Magnetic resonance apparatus and method

Abstract

A nuclear magnetic resonance coil and a method of using the magnetic resonance coil, the coil being configured in a first mode to receive and conduct current at a drive port to generate a static magnetic field in a space adjacent to the coil, and in a second mode to receive and output nuclear magnetic resonance signals generated in the space at a receive port.

Description

[Technical field]

[0001] FIELD OF THE DISCLOSURE The embodiments described herein relate generally to magnetic resonance devices and, more particularly, to magnetic resonance devices and methods in which the coils used to receive magnetic resonance signals also serve to provide a polarizing quasi-static magnetic field. [Background technology]

[0002] Known magnetic resonance devices use either coils or permanent magnets to generate a static magnetic field, which is intended to be uniform over a volume of interest, and a single, electrically independent coil is used to detect the magnetic resonance signals generated in the volume of interest. [Brief description of the drawings]

[0003] [Figure 1] FIG. 1 illustrates an NMR system, according to one embodiment. [Diagram 2] FIG. 2 illustrates a coil activation sequence used in one embodiment. [Figure 3A] FIG. 1 illustrates spin polarization under the influence of a field B0prepolarise. [Figure 3B] FIG. 1 illustrates spin polarization under the influence of a field B0 measurement. [Figure 3C] FIG. 13 shows the spin polarization behavior after application of field B1. [Figure 4] FIG. 13 shows a coupling circuit for a dual-purpose coil. [Diagram 5] FIG. 13 shows another coupling circuit for a dual-purpose coil. [Figure 6] FIG. 13 is a cross-sectional view of an axisymmetric simulation of a dual-purpose coil according to an embodiment. [Figure 7] FIG. 2 is a diagram showing the properties of materials used in the magnetic core of one embodiment. [Figure 8A] 1 is a schematic diagram showing a cross-sectional view of the right half of an axisymmetric NMR system, according to one embodiment. [Figure 8B]FIG. 1 is a three-dimensional isometric view of an NMR system, according to one embodiment. [Figure 8