity of the B0 field generated by magnet 122 .

The gradient coils 128 can be arranged to provide gradient fields and, for example, can be arranged to generate gradients in the B0 field along three substantially 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 according to frequency or phase. For example, the gradient coils 128 can be configured to vary the frequency or phase as a linear function of 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 perform frequency encoding in that direction, a second gradient coil can be configured to selectively vary the B0 field in a second (Y) direction substantially orthogonal to the first direction to perform phase encoding, and a third gradient coil can be configured to selectively vary the B0 field in a third (Z) direction substantially orthogonal to the first and second directions to enable slice selection for volumetric imaging applications.

As described above, MRI is performed by using a transmitting coil and a receiving coil (commonly referred to as a radio frequency (RF) coil) to excite and detect the transmitted MR signals, respectively. The transmit/receive coil can include a separate coil for transmitting and receiving, multiple coils for transmitting and/or receiving, or the same coil for transmitting and receiving. Therefore, the transmit/receive component can include one or more coils for transmitting, one or more coils for receiving, and/or one or more coils for transmitting and receiving. The transmit/receive magnetic component is also commonly referred to as Tx/Rx or Tx/Rx coil to collectively refer to various configurations of the transmit and receive magnetic components of an 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 B1. The (one or more) transmit coils can be configured to generate any suitable type of RF pulses. For example, the transmit coil may be configured to generate any of the pulse sequences described in U.S. patent application Ser. No. 14/938,430, filed Nov. 11, 2015, entitled “Pulse Sequences for Low Field Magnetic Resonance” (the '430 application), the entire contents of which are incorporated herein by reference.

Each of the magnetic components 120 can be constructed in any suitable manner. For example, in some embodiments, one or more (e.g., all) of the magnetic components 120 can be manufactured, constructed, or produced using the techniques described in U.S. patent application Ser. No. 14/845,652, filed on September 4, 2015, entitled “Low-field Magnetic Resonance Imaging Methods and Apparatus” (the '652 application), the entire contents of which are incorporated herein by reference. However, the techniques described herein are not limited in this respect, as the magnetic components 120 can be provided using any suitable technique.

The power management system 110 includes electronics that 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 may include one or more power supplies, gradient power components, transmit coil components, and/or any other suitable power electronics required to provide suitable operating power to energize and operate the components of the low-field MRI system 100.

As shown in FIG1 , the power management system 100 includes a power supply 122, one or more power components 114, a transmit/receive switch 116, and a thermal management component 118. The power supply 112 includes electronics that provide operating power to the magnetic components of the MRI system 100. For example, the power supply 112 may include electronics that provide operating power to one or more BO coils (e.g., the BO magnet 122) to generate the main magnetic field for a low-field MRI system. In some embodiments, the power supply 112 is a unipolar continuous wave (CW) power supply, however, any suitable power supply may be used. The transmit/receive switch 116 may be used to select whether the RF transmit coil or the RF receive coil is being operated.

The one or more power components 114 may include: one or more RF receive (Rx) preamplifiers that amplify MR signals detected by one or more RF receive coils (e.g., coil 126); one or more RF transmit (Tx) power components configured to supply power to one or more RF transmit coils (e.g., coil 126); one or more gradient power components configured to supply power to one or more gradient coils (e.g., gradient coil 128); and one or more shim power components configured to supply power to one or more shim coils (e.g., shim coil 124).

The thermal management component 118 provides cooling for components of the low-field MRI system 100 and can be configured to do so by facilitating the transfer of thermal energy generated by one or more components of the low-field MRI system 100 away from those components. The thermal management component 118 can include, but is not limited to, components that perform water-based or air-based cooling, which can be integrated with or disposed proximate to heat-generating MRI components, including, but not limited to, B0 coils, gradient coils, shim coils, and/or transmit/receive coils. The thermal management component 118 can include any suitable heat transfer medium, including, but not limited to, air and liquid coolant (e.g., water), to transfer heat away from the components of the low-field MRI system 100.

As shown in FIG1 , 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 the power management system 110. The controller 106 can be configured to implement one or more pulse sequences, which determine the instructions sent to the power management system 110 to operate the magnetic component 120 in a desired sequence. For example, for embodiments in which the MRI system 100 operates at low field, the controller 106 can be configured to control the power management system 110 to operate the magnetic component 120 according to the following pulse sequences: a zero echo time (ZTE) pulse sequence, a balanced steady-state free precession pulse sequence (bSSFP), a gradient echo pulse sequence, a spin echo pulse sequence, an inversion recovery pulse sequence, arterial spin labeling, diffusion-weighted imaging (DWI), and/or any other pulse sequence specific to operation in a low-field environment. Pulse sequences used for low-field MRI can be applied to different contrast types, such as T1-weighted and T2-weighted imaging, diffusion-weighted imaging, arterial spin labeling (perfusion imaging), proton precession (Overhauser) imaging, etc. However, any pulse sequence may be used, as the present aspects are not limited in this regard.As aspects of the disclosure provided herein are not limited in this regard, the controller 106 may be implemented as hardware, software, or any suitable combination of hardware and software.

In some embodiments, the controller 106 can be configured to implement the pulse sequence by obtaining information about the pulse sequence from a pulse sequence repository 108, which stores information for each of one or more pulse sequences. The information about a particular pulse sequence stored by the pulse sequence repository 108 can be any suitable information that enables the controller 106 to implement the particular pulse sequence. For example, the information about the pulse sequence stored in the pulse sequence repository 108 can include: one or more parameters for operating the magnetic component 120 according to 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 management system 110 according to the pulse sequence, one or more programs including instructions that, when executed by the controller 106, cause the controller 106 to control the system 100 to operate according to the pulse sequence, and/or any other suitable information. The information stored in the pulse sequence repository 108 can be stored on one or more non-transitory storage media.

As shown in FIG1 , the controller 106 also interacts with a computing device 104, which is programmed to process the received MR data. For example, the computing device 104 may process the received MR data using any suitable one or more image reconstruction processes to generate one or more MR images. The controller 106 may provide information regarding one or more pulse sequences to the computing device 104 for use in processing the data. For example, the controller 106 may provide information regarding one or more pulse sequences to the computing device 104, and the computing device may perform an image reconstruction process based at least in part on the provided information.

The computing device 104 can be any electronic device that can process the acquired MR data and generate one or more images of the subject being imaged. In some embodiments, the computing device 104 can be a fixed electronic device, such as a desktop computer, a server, a rack-mounted computer, a workstation, or any other suitable fixed electronic device that can be configured to process MR data and generate one or more images of the subject being imaged. Alternatively, the computing device 104 can be a portable device, such as a smart phone, a personal digital assistant, a laptop computer, a tablet computer, or any other portable device that can be configured to process MR data and generate one or more images of the subject being imaged. In some embodiments, the computing device 104 can include multiple computing devices of any suitable type, as the disclosed aspects provided herein are not limited in this regard. The user 102 can interact with the computing device 104 to control various aspects of the low-field MR system 100 (e.g., programming the system 100 to operate according to a specific pulse sequence, adjusting one or more parameters of the system 100, etc.) and/or view images obtained by the low-field MR system 100.

FIG2A and FIG2B illustrate an exemplary dual-plane geometry for a BO magnet. BO magnet 222 is schematically illustrated by magnets 222a and 222b, which are arranged substantially parallel to one another to generate a BO field generally along axis 245 regardless of orientation (the orientation desired to provide a field of view between magnets 222a and 222b (i.e., a region between the magnets where the BO field uniformity is suitable for MRI)). This dual-plane arrangement allows for a generally "open" magnetic resonance imaging system. An RF coil (or coils) is schematically illustrated as RF coil 226, which is arranged to generate a pulsed oscillating magnetic field generally along axis 225 (i.e., the principal axis of RF coil 226) to stimulate an MR response and detect MR signals. The exemplary RF coil 226 is arranged to detect MR signal components oriented substantially along principal axis 225 (i.e., linearly polarized components of the MR signal aligned with the principal axis of the coil). As described above, the relatively low operating frequencies of low-field MRI allow for coil designs that are unsuitable for use in high-field environments. The inventors have developed RF coil designs that improve the coil's ability to transmit RF pulse sequences and/or detect transmitted MR signals, some of which are discussed in further detail below. The inventors have also developed techniques for using magnetic synthesis to optimize the placement of one or more conductors of an RF coil according to desired criteria, some examples of which are also described in further detail below.

3A and 3B show several views of a radio frequency (RF) head coil configured to transmit a suitable RF pulse sequence in a low-field MRI system and detect MR signals emitted in response to the RF pulse sequence. The transmit/receive coil 300 can, for example, correspond to the RF coil 226 shown in FIG. 2 and is particularly configured to obtain MR images of the head. As shown, the transmit/receive coil 300 includes a base 350 formed to accommodate the head of the subject to be imaged. The base can be formed with slots, and the conductor 330 is arranged (e.g., wound) in the slots according to the desired geometry. The base includes a helmet portion that accommodates the head and a support base so that the patient can comfortably lean his head against the helmet in a static position.

As shown, conductor 330 is wound around base 350 in a spiral geometry so that, when in operation, the coil generates a magnetic field oriented along axis 305 and can detect magnetic fields oriented along the same axis. Thus, axis 305 corresponds to the principal axis of the coil formed by conductor 330. Conductor 330 comprises a single, continuous wire forming a single-channel transmit and receive coil. The exemplary transmit/receive coil 300 in Figures 3A and 3B has a conduction path of approximately 14 meters. As described above, the high frequencies of high-field MRI (e.g., greater than 64 MHz) require that the conduction path of the RF coil be very short (e.g., on the order of centimeters) to operate properly. Therefore, the length of the conductors in this exemplary transmit/receive coil far exceeds the limitations imposed by the high frequencies of high-field MRI systems (by an order of magnitude or more). However, the illustrated configuration is not optimized, and therefore, the performance of the head coil may be suboptimal, resulting in lower-quality images.

The inventors have developed coil configurations for improving RF coil efficacy (e.g., improving the RF pulses delivered to a subject and/or improving the sensitivity of detecting MR signals emitted in response to an RF pulse sequence). Consequently, an increased signal can be detected, resulting in a higher signal-to-noise ratio (SNR), a factor that is particularly important in low-field MRI, where MR signals are relatively weak compared to high-field MRI. The inventors have also developed techniques for determining the overall optimal arrangement (e.g., configuration) of conductors on an RF coil to improve the resulting coil's ability to detect emitted MR signals and/or transmit RF energy in a low-field environment. As discussed in further detail below, the techniques described herein can be applied to any curved surface of interest to provide RF coils having any desired geometry for any portion or portions of body tissue (e.g., head, torso, arm, leg, knee, etc.).

FIG4A illustrates a method for determining an RF coil configuration according to some embodiments. In act 410, a model of an RF coil is provided. The term "model" herein refers to any mathematical representation of an RF coil or a representation from which a representation of the RF coil can be derived. For example, a model of an RF coil can include a geometric representation such as a triangular mesh or other representation constructed from geometric primitives. Additionally, the model can be described by implicit surfaces and/or can include other types of suitable mathematical representations or combinations thereof. A suitable model typically allows for magnetic field synthesis to be performed using the model, for example, by enabling simulation of the operation of the modeled RF coil to synthesize the magnetic field generated within a region of interest under the simulated operation. A model typically has one or more parameters that, when set to a given set of corresponding values, characterize a specific configuration of the model. Changing the value of one or more of the parameters changes the configuration of the model. An optimized RF coil configuration can be derived from the optimized model configuration by finding the optimal model configuration (e.g., a set of one or more parameters describing the model) according to a given criterion, as discussed in further detail below.

In act 420A, a configuration of the RF coil is determined using the model of the RF coil. For example, the model can be used to perform optimization to determine an RF coil configuration that satisfies at least one constraint and produces a magnetic field that meets at least one criterion when the model is simulated. In some embodiments, the at least one criterion for the magnetic field includes magnetic field uniformity. For example, the optimization can be customized to identify a model configuration that produces a magnetic field that meets a uniformity criterion (e.g., less than a specified percentage of non-uniformity) within a region of interest when the RF coil model is simulated. In some embodiments, the at least one criterion includes a magnetic field strength criterion. Any suitable criterion or combination of criteria that facilitates determining a desired RF coil configuration from an optimized configuration of the model can be used. In some embodiments, different parameters are used to describe the model configuration and the RF coil configuration. For example, the model configuration can represent a surface potential with parameters corresponding to current density, and the RF coil configuration can represent the arrangement of conductors (e.g., wires) in three-dimensional space. In some embodiments, an optimal model configuration can be identified (e.g., by determining an optimal set of parameters according to a given criterion), and the RF coil configuration can be determined based on the optimized model configuration. Determining the RF coil configuration can involve a second optimization, but in other embodiments, the RF coil configuration is determined in other ways (e.g., determining the optimal coil configuration can involve multiple stages). According to some embodiments, the optimal RF coil configuration is determined in conjunction with optimizing the model configuration. For example, the model configuration and the RF coil configuration can be parameterized similarly, such that the optimal RF coil configuration is generally determined by optimizing the model configuration based on how the RF coil is modeled.

As described above, optimization can include finding optimal parameter values that satisfy a given criterion in view of at least one constraint. The at least one constraint can be any constraint or combination of constraints that facilitates a configuration (model and/or coil configuration) that satisfies one or more design specifications for the RF coil. In some embodiments, the at least one constraint includes the resistance of the RF coil configuration. For example, the optimization can impose a maximum resistance for the RF coil configuration or otherwise minimize the coil resistance when determining the optimal configuration relative to the given criterion. In some embodiments, the at least one constraint includes the inductance of the RF coil. For example, the optimization can impose a maximum inductance for the RF coil configuration or otherwise minimize the coil inductance when determining the optimal configuration relative to the given criterion. Any other constraint or combination of constraints can additionally or alternatively be used to determine the coil configuration, some examples of which are described in further detail below.

As a result of executing step 420A, a configuration of the RF coil is obtained. As described above, the coil configuration can be defined by a set of parameters describing the RF coil. According to some embodiments, the RF coil configuration describes the three-dimensional geometry of one or more conductors (e.g., describing how the one or more conductors are arranged in three-dimensional space). For example, the configuration can describe the number of turns or loops, the spacing between turns of at least one conductor of the RF coil, and/or any other description of how the at least one conductor is arranged. The configuration can be any description of how the one or more conductors of the RF coil are arranged on a curved surface of interest and/or any description of the properties and/or attributes of the one or more conductors, as these aspects are not limited in this regard. According to some embodiments, the coil configuration is determined from a model configuration obtained by optimizing one or more parameters of the RF coil model. For example, the RF coil configuration can include parameters controlling the number of turns of the RF coil, the spacing between turns, and/or the position of the conductors (e.g., one or more wires in the coil). Generally, the one or more parameters of the RF coil configuration at least partially define the distribution and/or arrangement of the physical conductors on the curved surface of interest of the physical RF coil. More details regarding some examples of coil optimization techniques are discussed below.

FIG4B illustrates a method for optimizing an RF coil configuration according to some embodiments. In action 410B, a model of the RF coil is obtained. The model may be obtained or provided using any of the techniques discussed above in conjunction with FIG2A or using any suitable technique for providing a representation of an RF coil.

In action 420B, the operation of the modeled RF coil is simulated for a specific configuration of the model. For example, given a specific model configuration, the magnetic field generated by simulating the operation of the model is synthesized. According to some embodiments, the simulation may involve synthesizing the magnetic field generated at a set of points within the region of interest by simulating currents on the curved surface of the model of the RF coil. In action 422, the synthesized magnetic field is compared with a given criterion to evaluate whether the configuration is satisfactory from an optimization perspective (e.g., whether it meets the given criterion). According to some embodiments, the criterion may take the form of a function having one or more constraints and/or one or more variables to be minimized or maximized. For example, the optimization of the function may seek to maximize the magnetic field generated within the region of interest while minimizing the inductance and/or resistance of the RF coil (or constraining the inductance and/or resistance to be below corresponding specified values). However, any set of variables considering any set of constraints may be used, as the techniques described herein are not limited to use with any specific optimization or optimization scheme.

A specific coil design and design constraints may at least partially dictate what factors to consider when optimizing the configuration of an RF coil. Non-limiting factors that may be evaluated in an optimization equation for an RF coil design (e.g., in the form of variables to be minimized or maximized, or as constraints) include any one or a combination of magnetic field strength, magnetic field uniformity, coil efficiency/sensitivity, coil inductance, coil resistance, wire length, wire thickness, wire spacing, and the like. The relative importance of any one or a combination of these factors may be weighted to arrive at an optimal configuration based on the given design constraints.

If, in act 422, it is determined that the solution (e.g., the evaluation of a given function) resulting from simulating operation of the model with the current configuration is optimal according to a predetermined criterion, processing proceeds to act 460, where an RF coil configuration is determined based on the model configuration. For example, the optimized model configuration can be used to determine a coil configuration that, when in operation, will produce a magnetic field substantially similar to the magnetic field simulated from the model configuration. According to some embodiments, the coil configuration is determined from the optimized model configuration by determining a line profile of the RF coil based at least in part on the model configuration. For example, a line profile of the optimized RF coil can be determined using a profiling technique (examples of which are discussed below), and the line profile can then be used to generate an actual physical RF coil as shown in act 470, further aspects of which are described below. That is, the profile describes the coil configuration and can be used as a pattern for arranging the physical conductors of the RF coil.

If it is determined in act 422 that the solution is not optimal according to a predetermined criterion (e.g., does not meet a given criterion), processing proceeds to act 430, where one or more parameters of the model may be modified to produce an updated model configuration. To optimize the model configuration, processing returns to act 420 to simulate the operation of the RF coil using the updated model configuration, and the process iterates until an optimal configuration is determined (e.g., a set of one or more parameters controlling the model configuration is optimal according to a given criterion). The manner in which the configuration is updated for the next iteration may be selected according to any suitable optimization scheme. By repeating acts 420, 422, and 430, the configuration of the RF coil model may be optimized according to certain criteria characterized by the criteria (e.g., by optimizing a suitable function). From the final model configuration, an overall optimal RF coil configuration may be obtained. It should be understood that optimizing the model configuration and/or RF coil configuration need not result in a globally or absolutely optimal solution, but only requires convergence to some sufficiently "optimal" criterion. Therefore, for a given model and formulation, there may be many "optimal" solutions.

FIG5 illustrates an exemplary implementation of the general method described in FIG4 , according to some embodiments. In act 510 , a model of an RF coil is configured using a curved three-dimensional mesh corresponding to a region of interest (e.g., the field of view of a low-field MRI system) where RF energy is to be provided and MR signals are to be detected. According to some embodiments, the mesh comprises a plurality of surface elements connected by nodes at the element's vertices. Non-limiting examples of meshes having triangular surface elements that can be used as the basis for an RF coil model, according to some embodiments, are illustrated in FIG6A and FIG6B . Specifically, FIG6A illustrates an exemplary mesh 600A corresponding to a head coil. Mesh 600A is formed from a plurality of triangles (e.g., triangle 610 ) connected by shared edges with one or more adjacent triangles. Each triangle vertex or node (e.g., node 620 ) is shared by one or more adjacent triangles, but any suitable configuration of surface elements may be used to form the mesh. In some embodiments, the mesh includes approximately 1000-4000 triangles, but it will be appreciated that any suitable number of triangles may be used, and the number and/or shape of triangles in the mesh may depend at least in part on the surface being modeled.

FIG6B shows an exemplary mesh 600B corresponding to an RF coil suitable for imaging a leg, such as a knee or other portion of a leg. Similar to the mesh in FIG6A , the desired surface is triangulated to form a plurality of triangles (e.g., triangle 610) interconnected at shared vertices or nodes (e.g., node 620). It should be understood that the exemplary surfaces shown in FIG6A and FIG6B are merely illustrative, and any desired primitives can be used to define a mesh for any geometric shape. In other words, a surface primitive having any geometric shape (e.g., triangle, square, hexagon, octagon, etc.) can be used to define a mesh on any surface. It should also be understood that using a mesh is merely one example of a geometric representation that can be used to provide a model of an RF coil.

Meshes such as those shown in FIG6A and FIG6B provide a flexible representation for modeling RF coils because any arbitrary curved surface can be represented using the mesh, thereby facilitating modeling RF coils for any desired portion of body tissue of a human body and/or any combination of body tissue portions to produce RF coils optimized for use with any desired portion of body tissue of a human body, including but not limited to the head, neck, torso, one or more appendages or portions thereof (e.g., arms, legs, hands, feet or portions thereof).

Referring again to FIG. 5 , in act 520 , the operation of the model of the RF coil can be simulated. For example, using the exemplary triangular mesh 600A or 600B shown in FIG. 6A and FIG. 6B , the operation of the model can be simulated at least in part by simulating current loops around each node in the mesh (e.g., simulating current loops through adjacent triangles by simulating current loops around shared nodes of adjacent triangles) and calculating the magnetic field generated by the corresponding current loops at designated target points selected within the region of interest. Specifically, a large number of target points can be selected (e.g., 100-1000 designated points within the interior of the triangular mesh) at which the magnetic field resulting from the simulated current loops around the nodes of the triangular mesh can be calculated. Generally, the target points are selected and distributed in a manner suitable for characterizing the magnetic field of the entire region of interest. The region of interest can be associated with the field of view of an imaging system, for example, but can also correspond to other regions of interest.

According to some embodiments, current loops are simulated at each node in the grid, and the resulting magnetic field generated by each current loop at each target point is determined to obtain information about the impact of each current loop on each target point. For example, operation of the simulation model in this manner can be used to obtain a data matrix corresponding to the magnetic field generated at each target point in response to each corresponding current loop being simulated. This data can in turn be operated on by a suitable optimization algorithm, an example of which is described in further detail below. According to some embodiments, the strength of each current loop forms at least in part a set of parameters that are varied during the optimization. That is, a suitable optimization algorithm selects the strength of each current loop to, for example, maximize or minimize a given function (e.g., a potential function defined by the current loops on the grid surface) or other appropriately tailored optimization to achieve desired magnetic field characteristics at each target point in the region of interest.

After simulating the operation of the RF coil using the surface current loops, the exemplary process of FIG. 5 proceeds in a manner similar to that discussed above in conjunction with the exemplary process of FIG. 4 . For example, according to some embodiments, the operation of the model can be performed by simulating current loops around nodes of surface elements (e.g., vertices of a triangular mesh) to optimize a potential function defined on the surface of the mesh. In such an exemplary optimization, the optimized surface potential is obtained, at least in part, by iterating through actions 520, 522, and 530 shown in FIG. 5 , by varying the strength of the current loops simulated on the triangular mesh until the generated magnetic field meets a given criterion. FIG. 7A and FIG. 7B illustrate model configurations 705a and 705b, respectively. In the exemplary embodiments shown in FIG. 7A and FIG. 7B , the model configuration is characterized in part by surface potentials that have been optimized using the techniques described herein. In particular, the shading in FIG. 7A and FIG. 7B depicts the magnetic scalar surface potential (e.g., a stream function of current density, as discussed in further detail below in conjunction with the exemplary optimization shown in FIG. 12 ), whose value is determined during the optimization. Based on the surface potential function, the coil configuration can be determined, as discussed in further detail below. In the exemplary embodiment shown in Figures 7A and 7B, the surface potential function corresponds to the integrated current density on the surface, which is obtained at least in part by varying the current intensity parameters of the current loops at the nodes in the mesh until the simulation of the model configuration has been optimized to meet a given criterion taking into account one or more constraints.

In act 560, an RF coil configuration can be determined based on the model configuration. For example, the model configurations 705a and 705b shown in Figures 7A and 7B (e.g., potential functions) can be converted into contours indicating the arrangement of conductors for the RF head coil and leg coil, respectively. Figures 8A and 8B illustrate a coil configuration 815, which is characterized by contour lines (e.g., exemplary contour lines 880) for the conductors of the head coil determined based on the model configuration 705a shown in Figure 7A. For example, for the exemplary coil configuration shown in Figures 8A and 8B, the contour lines are selected to produce the current density that optimizes the model configuration (i.e., the differential of the surface potential function shown in Figures 7A and 7B). Because the contour lines of the coil configuration represent the current path of the coil that can ultimately be achieved by a single conductor (e.g., a single conductor is wound according to the contour of the coil configuration to form multiple turns or loops), each contour line has the same current. Therefore, to achieve the varying current densities described by the model configuration, the spacing of the contour lines varies accordingly. Specifically, regions of higher current density will produce contours that are spaced closer together, while regions of lower current density will produce contours that are spaced further apart. Thus, the coil configuration can be determined from the model configuration by finding contour lines of equal potential through the surface potential function of the model configuration (e.g., contour lines of equal scalar value passing through the surface potential function as shown in FIG7A and FIG7B ). Determining the coil configuration from the model configuration in this manner can be implemented, at least in part, using any suitable contouring or leveling algorithm.

FIG8A shows a coil configuration 815 overlaid on a phantom configuration from which the contour is determined, and FIG8B shows coil configuration 815 alone. The contours of the conductors used in the RF coil are selected to substantially produce the magnetic field generated when the phantom is simulated using the optimized phantom configuration. In this way, a generally optimal coil configuration 815 can be determined. In other words, the exemplary coil configuration 815, characterized by the arrangement of the contours in space, defines a conductor pattern optimized according to desired criteria. As shown, the contour of coil configuration 815 has a main axis 825 that is substantially aligned with the longitudinal axis of the body. Main axis 825 is also an exemplary reference axis around which the coil configuration forms multiple turns.

As shown in Figures 8A and 8B , in the resulting RF coil configuration, the spacing between contours (e.g., the spacing between turns in the conductor of the RF coil) is uneven, with the contours being closer together toward the base of the RF coil configuration. Thus, compared to the coil shown in Figure 3 , which has a configuration based on human intuition with substantially uniform spacing between turns of the coil's conductor across the curvature of the helmet, the optimized coil configuration shown in Figures 8A and 8B has uneven spacing between multiple contours, resulting in the resulting RF coil having uneven spacing between the multiple turns or loops of the conductor forming the RF coil. This configuration provides an optimal solution that would not be achievable using human intuition or empirical trial and error alone. Furthermore, while the coil configuration in Figure 3 has substantially regular contours, the optimized coil configuration results in multiple irregular contours. Thus, the optimization yields configuration solutions that would not be achievable using human intuition alone. 9A and 9B illustrate an optimized RF leg coil configuration 915 determined from the model configuration 705b shown in FIG. 7B , the coil configuration forming a plurality of turns about a main axis 925 that is substantially aligned with the longitudinal axis of the target body tissue (e.g., a patient's leg) when the target body tissue is positioned within the coil.

An RF coil configuration (e.g., the exemplary coil configurations shown in Figures 8A, 8B, 9A, and 9B) can then be used to generate an RF coil according to the determined configuration. For example, to generate an RF coil, the RF coil configuration typically needs to be transferred to a support structure, such as a helmet to be worn by a subject, for head coil configuration 815 shown in Figures 8A and 8B. According to some embodiments, the RF coil configuration is used to generate the RF coil by applying the contour of the RF coil configuration to a substrate, which in turn is used to secure the arrangement of conductors to the curved surface of the RF coil. Figures 10A and 10B show different views of a geometric rendering of a helmet 1000, with slots (e.g., slot 1080) formed in substrate 1050 corresponding to the positions calculated for the conductors during the exemplary optimization described above. For example, slot 1080 can be positioned corresponding to the contour of coil configuration 815 determined from model configuration 705a. In particular, the contour of the coil configuration can be mapped to the curved surface of the support structure or substrate to provide locations for applying the coil conductors (e.g., providing slot locations for the coil conductors). The dimensions of the slots (e.g., the width and depth of the slots) can be selected to accommodate conductors for forming radio frequency coils. Once the curved surface is rendered (e.g., based on a surface mesh and an optimized coil configuration), it can be manufactured (e.g., using a 3D printer) to quickly and cost-effectively produce a helmet, such as an RF head coil for low-field MRI. As shown in FIG10A , slots 1085 are provided to connect slots 1080 corresponding to the contours of the conductors (e.g., corresponding to the contours of the coil configuration obtained from the optimized model configuration). Slots 1085 allow a single conductor to be wound around the base 1150 within the provided slots to provide multiple turns of the conductor, as discussed in further detail below. When the conductor is positioned within the slots, as shown in FIG10A , the conductor will form multiple turns around the main axis 1025. When the patient's head is positioned within the helmet 1000, the main axis can be oriented substantially aligned with the longitudinal axis of the patient's body.

FIG11 illustrates an exemplary substrate or support 1100 to which the RF coil configuration 915 of FIG9A and FIG9B has been applied to create a support for a leg coil. Specifically, support 1100 includes slots (e.g., slots 1180) formed in substrate 1150 corresponding to the outline of the exemplary RF coil configuration 915 shown in FIG9A and FIG9B. Support 1100 includes slots 1185 configured to connect slots 1180 to facilitate continuous positioning of conductors within slots 1180 in a plurality of turns according to a desired coil configuration. When positioned within the slots, the conductors form a plurality of turns about a primary axis 1125.

Once a support structure (e.g., helmet 1000, support 1100, or other geometry configured for a particular body tissue) is created, conductors (e.g., wires) can be applied to the structure (e.g., by positioning the conductors within slots) to create an RF coil with an optimized coil configuration. For example, a single conductor can be positioned within a slot (e.g., a wire can be placed within slots 1080 and 1180, respectively, as shown in Figures 10 and 11) to at least partially create an RF coil with improved transmit/receive characteristics. The slots are formed in a support structure fabricated according to the geometry of the corresponding RF coil. It should be understood that the coil configuration can be applied to the support structure using suitable techniques and is not limited to slots provided in a base support structure. That is, the conductors can be coupled to the support structure in any suitable manner according to the desired coil configuration, as these aspects are not limited in this regard. RF coils created using optimized coil configurations may exhibit increased sensitivity to transmitted MR signals, thereby improving the signal-to-noise ratio (SNR) of low-field MRI systems. Additional examples of RF coils fabricated using the techniques described herein are discussed in detail below.

The ease with which such a support structure can be manufactured facilitates the creation of customized RF coils for specific individuals and/or specific parts of the body. With respect to customizing an RF coil for a specific individual, measurements for the specific individual can be obtained using a laser or other range-finding device and/or by manually measuring, for example, important dimensions of a portion of the body tissue being imaged using a caliper. The measurements and/or range data can be used to create a surface for modeling the RF coil (e.g., the measurement data can be used to render a mesh corresponding to the geometry of the body tissue of interest for a specific patient). The optimization techniques described herein can then be performed to locate the optimal RF coil configuration, which in turn can be used to generate (e.g., by 3D printing) a support for the optimal coil configuration customized for the specific patient. Thus, an optimized coil configuration can be determined, and the corresponding coil can be generated relatively quickly and efficiently for any geometry of interest.

As described above, the design of an RF coil may involve meeting specific design constraints and/or requirements. According to some embodiments, coil inductance and/or coil resistance are evaluated to constrain the optimization of the RF coil configuration. As described above, to operate properly, an RF transmit/receive coil must resonate. Therefore, an increase in inductance requires an increase in capacitance in the tuning circuit coupled to the coil to achieve resonance. Increased resistance affects the coil's quality (Q) factor by increasing the bandwidth of the coil resonance, making the coil less effective at exciting the MR effect and less sensitive at detecting transmitted MR signals. A particular system may have design requirements that dictate the coil's inductance and/or resistance (e.g., to achieve a coil with a specified Q factor or to match a specified tuning circuit). Therefore, by evaluating the coil inductance and/or coil resistance (e.g., by minimizing or limiting their values), the RF coil configuration can be optimized within the given design constraints.

According to some embodiments, a regularization scheme is used that includes additional terms corresponding to one or more design constraints (e.g., coil resistance, coil inductance, field uniformity, etc.). For example, coil inductance and/or coil resistance can be included as additional terms in the optimization. In conjunction with the example RF coil model shown in FIG6 , coil resistance and/or inductance can be calculated for each simulated current loop. Thus, data corresponding to magnetic field strength and one or more additional constraints, such as coil resistance or inductance, can be generated. For example, a magnetic field strength matrix can be calculated as a first term, and a coil resistance matrix can be calculated as a second term, wherein the optimization operation is performed to achieve the desired magnetic field characteristics while minimizing the coil resistance. It should be understood that additional terms for any desired constraints can be included in the optimization. The selected terms can be weighted as needed so that the optimization produces the desired value (e.g., the value of a function of the surface of the mesh that produces the optimal result according to the specified constraints).

It should be understood that any number or type of constraints can be included in the optimization to meet the requirements of a particular design. For example, a given design may require the use of wires of a given thickness or width. To prevent the optimization from selecting a configuration in which the wires are placed too close to each other (e.g., a solution in which the spacing between wires (e.g., turns of conductors) is less than the width of the wire at one or more locations on the curved surface), a term can be included in the optimization that imposes a minimum spacing between the wires that form one or more conductive paths of the coil. For designs with wire conductors of fixed thickness, coil resistance constraints can be implemented by including a term corresponding to the wire length in the optimization, as discussed in further detail below.

An exemplary implementation of a method for determining the configuration of an RF coil is described in more detail below in conjunction with the illustrative and non-limiting process shown in FIG12 . It should be understood that the implementation described below is merely one example of how to optimize the configuration of an RF coil, and that any other suitable technique may be used, as determining the configuration of an RF coil using a model of the RF coil is not limited to any particular implementation. In act 1210 , a surface geometry to be modeled is received. As described above, any surface geometry can be used to generate an RF coil according to the techniques described herein. In act 1212 , a model of the surface geometry is created. In this exemplary model, the surface geometry can be considered to be a thin conductive surface S defined by the unit normal vector of the surface at a point r′. The current flowing on S is represented at r′ by the current density vector J(r′). When the current density is constrained to the surface S and is non-dispersive, a potential function, a stream function, can be defined on the surface S. The current density on the surface S generates a magnetic field B(r) over a region of interest V separated from the surface S. The relationship between the generated magnetic field B(r) and the current density on the surface S can be expressed as:

The optimization can be performed at least in part by solving an inverse problem to find the current density J(r′) on the surface S that will provide a given magnetic field B(r) over the region of interest V. To solve this inverse problem, the problem can be discretized. For the surface S, the current density J(r′) can be discretized using a mesh defined by a set of flat triangular surface elements with nodes at the corners of the surface elements (e.g., as shown in FIG6 ). As described above, shapes other than triangles of surface elements can alternatively be used to form a mesh for discretizing the surface S. The stream function ψ(r′) of the current density can be discretized as a set of basis functions for each node I _n in the mesh:

In (2), _ψn (r') is the stream function basis function for the nth node of the mesh. The above exemplary stream function for a node describes the current loop flowing through all triangle elements of the mesh that share that node on the surface S. Nodes on the edges of the mesh can be forced to have the same stream function value to prevent current from flowing into and out of the edge. In the inverse solution, the stream function value at each node of the mesh serves as a free parameter that can be optimized using the techniques described herein.

The process then proceeds to act 1214, in which the magnetic field B(r) in the region of interest V is discretized. The magnetic field can be discretized by defining a set of target points located within the region V. The target points can have any position in space except on the curved surface S and together define the target region of interest V. In some embodiments described in more detail below, the set of target points can include a first target point corresponding to a first region of desired maximum magnetic field and a second target point corresponding to a second region of desired minimum (e.g., zero) magnetic field. For example, the first target point can be located in a volume inside the curved surface S, while the second target point can be located outside the curved surface S. Including the second target point enables the design of an RF coil that provides shielding benefits in addition to optimizing the coil design to provide a desired magnetic field in the field of view of a low-field MRI system in the region to be imaged, such as a low-field MRI system.

Then, the processing of Figure 12 proceeds to action 1216, in which the model configuration is optimized by determining the optimal value of the current loop on the surface S modeled by the stream function at each node of the grid for the desired magnetic field at the target point set. In addition to the desired magnetic field, some embodiments also include other parameters that are desired to be minimized during optimization, such as the energy stored in the coil (inductance) or the resistive power dissipation in the coil. Boundary conditions can also be imposed as constraints during optimization. For example, in order to conserve the current on the surface S, one or more constraints can be used to impose the same condition for the potential of all points along the edge of the surface. For example, in order to conserve the current of the surface of the head coil shown in Figure 6A, the same condition for the potential of points along a single edge can be forced as a constraint in the optimization. Similarly, the points on the edge of either end of the leg coil surface shown in Figure 6B can also be constrained to have equal potential with other points along the same edge, but the potential along the two edges is allowed to be different. It should be understood that the surface can be formed by any number of separate surfaces, and each of the separate surfaces can have any number of edges. Minimizing the exemplary function U using a suitable optimization scheme can be expressed as follows:

In (3), the first term describes the difference between the measured field and the target field, the second term models the inductance _Lmn , and the third term models the coil resistance _Rmn . The inductance and resistance terms can be weighted using regularization terms α and β determined based on the desired characteristics of the RF coil being designed. In some embodiments, the minimum of the function U can be identified by differentiating the function with respect to _In to produce a system of linear equations that can be combined into a matrix equation: ZI=b, where the matrix Z is calculated by differential optimization and the vector b contains the magnetic field values. This matrix equation can then be inverted to solve for I, which contains the stream function value _In at each node n in the grid. The node stream function values _In can then be linearly combined to reconstruct the stream function of the current density on the surface S. Therefore, the optimization of the surface potential function described above can be used to determine an optimized model configuration, such as the optimized model configurations 705a and 705b shown in Figures 7A and 7B, respectively. However, it should be understood that the above methods are merely exemplary and that any functions and constraints may be optimized to obtain an optimized model configuration and will depend on the nature and characteristics of the model and the requirements of the design.

According to some embodiments, additional constraints can be added to the optimization problem, including but not limited to requiring minimum spacing of wires (e.g., between adjacent turns of a conductor) and/or reducing the total length of the coil conductors (e.g., wire length). As another example, in the case of multi-channel receive coils (e.g., for performing parallel MRI), an additional constraint can be included in the optimization solution to minimize the mutual inductance between a given coil and another coil (e.g., a constraint requiring or attempting to reduce the mutual inductance between pairs of coils to zero or satisfactorily close to zero). Such constraints facilitate the design of multiple receive coil arrays that are substantially decoupled from each other during receive operation.

Introducing additional constraints in the optimization may complicate or compromise the ability to solve the above matrix equation using a simple inversion method. Therefore, some embodiments use convex optimization methods rather than matrix inversion to minimize the function U. For example, the coil design can be optimized by minimizing || _Bψ - _bt || ₂ +α||ψ|| ₂ using Tychonoff regularization of the root mean square (RMS) residual, where _bt is the target field and α is the regularization parameter. In embodiments using convex optimization, any suitable convex optimization solver can be used, as the present invention is not limited in this respect. It should be understood that other optimization methods may also be suitable, including but not limited to gradient descent, genetic algorithms, particle swarm optimization, simulated annealing, Monte Carlo methods, etc.

Returning to the process of FIG. 12 , after determining the optimal solution for the model configuration, processing can proceed to act 1218 , where the stream function of the current density output from act 1216 is used to generate a coil configuration, such as a representation of a conductor profile, that produces the desired magnetic field for the optimized coil design when current is supplied. In some embodiments, a contouring method is used to determine the position of the conductors (e.g., wires) for the optimized coil configuration on the surface S. Contouring can be performed in any suitable manner. For example, each element (e.g., a triangle) of the mesh used to approximate the surface S can be transformed into parameter (u, v) space via a linear transformation. The values of the stream function at the corners of the element (e.g., for the triangle element (ψ ₁ , ψ ₂ , ψ ₃ )) can be used to define a plane of stream functions in the element in (u, v, ψ) space. The intersection of this plane with the constant ψ plane—representing the contour level ψC _n —gives the equation for the conductor path in that element. The portion of these lines that falls within the u and ν limits of the unit element is the line path of that element. This process can be performed on all elements and converted back to (x, y, z) space to obtain the conductor path of the coil configuration. In some embodiments, constraints are added during contouring to constrain the solution based on one or more physical properties, such as the width dimension of the conductor as described above (e.g., the cross-sectional diameter of the wire).

Once the conductor path for the RF coil is known, the process of FIG. 12 proceeds to act 1220, where a support structure for the RF coil is generated and the coil configuration is applied to the support structure. In some embodiments, a three-dimensional (3D) printer or other suitable device can be used to generate the support structure for the optimized RF coil design described above. The support structure can include one or more channels, slots, or conduits corresponding to the positions of the conductor paths derived from the determined RF coil configuration (e.g., the conductor paths derived from contouring the optimized stream function values described above). In other words, the coil configuration can be used to determine the positions of the slots configured to accommodate the coil conductors according to the coil configuration. The support structure can be provided in other ways to facilitate the application of one or more conductors according to the RF coil configuration determined by optimization. Processing then proceeds to act 1222, where one or more conductors (e.g., wires) are positioned along the path on the support structure to create an RF coil based on the optimized configuration. A suitable resonant circuit can then be coupled to the coil to produce an RF coil optimally configured for transmission and/or reception, for example, as part of a low-field MRI system. In particular, the coil can be tuned to resonate at a target frequency in a low field regime.

As described above, in low-field environments, the relatively low transmit frequency allows for conductor lengths to be significantly increased relative to conductor lengths in high-field regimes. For example, the conductor path length shown in the exemplary RF coil configuration applied to the support structure shown in Figures 10A and 10B is approximately 4 meters, which exceeds the maximum length limit in high-field environments by an order of magnitude or more. According to some embodiments, the conductor length is greater than 1 meter, greater than 2 meters, greater than 4 meters, greater than 7 meters, greater than 10 meters, and so on. Thus, transmit/receive coils that operate optimally according to desired standards can be relatively simple and cost-effective to design and produce, and can operate with relatively high efficiency.

In addition to the design flexibility afforded by the increased conductor length, this substantial relaxation of these constraints allows RF coils to be formed using a single conductor wound in multiple turns, using single-strand or multi-strand wire of appropriate gauge, such as Litz wire. For example, the configuration shown in Figures 10A and 10B includes 20 turns or loops of the conductor. However, any number of turns can be selected or determined through optimization and may depend on the coil geometry and its desired operating characteristics. Generally speaking, increasing the number of turns or loops of a coil conductor increases the coil's sensitivity. However, the inventors have recognized that, to some extent, increasing the number of turns can actually degrade the performance of the RF coil. In particular, a coil comprising multiple turns or loops will resonate (self-resonate) without being tuned, at least in part due to parasitic capacitance generated by the relationship between the conductors in the multiple turns or loops of the coil. The effect of self-resonance is to reduce the coil's Q factor and degrade its performance. This effect can be particularly detrimental when the self-resonance is close to the frequency at which the RF coil is tuned to resonate (i.e., the coil's target resonant frequency corresponding to the strength of the MRI system's B0 field). Because the frequency of self-resonance decreases as the number of turns increases, this phenomenon can impose an effective limit on the number of turns of the conductor before coil performance degrades unsatisfactorily. According to some embodiments, the number of turns of the conductor of the coil is limited to ensure that the frequency of self-resonance is at least twice the target resonant frequency to which the RF coil is tuned. According to some embodiments, the number of turns of the conductor of the coil is limited to ensure that the frequency of self-resonance is at least three times the target resonant frequency to which the RF coil is tuned, and according to other embodiments, the number of turns of the conductor of the coil is limited to ensure that the frequency of self-resonance is at least five times the target resonant frequency.

The limit on the number of turns required to ensure that the self-resonant frequency is a desired distance away from the target resonant frequency depends on many factors, including the geometry and size of the coil (for example, the geometry of the head coil may result in different limits compared to the geometry of the leg coil to achieve the same separation of the self-resonant frequency from the target resonant frequency) and the type of conductor used (for example, the gauge of the wire, whether the wire is single-strand or multi-strand, etc.). It should be understood that the limit on the number of turns can be selected to be any number depending on the requirements of the coil, including not limiting the number of turns of the conductor of the coil.

The inventors have developed transmit/receive coil configurations that improve the coil's effectiveness in transmitting RF pulses and/or detecting MR signals emitted in response. As described above, the aforementioned exemplary coils are configured to detect linearly polarized components of MR signals oriented along the coil's primary axis (e.g., axis 1325 shown in FIG. 13A ). However, the circularly polarized MR signals emitted, for example, in the configuration shown in FIG. 13A also include linearly polarized components oriented in an orthogonal direction, shown as axis 1335 (perpendicular to the plane of the figure), which are not detected by the aforementioned exemplary coils. For example, as shown in FIG. 13C and FIG. 13D , the aforementioned exemplary head coil and exemplary leg/knee coil configurations are configured to detect MR signal components oriented along axis 1325, but not along axis 1335. FIG. 13B illustrates a BO magnet 1324 having a cylindrical geometry oriented in the same coordinate system as the planar BO magnet shown in FIG. 13A . For example, the BO magnet can be a solenoid electromagnet that generates a BO magnetic field along axis 1325. Thus, since the primary axis of the coil configuration is aligned with the B0 field, the exemplary coil configurations shown in FIG. 13C and FIG. 13D generally cannot be used in such configurations. The inventors have recognized that the RF coil can be configured to detect MR signal components oriented along axis 1335 and/or axis 1345, and that such configurations can, but are not necessarily, optimized using the same techniques as described above. Thus, the RF coil can be configured to detect MR signals using any B0 magnet geometry (e.g., planar, cylindrical, solenoid, etc.) by configuring the RF coil with a primary axis appropriately oriented relative to the direction of the B0 field.

By way of illustration, Figures 14A and 14B illustrate an exemplary phantom configuration 1405 and an RF coil configuration 1415 determined from the phantom configuration 1405, which are adapted (e.g., optimized) to detect MR signal components oriented along a principal axis 1435 of a head coil, and Figures 15A and 15B illustrate an exemplary phantom configuration 1505 and an RF coil configuration 1515 determined from the phantom configuration 1505, which are adapted (e.g., optimized) to detect MR signal components along a principal axis 1535 of a leg/knee coil. Principal axis 1435 and principal axis 1535 also correspond to exemplary reference axes (it will be appreciated that there are multiple reference axes) about which the respective configurations form multiple turns. As shown, principal axis 1435 and principal axis 1535 are substantially orthogonal to the longitudinal axis of the target body tissue when the target body tissue is located within the respective coils.

As shown in Figures 14B and 15B, the primary axes 1435 and 1535 are orthogonal to the axes 1445 and 1545, respectively, along which the B0 field, such as that generated by a dual-plane B0 magnet, can be oriented. Similarly, as shown in Figures 14B and 15B, the primary axes 1435 and 1535 are orthogonal to the axes 1425 and 1525, respectively, along which the B0 field, such as that generated by a solenoidal B0 magnet, can be oriented. Thus, the coil configurations 1415 and 1515 can be used to generate coils that transmit RF pulses and/or detect MR signals under a variety of B0 magnet geometries. In a manner similar or identical to that described above, the exemplary RF coil configurations 1415 and 1515 can then be applied to a supporting substrate by creating slots or other structures for accommodating the conductors of the coils according to the respective coil configurations (e.g., the exemplary head coil substrate 1650a and leg coil substrate 1650b shown in Figures 16A and 16B, respectively) and placing conductors (e.g., wires) within the slots or otherwise securing the conductors to the substrate in an arrangement described by the coil configurations (e.g., the outlines of the exemplary coil configurations 1415 and 1515, respectively), thereby forming a plurality of turns about exemplary reference axes 1635a and 1635b, which also correspond to the main axes of the respective coils.

FIG17 illustrates an exemplary head coil for, for example, obtaining one or more images of a patient's brain, in which conductors are placed within slots formed in a support base in the form of a helmet configured to accommodate a person's head. Specifically, head coil 1700 includes a base 1750 having slots or channels 1780 arranged according to a desired coil configuration, within which conductors 1725 are placed to form multiple turns or loops of the coil (e.g., turn 1727 shown). Slots 1785 are arranged to connect slots 1780, allowing conductors 1725 to be wound around the support base from one contour or loop to the next according to the desired coil configuration. Exemplary head coil 1700 includes 20 turns (10 turns per hemisphere) formed by the coil configuration's conductor loops around a main axis and an exemplary reference axis 1735. As described above, relatively long conductor lengths can be used in low-field environments, enabling a single conductor to be wound around a curved surface of interest according to the desired coil configuration. It should be understood that, in some embodiments, coil configurations can be implemented using multiple conductors that may be independent of one another or connected together. According to some embodiments, conductor 1725 is formed of a suitable gauge wire. For example, conductor 1725 can be a single strand wire or can be a multi-strand wire such as Litz wire. It should be understood that conductor 1725 can be any suitable conductor, as the present invention is not limited to use with any particular type of conductor.

As previously discussed, the inventors have recognized that multiple coil configurations can be used in combination to improve the SNR of RF coils. For example, a pair of coils configured with different principal axes can be used to obtain dual measurements of MR signals. According to some embodiments, the RF transmit/receive component is configured to include a first coil and a second coil configured with orthogonal or substantially orthogonal principal axes, respectively, to improve the SNR of the RF component. For example, an exemplary head coil configuration suitable for detecting MR signal components oriented along the principal axis 1325 of the exemplary coil configuration shown in FIG13B and an exemplary head coil configuration suitable for detecting MR signal components oriented along the principal axis 1435 of the exemplary coil configuration shown in FIG14B can be used together to detect MR signal components oriented along both principal axes. As discussed in further detail below, by using such a dual coil arrangement, the SNR of MR signal detection can be improved.

As examples, Figures 18A and 18B illustrate coil configurations that can be combined to provide a head coil capable of detecting MR signal components oriented along multiple axes, according to some embodiments. Specifically, Figure 18A illustrates an exemplary coil configuration 1815a arranged to detect MR signal components oriented substantially along a principal axis 1825, and Figure 18B illustrates an exemplary coil configuration 1815b arranged to detect MR signal components oriented substantially along a principal axis 1835 that is orthogonal to principal axis 1825. Figure 18C illustrates a multi-coil configuration 1815c resulting from combining coil configurations 1815a and 1815b, which is arranged to detect MR signal components oriented substantially along principal axes 1825 and 1835.

As another example, Figures 19A and 19B illustrate exemplary coil configurations 1915a and 1915b, respectively, configured to detect MR signal components oriented along orthogonal principal axes 1925 and 1935. Exemplary coil configurations 1915a and 1915b can be combined to form coil configuration 1915c, shown in Figure 19C, thereby providing a multi-coil configuration arranged to detect MR signal components oriented along multiple orthogonal axes. By configuring multiple coils to detect MR signal components oriented along substantially orthogonal axes, inductive coupling between coils can be optimally avoided. According to some embodiments, using dual coils configured with orthogonal principal axes can improve the SNR of MR signal detection by a square root of two. Specifically, each of the dual coils can obtain independent measurements of the same MR signal offset by 90°, resulting in an SNR improvement of a square root of two.

In the example coil configurations shown in Figures 18C and 19C, the dual coil configurations are oriented substantially orthogonal to each other and to the B0 field. That is, the primary axes of the dual coils are orthogonal to each other and to axes 1845, 1945 aligned with the B0 field. However, the inventors recognize that other arrangements may also be used. For example, Figure 20 shows a combined coil configuration 2015c for an exemplary head coil. Combined coil configuration 2015c includes: coil configuration 2015a having conductors arranged to detect MR signal components generally oriented along axis 2025; and coil configuration 2015b having conductors arranged to detect MR signal components generally oriented along axis 2035. In the example configuration shown in Figure 20, axes 2025 and 2035 are orthogonal to each other and oriented at 45° relative to axes 2045a and 2045b. A possible B0 field can be generated in the direction of axes 2045a and 2045b, for example, by a low-field MRI device. It should be understood that other configurations are possible, as the present invention is not limited in this respect. For example, multiple coils can be configured to detect MR signals in non-orthogonal directions. However, in this case, care should be taken to produce a coil configuration with suitably low mutual inductance. As discussed in further detail below, the inventors have recognized that the optimization techniques described herein can be used to determine a coil configuration that minimizes mutual inductance between the coils. In this manner, multiple coils that are not in an orthogonal relationship with each other can be used.

To implement multiple coil configurations (e.g., exemplary coil configurations 1815c, 1915c, 2015c, etc.) to provide an RF transmit/receive component comprising multiple coils (e.g., a pair of orthogonal coils), the inventors recognized that the conductors forming the coils of the respective configurations can be offset from one another. To separate a pair of coils disposed around a region of interest, the conductors of the coils can be offset from one another relative to the region of interest. For example, the conductors of a first coil can be disposed around the region of interest, and the conductors of a second coil can be disposed around the region of interest at a distance further from the region of interest. According to some embodiments, a support structure includes: an inner substrate layer having a curved surface around the region of interest to which the first coil is applied; and an outer substrate layer having a curved surface around the region of interest to which the second coil is applied. The inner and outer substrate layers can be offset from one another, for example, along a direction normal to the curved substrate surface to which the coils are applied. In this regard, the outer substrate layer is disposed above the inner substrate layer relative to the region of interest. Some non-limiting examples of dual-coil RF components having a first coil disposed in a first substrate layer of a support structure and a second coil disposed in a second substrate layer of the support structure offset from the first substrate layer are described in further detail below. However, it should be understood that the plurality of coils may be applied in other manners, as the present invention is not limited in this regard.

As an example of a dual-coil RF component, FIG21 shows a helmet 2100 in which a pair of coil configurations are applied to corresponding base layers of the helmet. Specifically, coil configuration 2115a (e.g., a coil configuration similar or identical to coil configuration 1815a shown in FIG18A ) is applied to the outer base layer 2155a of the support structure of the helmet 2100 via slots, which are suitable for accommodating coil conductors arranged according to the corresponding coil configurations. Base layer 2155a is shown in FIG21 , wherein one of the hemispheres is removed to show the inner base layer below. In this regard, coil configuration 2115b (e.g., a coil configuration similar or identical to coil configuration 1815b shown in FIG18B ) is applied to the inner base layer 2155b of the support structure of the helmet 2100 via slots, which are suitable for accommodating coil conductors arranged according to the corresponding coil configurations. As shown in FIG21 , the direction in which the outer base layer 2155a is offset from the inner base layer 2155b is perpendicular to the base curve. In this illustrative example, the outer base layer 2155a overlaps the inner base layer 2155b.

As shown in exemplary helmet 2100, the inner and outer base layers form corresponding curved surfaces around a region of interest within the helmet. When the helmet is worn by a patient and operated within the appropriate BO field of an MRI system, the region of interest will include the MRI system's field of view (i.e., an area with sufficient homogeneity to perform MRI within the BO field). Accordingly, exemplary base layers 2155a and 2155b shown in FIG21 are offset relative to the region of interest, with outer base layer 2155a positioned further from the region of interest than inner base layer 2155b. Consequently, when operating within the appropriate BO field of an MRI system, coils applied to outer base layer 2155a will be further from the field of view than coils applied to inner base layer 2155b. When a conductor is placed within the slots of base layer 2155a, the conductor forms multiple turns around a principal axis 2125 (e.g., aligned with the longitudinal axis of the body), and when a conductor is placed within the slots of base layer 2155b, the conductor forms multiple turns around a principal axis 2135 (e.g., substantially orthogonal to the longitudinal axis of the body).

When arranged in a close proximity, coils arranged in separate layers may exhibit capacitive coupling. Capacitive coupling between coils arranged in separate layers can be reduced or avoided by increasing the distance between coils in different layers along the normal direction of the curved surface of the support structure. For example, by increasing the offset of the coils in the outer layer along the normal direction of the curved surface, capacitive coupling can be reduced or eliminated. However, the increased offset will generally also reduce the sensitivity of the coils in the outer layer due to the increased distance from the area of interest, so the offset can be appropriately and/or desired to appropriately balance capacitive coupling and coil sensitivity. Alternatively or in addition, a decoupling network can be included to reduce or eliminate capacitive coupling between coils arranged in separate layers.

In FIG21 , an opening or slot 2175 can be provided to facilitate connection to the hemisphere of the outer layer 2155a and/or to accommodate the end of the coil conductor, which, when located within the slot, facilitates connection to the transmit and/or receive circuitry that operates the RF head coil. In this manner, multiple coil configurations can be applied to the support structure to produce an RF head coil with an improved SNR.

Figures 22A and 22B illustrate an alternative technique for applying multiple coil configurations to the support structure of helmet 2200. In Figure 22A, coil configuration 2215a (e.g., similar or identical to coil configuration 1815a shown in Figure 18A) is applied to an inner substrate layer 2255a of the support structure of helmet 2200 via slots adapted to accommodate coils and position conductors according to the respective coil configurations. Figure 22B shows a hemisphere of outer substrate layer 2255b removed to illustrate inner substrate layer 2255a of Figure 22A and coil configuration 2215b applied to the inner surface of outer substrate layer 2255b. Specifically, coil configuration 2215b (e.g., similar or identical to coil configuration 1815b shown in Figure 18B) is applied to the inner side (e.g., on the concave side) of outer layer 2255b via slots adapted to accommodate and position coil conductors according to the respective coil configurations. The opening 2275 is configured to accommodate conductor terminals for connecting to the transmitting and/or receiving circuit system, and can also be adapted to attach two portions of the outer layer 2255b. It should be understood that the coil configuration can be applied to the inner or outer layers in Figures 21 and 22, and that the arrangement shown is selected for illustration only. In addition, it should be understood that the coil configuration can be applied to the concave or convex side of the inner or outer substrate layer, and the arrangement shown is shown to illustrate that the coil configuration can be applied to either side of the substrate curve.

FIG23A and FIG23B illustrate an RF head coil 2300, which includes a first RF coil 2310a formed from conductors 2327a arranged according to a first configuration (e.g., by placing conductors 2327a within slots patterned in an inner layer according to configuration 2215a shown in FIG22 ); and a second RF coil 2310b formed from conductors 2327b arranged according to a second configuration (e.g., by placing conductors 2327b within slots patterned in an outer layer according to configuration 2215b shown in FIG22 ), as shown in FIG23A . FIG23B shows the distal ends of conductors 2327a and 2327b emerging from openings in the support structure of the head coil 2300 for connection to transmit and/or receive circuitry, enabling the RF head coil to be operated, for example, to obtain one or more MRI images (e.g., one or more images of a patient's brain). For example, the RF head coil 2300 can be connected to a low-field MRI system to obtain MR signals with an improved signal-to-noise ratio (SNR), thereby improving the quality of the acquired images.

It should be understood that providing conductors for an RF coil by arranging slots, channels, or conduits in a desired configuration is merely one example of creating an RF coil, which may be suitable, for example, when using 3D printing or similar techniques to create the support structure. However, any method or technique may be used to arrange the conductors in a desired configuration to create the RF coil. For example, one or more conductors may be encapsulated within the support structure material using a molding process or other manufacturing process, or one or more conductors may be secured to the support structure in other ways, such as by fasteners, adhesives, etc. Any suitable technique for arranging the conductors in a desired configuration may be used, as the present invention is not limited in this respect.

FIG24 shows a support structure 2400 for a leg coil employing paired coil configurations. Specifically, coil configuration 2415a (e.g., a coil configuration similar or identical to coil configuration 1915a shown in FIG19A ) is applied to an outer layer 2455a of support structure 2400 via slots adapted to accommodate and secure the coil conductor position according to the respective coil configuration. Coil configuration 2415b (e.g., a coil configuration similar or identical to coil configuration 1915b shown in FIG19B ) is applied to an inner layer 2455b of support structure 2400 via slots adapted to accommodate and secure the coil conductor position according to the respective coil configuration. Structure 2475 provides a mechanism for arranging the ends of the conductors, which, once placed within the slots of the coil configuration, are connected to the transmit and/or receive circuitry to operate the RF coil. In this manner, multiple coil configurations can be applied to a support structure to produce an RF leg coil with an improved SNR.

FIG25 illustrates an exemplary RF coil 2500 suitable for use with a leg. RF coil 2500 includes a first RF coil 2510a formed from a conductor 2527a arranged in a first configuration (e.g., by placing conductor 2527a in an outer layer according to configuration 1915a shown in FIG19A ) forming a plurality of turns about an exemplary reference axis 2525 (e.g., a principal axis substantially aligned with the longitudinal axis of the leg within the coil); and a second RF coil 2510b formed from a conductor 2527b arranged in a second configuration (e.g., by placing conductor 2527b in an inner layer according to configuration 1915b shown in FIG19B ) forming a plurality of turns about an exemplary reference axis 2535 (e.g., a principal axis substantially orthogonal to the longitudinal axis of the leg within the coil). RF coil 2500 can be used as part of a low-field MRI system to obtain one or more images of a portion of the leg, such as a knee. Connector 2575 routes the ends of conductors 2527a and 2527b and provides connections for electrically connecting the conductors to, for example, transmit and/or receive circuitry of a low-field MRI system. It will be appreciated that the techniques described above can be used to create RF coils for any portion of body tissue, and that the exemplary head and leg coils depicted are merely examples for illustrating the methods and apparatus developed by the inventors and discussed herein.

In embodiments having a radio frequency component comprising multiple coils, one or both of the coils can be used to transmit RF pulses to a region of interest to induce an MR response. For example, in some embodiments, only one of the multiple coils is used as a transmit coil, and each of the multiple coils is used as a receive coil. According to some embodiments, each of the multiple coils is used as both a transmit coil and a receive coil. Thus, the multiple coils can be used in any arrangement to provide the transmit/receive component of a magnetic resonance imaging system, such as a low-field MRI system.

In embodiments that include multiple coils (e.g., an RF transmit/receive assembly using pairs of coils in a quadrature relationship, as illustrated by the exemplary RF coils 2300 and 2500 shown in Figures 23 and 25), MR signals will generate electrical signals in each of the multiple coils. These signals can be combined in any number of ways to improve the SNR. For example, the electrical signals can be combined in the analog or digital domain. In the analog domain, the electrical signals generated in each of the multiple coils can be appropriately phase-shifted and combined. For example, using the exemplary coils described above, the electrical signals generated in each pair of coils based on the corresponding MR signals will have a 90° phase difference due to the orthogonality of the respective configurations. Therefore, the electrical signal from one coil can be phase-shifted by 90° and combined with the electrical signal generated by the other coil to produce a combined signal with an increased SNR. In the digital domain, MR signals can be acquired via separate channels (e.g., a separate signal can be acquired from each coil) and digitized. The digitized signals can then be digitally processed and combined in the digital domain by phase-shifting them. One advantage of obtaining separate signals and processing them in the digital domain is that the separate signals can be noise corrected before being combined.However, MR signal components detected by multiple coils may be combined and processed in any suitable manner, as the present invention is not limited in this respect.

As described above, the aforementioned techniques can be used to optimize the coil configuration of a radio frequency component comprising a plurality of coils, for example using magnetic synthesis to determine a coil configuration that is generally optimal with respect to one or more parameters. According to some embodiments, the mutual inductance between the plurality of coils can be included as a term in the optimization scheme for minimizing the mutual inductance between the coils. The mutual inductance term may be particularly beneficial in the following embodiments, where the coil configurations are not orthogonally oriented to one another (e.g., coil configurations having principal axes that are not orthogonal to one another), either due to design or because the orthogonality cannot be achieved to the desired degree. Minimizing (or eliminating) the mutual inductance between the coils facilitates the radio frequency component to have an improved SNR and/or sensitivity, thereby improving the quality of MR signal detection.

The low-field MRI system may include RF components provided according to any one or a combination of the aforementioned techniques to facilitate acquisition of clinically useful images at low fields. For example, the low-field MRI system may include: a B0 magnet 122 configured to generate a low-field B0 magnetic field; and a transmit/receive element 125 that may be optimized using any one or a combination of the techniques described herein to increase sensitivity and/or configured to improve the SNR of MR signal detection to facilitate acquisition of clinically useful images of one or more desired portions of body tissue.

The inventors also recognized that coils can be operated so that they generate more than one type of magnetic field in an MRI system. For example, the inventors have developed a system that drives one or more coils with the multifunctional capability to generate one or more gradient magnetic fields and to generate and/or receive one or more RF magnetic fields. According to some embodiments, the multifunctional coil is configured to operate as at least one transmit/receive coil and at least one gradient coil. The inventors also recognized that the configuration of such a multifunctional coil can be optimized using the optimization techniques described herein. Further details on the design and optimization of the multifunctional coil are provided below.

FIG26A illustrates a system configured to generate a multifunctional coil operable to generate multiple types of magnetic fields, according to some embodiments. The exemplary system schematically depicted in FIG26A includes a controller 2675 coupled to a coil 2600 to cause the coil to generate at least a gradient magnetic field and an RF magnetic field. According to some embodiments, the controller 2675 includes a gradient amplifier 2620 coupled to the coil 2600 via a low-pass filter 2630. In operation, a console 2685 can issue a gradient command input 2610 to cause the gradient amplifier 2620 to drive the coil 2600 to generate one or more gradient fields according to a desired pulse sequence (e.g., a pulse sequence designed to acquire MR data for generating one or more images). In this manner, the coil 2600 can operate as a gradient coil (e.g., Gx, Gy, etc.), for example, in a low-field MRI system.

Controller 2675 also includes an RF amplifier 2650, which is coupled to coil 2600 via a high-pass filter 2640. Console 2685 can also issue RF command input 2660 to cause RF amplifier 2650 to drive coil 2600 to generate an RF magnetic field according to a desired pulse sequence. By doing so, coil 2600 can also operate as an RF coil. According to some embodiments, controller 2675 can also use coil 2600 to detect MR signals emitted in response to the RF magnetic field generated by coil 2600, allowing coil 2600 to operate as both a transmit coil and an RF receive coil. For example, FIG26B illustrates a multifunction coil 2600 driven by controller 2675, wherein controller 2675 has both a transmit path 2680 and a receive path 2690, enabling the use of multifunction coil 2600 as both a transmit coil and a receive coil. The T/R switch 2687 switches between the transmit path 2680 and the receive path 2690 to allow selective operation of the multi-function coil to generate an RF magnetic field and detect MR signals transmitted in response to an RF transmit cycle.

It should be understood that coil 2600 can be used as an RF receive coil, either when operating as an RF transmit coil or when not operating as an RF transmit coil, and further, coil 2600 can be used as an RF transmit coil, either when operating as an RF receive coil or when not operating as an RF receive coil. Thus, controller 2675 is configured to operate coil 2600 as both a gradient coil and an RF coil, enabling coil 2600 to provide multiple functions in an MRI system, such as a low-field MRI system. It should be understood that the controllers shown in FIG. 26A and FIG. 26B are merely exemplary and may include additional components and/or may exclude one or more of the components shown, as a suitable controller for implementing a multifunctional coil may include any combination of components configured to cause the coil to generate multiple types of magnetic fields.

According to some embodiments, a multifunction coil (e.g., coil 2600) operates as both a Gx gradient coil and an RF transmit/receive coil. According to some embodiments, a multifunction coil operates as both a Gy gradient coil and an RF transmit/receive coil. It should be understood that more than one multifunction coil can be used in an MRI system. For example, according to some embodiments, a first multifunction coil is configured to operate as a Gx gradient coil and a second multifunction coil is configured to operate as a Gy gradient coil, where both the first and second multifunction coils also operate as RF transmit/receive coils. Multiple multifunction coils operated in this manner can be used to implement multiple transmit/receive channels, which can be used to improve signal-to-noise ratio (SNR), reduce acquisition time, or both. For example, MR data acquired from multiple receive coils can be combined to increase the SNR. When both the Gx gradient coil and the Gy gradient coil also function as receive coils, there is a 90-degree phase difference between the receive channels (i.e., because the Gx gradient coil and the Gy gradient coil are substantially orthogonal to each other and to the B0 magnetic field). This orthogonality can be exploited to improve the SNR by as much as the square root of two. Alternatively, or in addition to increasing the SNR, multiple transmit/receive coils may be used to perform parallel MR to reduce the acquisition time required to obtain MR data for generating one or more images.

FIG27 illustrates a system for providing a multifunctional coil system in conjunction with a specific configuration of a gradient coil assembly. It should be understood that while the gradient coil assembly shown in FIG27 is labeled as a Gx gradient coil assembly, this is not limiting, as the same techniques can be applied to a Gy gradient coil assembly. In FIG27 , the gradient coil assembly is configured as coil pairs, with each pair having coils connected with opposite polarity or using a 180-degree coaxial line phase shifter circuit so that they are driven 180 degrees out of phase. In FIG27 , an exemplary controller 2775 is configured to transmit and/or receive RF magnetic fields using the gradient coil assembly, which also operates as an RF coil. According to some embodiments, the gradient coil assembly operates as a single RF coil. One technique for achieving this is to treat the gradient coil assembly as a single continuous coil by coupling the corresponding balun to a 1:4 RF splitter/combiner in addition to each high-pass filter. In this manner, the gradient coil assembly configured in FIG27 can also be driven as a transmit and/or receive coil. Alternatively, each coil in the gradient coil set may be treated individually from an RF perspective by driving each coil using a respective RF amplifier and high pass filter, so that the gradient coil set may operate as four separate transmit and/or receive coils.

The inventors have recognized that multifunctional coil technology can facilitate low-field MRI systems with reduced cost and/or size. For example, the technology described herein can be applied to the low-field MRI systems for imaging the head shown in Figures 22A-C of the '652 application. These systems include a head component (e.g., a helmet) configured to accommodate the head of the person being imaged. The head component can incorporate one or more coils of a low-field MRI system (e.g., a B0 magnet, one or more gradient coils, one or more transmit/receive coils, etc.). The inventors have recognized that the head imaging system shown can be produced by configuring at least one coil that generates at least two types of magnetic fields to be incorporated into or accommodated in the head component (i.e., the head component can accommodate one or more multifunctional coils). According to some embodiments, the head component includes coils configured to transmit and/or receive RF magnetic fields and generate at least one gradient magnetic field. As previously described, such a multifunctional coil can be implemented by coupling a controller to the multifunctional coil to operate the coil as both an RF coil and a gradient coil (e.g., by coupling a first amplifier and a high-pass filter to the coil to drive the coil to generate and/or receive an RF magnetic field, and coupling a second amplifier and a low-pass filter to the coil to drive the coil to generate at least one gradient magnetic field). In this manner, one or more multifunctional coils can be used to generate both transmit RF pulses and gradient pulses according to a desired pulse sequence, and detect MR signals transmitted in response.

By using the above-described techniques to implement a multifunction coil, the cost of the resulting system can be reduced because a single coil can be used to generate more than one type of magnetic field for an MRI system. Furthermore, the multifunction coil can reduce the system's footprint and/or facilitate designs where space available for incorporating magnetic components into the system is limited (e.g., in the head imaging system described above). Another benefit of some of the above-described embodiments relates to the ability to implement multiple transmit and/or receive channels using the MRI system's gradient coils.

The inventors have recognized that the optimization techniques described herein can be used to optimize the configuration of multifunction coils overall. As described above, the optimization can be formulated to determine a coil configuration that satisfies one or more constraints and, when simulated, produces a magnetic field that meets one or more criteria. By formulating the optimization to include regularization terms for both the gradient coils and the RF coils, a coil configuration can be determined that produces both gradient and RF magnetic fields that meet the specified criteria. Therefore, the optimization techniques described herein can be applied to generate both single-function and multifunction coils, among others.

U.S. Patent Application No. 14/845,949, filed on September 4, 2015, entitled “Noise Suppression Methods and Apparatus,” (the ‘949 application), describes, among other subject matter, techniques for using auxiliary sensors to facilitate characterizing the noise environment of a low-field MRI system, thereby suppressing noise received by one or more RF receive coils. The entire contents of the ‘949 application are incorporated herein by reference. The techniques described in the ‘949 application allow magnetic resonance imaging systems (e.g., low-field MRI systems) to be operated outside a shielded room. This facilitates the creation of MRI systems that can operate in any environment, enabling MRI to be used in many scenarios that conventional MRI cannot meet. Any of the noise cancellation techniques described in the ‘949 application can be used in conjunction with the coil configurations described herein. Furthermore, the inventors recognize that the optimization techniques described herein can also be applied to determine the optimal configuration of one or more auxiliary sensors (e.g., auxiliary coils) for noise suppression. In particular, one or more criteria and/or one or more constraints corresponding to the desired operation of the auxiliary coils can be incorporated into the aforementioned optimization scheme to determine the coil configuration for the auxiliary coils. As further discussed in the '949 application, some embodiments include using an RF coil as both an auxiliary coil and as a primary coil, and in this regard, represents another example of a multifunction coil. The optimization techniques described herein can also be used to determine the configuration of a multifunction coil that is configured to operate as both a primary coil and an auxiliary coil, or additionally as a gradient coil.

As described above, the optimization techniques described herein can be used to optimize the configuration of a head coil disposed on a curved surface of a helmet adapted to accommodate a patient's head. The inventors have recognized that one or more auxiliary coils can be placed on or near the helmet to facilitate noise suppression. For example, the head coil can be configured to optimally detect MR signals emitted from a patient in a field of view within the helmet. One or more auxiliary coils can be placed near (or on) the helmet so that they respond to the noisy environment and not to MR signals emitted from the field of view. Noise signals from the one or more auxiliary coils can be used to suppress noise in the MR signals detected by the head coil, for example, using any of the techniques described in the '949 application.

As described above, and described in detail in the '949 application, one or more auxiliary coils can be used to detect a noisy environment without detecting MR signals emitted from the field of view of an MRI system. This is typically achieved by placing one or more auxiliary coils close to a primary coil (e.g., the primary receive coil of the MRI system) such that the auxiliary coils respond to a noise environment as similar as possible to that of the primary coil, but are positioned outside the detection range of the transmitted MR signals, such that the auxiliary coils do not respond to the transmitted MR signals. In this manner, the one or more auxiliary coils characterize a noise environment substantially similar to that of the primary coil but do not respond to MR signals, allowing the noise environment characterized by the one or more auxiliary coils to be used to suppress the noise detected by the primary coil. However, when positioned close to one another in this manner, the primary and auxiliary coils may become inductively coupled, causing one or more auxiliary coils to respond to MR signals emitted from the field of view due to inductive coupling with the primary coil, even though the auxiliary coils are outside the range of the MR signals. Because the auxiliary coil response also includes MR signal content, the described noise suppression techniques will serve to suppress the MR signal content detected by the primary coil, thereby reducing the signal-to-noise ratio (SNR), rather than increasing it as might be expected.

The inventors have recognized that the optimization techniques described herein can be used to generate configurations of auxiliary coils that reduce or eliminate inductive coupling with the main coil. Using this technique, auxiliary coils can be positioned close to the main coil while avoiding undesirable inductive coupling. According to some embodiments, the configuration of one or more auxiliary coils can be optimized to reduce or eliminate inductive coupling with the main coil. For example, the optimization scheme can include: one or more terms that define the region in which the auxiliary coil is sensitive to noise, which excludes the region in which MR signals can be directly detected; and one or more terms for minimizing inductive coupling between the one or more auxiliary coils and the main coil (e.g., one or more terms that cause the resulting configuration to suppress or eliminate mutual inductance between the coils when operating in conjunction with the main coil). According to some embodiments, the configuration of the main coil and one or more auxiliary coils can be optimized together so that the resulting main coil has generally optimal performance relative to a specified standard for operation of the receive coil, and the resulting one or more auxiliary coils operate with minimal or no inductive coupling with the main coil.

It should be understood that the techniques described herein can be applied to determine a coil configuration optimized for any portion of human body tissue, and the head coil shown is merely exemplary. In particular, the optimization techniques described herein are not limited with respect to the specific surface for which the coil configuration is optimized. Therefore, the techniques described herein can be applied to any surface that can be modeled. For example, by modeling surfaces using triangular meshes, virtually any surface can be triangulated, thus eliminating any substantial limitations on the geometry of RF coils to which these techniques can be applied. Therefore, the techniques described herein can be used to determine the configuration of an RF coil for any portion of body tissue, including but not limited to a head coil, coils for the torso, arms, legs, hands, feet, etc., or any combination thereof. Furthermore, the optimization techniques can be applied to any combination of multifunctional coils for any desired portion of body tissue.

Having described several aspects and embodiments of the technology set forth in this disclosure, it will be understood that various variations, modifications, and improvements will readily occur to those skilled in the art. Such variations, modifications, and improvements are intended to be within the spirit and scope of the technology described herein. For example, a person of ordinary skill in the art will readily conceive of various other devices and/or structures for performing the functions and/or obtaining the results and/or one or more advantages described herein, and each such variation and/or modification is considered to be within the scope of the embodiments described herein. Those skilled in the art will recognize or be able to determine, using only routine experimentation, many equivalents to the specific embodiments described herein. Therefore, it will be understood that the foregoing embodiments are given by way of example only, and within the scope of the appended claims and their equivalents, embodiments of the present invention may be implemented in a manner different from that specifically described. In addition, any combination of two or more features, systems, products, materials, equipment, and/or methods described herein is included within the scope of this disclosure if such features, systems, products, materials, equipment, and/or methods do not contradict each other.

The above-mentioned embodiments can be realized in any of many ways. One or more aspects and embodiments of the execution of the process or method of the present disclosure can utilize program instructions that can be executed by a device (such as a computer, processor or other device) to execute or control the execution of the process or method. In this regard, various inventive concepts can be embodied as a computer-readable storage medium (or multiple computer-readable storage media) (such as a computer memory, one or more floppy disks, compact disks, optical disks, tapes, flash memories, circuit configurations in field programmable gate arrays or other semiconductor devices, or other tangible computer storage media) encoded with one or more programs, and the one or more programs execute the method for realizing one or more of the above-mentioned various embodiments when executed on one or more computers or other processors. The computer-readable medium can be portable so that the program stored thereon can be loaded onto one or more different computers or other processors to realize various aspects in the above-mentioned aspects. In some embodiments, the computer-readable medium can be a non-transient medium.

The terms "program" or "software" are used herein in a general sense to refer to any type of computer code or set of computer-executable instructions that can be used to program a computer or other processor to implement the various aspects as described above. Additionally, it should be understood that, according to one aspect, one or more computer programs that, when executed, perform the methods of the present disclosure need not reside on a single computer or processor, but rather can be distributed in a modular manner among multiple different computers or processors to implement various aspects of the present disclosure.

Computer-executable instructions can take many forms, such as program modules, that are executed by one or more computers or other devices. Typically, program modules include routines, programs, objects, components, data structures, etc. that perform specific tasks or implement specific abstract data types. In general, in various implementations, the functionality of program modules can be combined or distributed as needed.

Furthermore, the data structure can be stored in a computer-readable medium in any suitable form. To simplify the description, the data structure can be shown as having fields that are related by position in the data structure. Similarly, such relationships can be achieved by allocating memory for the fields in the computer-readable medium with positions that convey the relationship between the fields. However, any suitable mechanism can be used (including by using pointers, tags, or other mechanisms that establish relationships between data elements) to establish the relationship between the information in the fields of the data structure.

The above-mentioned embodiments of the present invention can be implemented in any of a variety of ways. For example, hardware, software or a combination thereof can be used to implement the embodiments. When implemented in software, the software code can be executed on any suitable processor or processor set, whether it is arranged in a single computer or distributed between multiple computers. It should be understood that any component or component set that performs the above-mentioned functions can generally be considered to be a controller that controls the functions discussed above. The controller can be implemented in a variety of ways, such as using dedicated hardware or using microcode or software to program the general hardware (for example, one or more processors) that performs the above-mentioned functions, and can be implemented in a combined manner when the controller corresponds to multiple components of the system.

Furthermore, it should be understood that a computer may be implemented in any of a variety of forms, such as, by way of non-limiting example, a rack-mounted computer, a desktop computer, a laptop computer, or a tablet computer. Additionally, a computer may be embedded in a device not generally considered a computer but having suitable processing capabilities, including a personal digital assistant (PDA), a smart phone, or any other suitable portable or fixed electronic device.

In addition, the computer can have one or more input and output devices. These devices can be used to present a user interface, among other things. Examples of output devices that can be used to provide a user interface include: a printer or display screen for visually presenting output, and a speaker or other sound generating device for audibly presenting output. Examples of input devices that can be used for a user interface include a keyboard and a pointing device, such as a mouse, a touchpad, and a digitizing tablet computer. As another example, the computer can receive input information through speech recognition or in other audible formats.

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

Furthermore, as described, some aspects may be embodied as one or more methods. The actions performed as part of a method may be ordered in any suitable manner. Thus, embodiments may be configured that perform actions in an order different from that shown, which may include performing some actions simultaneously, even though shown as sequential actions in illustrative embodiments.

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.

Unless clearly indicated to the contrary, the indefinite articles "a" and "an" as used herein in the specification and claims should be understood to mean "at least one."

As used herein in the specification and claims, the phrase "and/or" should be understood to refer to "either or both" of the elements so combined, i.e., elements that are present in combination in some cases and separately in other cases. Multiple elements listed with "and/or" should be interpreted in the same manner, i.e., "one or more" of the elements so combined. In addition to the elements specifically identified by the "and/or" clause, other elements may optionally be present, whether related or unrelated to those specifically identified. Thus, as a non-limiting example, when used in conjunction with open language such as "comprising," a reference to "A and/or B" may refer, in one embodiment, to only A (optionally including elements in addition to B); in another embodiment, to only B (optionally including elements in addition to A); in another embodiment, to both A and B (optionally including other elements), etc.

As used herein in the specification and claims, the phrase "at least one" with respect to a list of one or more elements should be understood to mean at least one element selected from any one or more elements in the list of elements, but not necessarily including at least one of each element specifically listed in the list of elements, and not excluding any combination of elements in the list of elements. This definition also allows for elements to be optionally present other than the elements specifically identified in the list of elements to which the phrase "at least one" refers, whether related or unrelated to those specifically identified elements. 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") may, in one embodiment, mean: at least one, optionally including more than one, A, with no B present (and optionally including elements other than B); in another embodiment, mean: at least one, optionally including more than one, B, with no A present (and optionally including elements other than A); in yet another embodiment, mean: at least one, optionally including more than one, A and at least one, optionally including more than one, B (and optionally including other elements), etc.

Furthermore, the phraseology and terminology used herein are for the purpose of description and should not be regarded as limiting. As used herein, the words "including," "comprising," or "having," "containing," "involving," and variations thereof are meant to encompass the items listed thereafter and equivalents thereof as well as additional items.

In the claims and throughout the foregoing description, all transitional phrases such as "comprising," "including," "carrying," "having," "containing," "involving," "holding," "composed of," and the like are to be construed as open-ended, i.e., meaning including, but not limited to, including. Only the transitional phrases "consisting of" and "consisting essentially of" are closed or semi-closed transitional phrases, respectively.

Claims (112)

1. A radio frequency component for magnetic resonance imaging, the radio frequency component comprising: A first coil, configured to generate a magnetic field component, includes a first conductor having a length of at least 1 meter and arranged in a plurality of turns, wherein the spacing between the plurality of turns is non-uniform. A second coil, comprising a second conductor arranged in a plurality of turns, is oriented in response to a magnetic resonance signal component, wherein the first and second coils are tuned to resonate at frequencies corresponding to a B0 field less than or equal to 0.2 T and greater than or equal to 20 mT; and A support structure for the first coil and the second coil, the support structure comprising: a first base layer to which the first coil is disposed; and a second base layer to which the second coil is disposed, wherein the support structure includes a helmet formed to accommodate a patient's head. 2. The radio frequency component of claim 1, wherein the first coil is oriented in response to a first magnetic resonance signal component oriented along a first principal axis, and wherein the second coil is oriented in response to a magnetic resonance signal component including a second magnetic resonance signal component oriented along a second principal axis different from the first principal axis. 3. The radio frequency component according to claim 2, wherein the first main axis and the second main axis are substantially orthogonal, such that the first coil and the second coil respond to substantially orthogonal magnetic resonance signal components. 4. The radio frequency component of claim 3, wherein the first coil and the second coil are configured to detect magnetic resonance signals emitted from a field of view of a B0 field oriented along a third axis. 5. The radio frequency component according to claim 4, wherein the first main axis and the second main axis are substantially orthogonal to the third axis. 6. The radio frequency component of claim 4, wherein the first main axis and the second main axis are oriented at approximately 45° relative to the third axis. 7. The radio frequency component of claim 1, wherein the support structure defines a three-dimensional surface, and wherein the second base layer is offset from the first base layer substantially along the surface normal direction of the three-dimensional surface. 8. The radio frequency component of claim 1, wherein the support structure defines a three-dimensional surface around the region of interest, and wherein the second substrate layer is offset from the first substrate layer relative to the region of interest. 9. The radio frequency component of claim 1, wherein the first substrate layer includes at least one slot in which the first conductor is placed, and wherein the second substrate layer includes at least one slot in which the second conductor is placed. 10. The radio frequency component according to claim 2, wherein the first main axis is substantially aligned with the longitudinal axis of the patient's body. 11. The radio frequency component of claim 1, wherein the second conductor has a length of at least 1 meter. 12. The radio frequency component according to claim 1, wherein the first conductor and/or the second conductor has a length of at least 2 meters. 13. The radio frequency component according to claim 1, wherein the first conductor and/or the second conductor has a length of at least 3 meters. 14. The radio frequency component of claim 1, wherein the first conductor and/or the second conductor are configured with at least 5 turns. 15. The radio frequency component of claim 1, wherein the first conductor and/or the second conductor are configured with at least 10 turns. 16. The radio frequency component of claim 1, wherein the first conductor and/or the second conductor are configured with at least 15 turns. 17. The radio frequency component of claim 1, wherein the first conductor and/or the second conductor are configured with at least 20 turns. 18. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a target frequency, and wherein the number of turns of the first conductor and the second conductor is limited such that each of the self-resonance of the first conductor and the self-resonance of the second conductor is at a frequency at least twice the target frequency. 19. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a target frequency, and wherein the number of turns of the first conductor and the second conductor is limited such that each of the self-resonance of the first conductor and the self-resonance of the second conductor is at a frequency at least three times the target frequency. 20. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a target frequency, and wherein the number of turns of the first conductor and the second conductor is limited such that each of the self-resonance of the first conductor and the self-resonance of the second conductor is at a frequency at least five times the target frequency. 21. The radio frequency component of claim 1, wherein the first conductor is arranged according to a first coil configuration determined by optimization performed at least in part based on a model using the first coil and/or the second conductor is arranged according to a second coil configuration determined by optimization performed at least in part based on a model using the second coil. 22. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2T and greater than or equal to 0.1T. 23. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.1T and greater than or equal to 50mT. 24. The radio frequency component of claim 1, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 50 mT and greater than or equal to 20 mT. 25. The radio frequency component of claim 1, wherein, alternatively, the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 20 mT and greater than or equal to 10 mT. 26. The radio frequency component of claim 1, wherein, alternatively, the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 10 mT. 27. The radio frequency component of claim 1, wherein the first conductor includes a first continuous line arranged in a three-dimensional geometry around the region of interest as a plurality of turns, and the second conductor includes a second continuous line arranged in a three-dimensional geometry around the region of interest as a plurality of turns. 28. The radio frequency component of claim 27, wherein the first continuous line and the second continuous line are single-strand lines. 29. The radio frequency component of claim 27, wherein the first continuous line and the second continuous line are multi-strand lines. 30. A radio frequency component for magnetic resonance imaging, the radio frequency component comprising: A first coil, configured to generate a magnetic field component, includes a first conductor having a length of at least 1 meter and arranged in a plurality of turns, the first coil being arranged around a region of interest, wherein the spacing between the plurality of turns is non-uniform; A second coil, comprising a second conductor having multiple turns, is arranged around the region of interest and offset from the first coil relative to the region of interest, wherein the first and second coils are tuned to resonate at frequencies corresponding to a B0 field less than or equal to 0.2T and greater than or equal to 20mT; and A support structure for the first coil and the second coil, the support structure comprising: a first base layer to which the first coil is laid; and a second base layer to which the second coil is laid, wherein the support structure includes a helmet formed to accommodate a person's head. 31. The radio frequency component of claim 30, wherein the second substrate layer is located above the first substrate layer. 32. The radio frequency component of claim 30, wherein the first substrate layer includes at least one first slot arranged according to a first coil configuration to receive the first conductor, and the second substrate layer includes at least one second slot arranged according to a second coil configuration to receive the second conductor. 33. The radio frequency component of claim 32, wherein the first coil is configured with a first main axis, and the second coil is configured with a second main axis substantially orthogonal to the first main axis. 34. The radio frequency component of claim 32, wherein the first substrate layer and/or the second substrate layer are produced via three-dimensional printing. 35. The radio frequency component of claim 32, wherein the first substrate layer includes a convex side and a concave side, and wherein the second substrate layer includes a convex side and a concave side. 36. The radio frequency component of claim 35, wherein the at least one first groove is disposed on the convex side of the first substrate layer. 37. The radio frequency component of claim 36, wherein the at least one second groove is disposed on the convex side of the second substrate layer. 38. The radio frequency component of claim 36, wherein the at least one second groove is disposed on the concave side of the second substrate layer. 39. The radio frequency component of claim 35, wherein the at least one first groove is disposed on the concave side of the first substrate layer. 40. The radio frequency component of claim 39, wherein the at least one second groove is disposed on the convex side of the second substrate layer. 41. The radio frequency component of claim 39, wherein the at least one second groove is disposed on the concave side of the second substrate layer. 42. The radio frequency component of claim 30, wherein the second conductor has a length of at least 1 meter. 43. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor has a length of at least 2 meters. 44. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor has a length of at least 3 meters. 45. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor are configured with at least 5 turns. 46. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor are configured with at least 10 turns. 47. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor are configured with at least 15 turns. 48. The radio frequency component of claim 30, wherein the first conductor and/or the second conductor are configured with at least 20 turns. 49. The radio frequency component of claim 30, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2T and greater than or equal to 0.1T. 50. The radio frequency component of claim 30, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.1T and greater than or equal to 50mT. 51. The radio frequency component of claim 30, wherein the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 50 mT and greater than or equal to 20 mT. 52. The radio frequency component of claim 30, wherein, alternatively, the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 20 mT and greater than or equal to 10 mT. 53. The radio frequency component of claim 30, wherein, alternatively, the first coil and the second coil are tuned to resonate at a frequency corresponding to a B0 field less than or equal to 10 mT. 54. A radio frequency coil for magnetic resonance imaging, the radio frequency coil comprising: At least one conductor, having a length of at least 1 meter, is arranged in a three-dimensional geometry with multiple turns around the region of interest in a configuration optimized to increase sensitivity to magnetic resonance signals emitted within the region of interest, wherein the spacing between the multiple turns is non-uniform, and wherein the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2 T and greater than or equal to 20 m T; and A support structure for the radio frequency coil, the support structure including a first base layer to which the radio frequency coil is laid, wherein the support structure includes a helmet formed to accommodate a person's head. 55. The radio frequency coil of claim 54, wherein the configuration of the at least one conductor with a three-dimensional geometry is determined by performing at least one optimization based at least in part on a model of the radio frequency coil. 56. The radio frequency coil of claim 55, wherein the number of turns and/or the spacing between at least one pair of adjacent turns is determined by performing the at least one optimization based at least in part on a model using the radio frequency coil. 57. The radio frequency coil of claim 55, wherein the optimization determines the configuration of the model of the radio frequency coil that satisfies at least one constraint. 58. The radio frequency coil of claim 57, wherein the at least one constraint includes the resistance of the radio frequency coil. 59. The radio frequency coil of claim 57, wherein the at least one constraint includes the length of the at least one conductor. 60. The radio frequency coil of claim 57, wherein the at least one constraint includes the inductance of the radio frequency coil. 61. The radio frequency coil of claim 55, wherein the optimization determines the configuration of the model of the radio frequency coil that generates a magnetic field satisfying at least one standard. 62. The radio frequency coil of claim 61, wherein the at least one criterion includes magnetic field uniformity within the region of interest. 63. The radio frequency coil of claim 61, wherein the at least one criterion includes the magnetic field strength within the region of interest. 64. The radio frequency coil of claim 55, wherein the at least one conductor has a length of at least 2 meters. 65. The radio frequency coil of claim 55, wherein the at least one conductor has a length of at least 3 meters. 66. The radio frequency coil of claim 55, wherein the at least one conductor has a helical geometry having a plurality of turns. 67. The radio frequency coil of claim 55, wherein the plurality of turns comprises at least 5 turns. 68. The radio frequency coil of claim 55, wherein the plurality of turns comprises at least 10 turns. 69. The radio frequency coil of claim 55, wherein the plurality of turns comprises at least 15 turns. 70. The radio frequency coil of claim 55, wherein the plurality of turns comprises at least 20 turns. 71. The radio frequency coil of claim 55, combined with at least one auxiliary coil to facilitate noise suppression, the at least one auxiliary coil being located on or near the helmet. 72. The radio frequency coil of claim 71, wherein the radio frequency coil is configured to detect magnetic resonance signals emitted from a field of view located within the helmet, and wherein the at least one auxiliary coil is positioned to respond to ambient noise but not to the magnetic resonance signals emitted from the field of view. 73. The radio frequency coil of claim 72, wherein optimization is used to determine the configuration of the at least one auxiliary coil. 74. The radio frequency coil of claim 73, wherein the configuration of the at least one auxiliary coil is optimized to reduce or eliminate inductive coupling with the radio frequency coil. 75. The radio frequency coil of claim 54, combined with at least one auxiliary coil to facilitate noise suppression. 76. The radio frequency coil of claim 54, wherein the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2T and greater than or equal to 0.1T. 77. The radio frequency coil of claim 54, wherein the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.1T and greater than or equal to 50mT. 78. The radio frequency coil of claim 54, wherein the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 50 mT and greater than or equal to 20 mT. 79. The radio frequency coil of claim 54, wherein, alternatively, the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 20 mT and greater than or equal to 10 mT. 80. The radio frequency coil of claim 54, wherein, alternatively, the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 10 mT. 81. A radio frequency coil configured for a part of a patient's body, the radio frequency coil comprising: At least one conductor having a length of at least 1 meter and arranged in multiple turns around a region of interest, and said at least one conductor being oriented to generate a magnetic field component substantially orthogonal to the longitudinal axis of a portion of the patient's body, wherein the spacing between the multiple turns is non-uniform, and wherein the radio frequency coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2 T and greater than or equal to 20 m T; and A support structure for the radio frequency coil, the support structure including a first base layer to which the radio frequency coil is laid, wherein the support structure includes a helmet formed to accommodate a person's head. 82. The radio frequency coil of claim 81, wherein the radio frequency coil is configured to respond to a magnetic resonance signal emitted from a B0 field oriented substantially parallel to the longitudinal axis of the body. 83. The radio frequency coil of claim 82, wherein the radio frequency coil is configured to respond to a magnetic resonance signal emitted from a B0 field generated by a B0 magnet of biplane geometry. 84. The radio frequency coil of claim 82, wherein the radio frequency coil is configured to respond to a magnetic resonance signal emitted from a B0 field oriented substantially parallel to the longitudinal axis of the body. 85. The radio frequency coil of claim 84, wherein the radio frequency coil is configured to emit a magnetic resonance signal in response to a B0 field generated from a B0 magnet of solenoid geometry. 86. A low-field magnetic resonance system, comprising: B0 magnet, the B0 magnet being configured to generate a B0 magnetic field of 0.2 Tesla (T) or less in the field of view; A first coil, comprising a first conductor having a length of at least 1 meter and arranged in a plurality of turns, the first coil being configured to generate a magnetic field component within the field of view, wherein the spacing between the plurality of turns is non-uniform; A second coil, configured to respond to a magnetic resonance signal component emitted from the field of view; and A support structure for the first coil and the second coil, the support structure comprising: a first base layer to which the first coil is laid; and a second base layer to which the second coil is laid, wherein the support structure includes a helmet formed to accommodate a person's head. 87. The low-field magnetic resonance system of claim 86, wherein the second coil is offset from the first coil relative to the field of view when positioned within the B0 magnetic field. 88. The low-field magnetic resonance system of claim 86, wherein the plurality of turns of the first coil are arranged in a three-dimensional geometry around the field of view when located within the B0 magnetic field. 89. The low-field magnetic resonance system of claim 88, wherein the second coil includes a second conductor having a plurality of turns, the second conductor being arranged in a three-dimensional geometry around the field of view when located within the B0 magnetic field. 90. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured to generate a B0 magnetic field less than or equal to 0.1 T and greater than or equal to 50 m T. 91. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured to generate a B0 magnetic field of less than or equal to 50 mT and greater than or equal to 20 mT. 92. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured to generate a B0 magnetic field of less than or equal to 20 mT and greater than or equal to 10 mT. 93. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured to generate a B0 magnetic field of less than or equal to 10 mT. 94. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured with a biplane geometry. 95. The low-field magnetic resonance system of claim 86, wherein the B0 magnet is configured in a solenoid geometry. 96. An apparatus for a magnetic resonance imaging system, the apparatus comprising: A first coil, comprising a first conductor having a length of at least 1 meter and arranged in a plurality of turns, the first coil being oriented in response to a first magnetic resonance signal component, wherein the spacing between the plurality of turns is non-uniform, and wherein the first coil is tuned to resonate at a frequency corresponding to a B0 field less than or equal to 0.2 T and greater than or equal to 20 m T. A support structure for the first coil, the support structure including a first base layer to which the first coil is laid, wherein the support structure includes a helmet formed to accommodate a person's head; and At least one controller is configured to operate the first coil to generate a radio frequency magnetic field and a gradient field. 97. The apparatus of claim 96, wherein the at least one controller is configured to operate the first coil according to a magnetic resonance pulse sequence. 98. The apparatus of claim 96, wherein the at least one controller is configured to use the first coil to detect a magnetic resonance signal. 99. The apparatus of claim 96, wherein the first coil is configured to operate in a low-field magnetic resonance imaging system. 100. The apparatus of claim 96, further comprising a second coil, wherein the at least one controller is configured to operate the second coil to generate a radio frequency magnetic field and a gradient field. 101. The apparatus of claim 100, wherein the first coil and the second coil generate at least a portion of a gradient magnetic field in a direction substantially orthogonal to the BO magnetic field of the system. 102. The apparatus of claim 100, wherein the first coil generates at least a portion of a gradient magnetic field in a first direction substantially orthogonal to the BO magnetic field of the system, and wherein the second coil generates at least a portion of a gradient magnetic field in a second direction substantially orthogonal to the BO magnetic field. 103. The apparatus of claim 102, wherein the at least one controller is configured to use the first coil and the second coil to detect magnetic resonance signals. 104. The apparatus of claim 103, wherein the magnetic resonance signals detected by the first coil and the second coil are combined to increase the signal-to-noise ratio. 105. The apparatus of claim 103, wherein the first coil and the second coil are operated to reduce the acquisition time. 106. The apparatus of claim 100, wherein the at least one controller comprises: A first amplifier is used to provide power to operate the first coil to generate the radio frequency magnetic field; A high-pass filter, the high-pass filter being connected between the first amplifier and the first coil; A second amplifier is used to provide power to operate the first coil to generate the gradient magnetic field; and A low-pass filter is connected between the second amplifier and the first coil. 107. The apparatus of claim 100, wherein optimization is used to determine the configuration of the first coil. 108. The apparatus of claim 107, wherein the optimization determines the configuration of at least one conductor of the first coil that satisfies at least one constraint and generates a magnetic field satisfying a predetermined criterion when simulating the operation of a model of the first coil. 109. The apparatus of claim 100 further includes at least one auxiliary coil for noise suppression, the at least one auxiliary coil being located on or near the helmet. 110. The apparatus of claim 109, wherein the first coil is configured to detect magnetic resonance signals emitted from a field of view located within the helmet, and wherein the at least one auxiliary coil is positioned to respond to ambient noise but not to magnetic resonance signals emitted from the field of view. 111. The apparatus of claim 109, wherein optimization is used to determine the configuration of the at least one auxiliary coil. 112. The apparatus of claim 109, wherein the configuration of the at least one auxiliary coil is optimized to reduce or eliminate inductive coupling with the first coil.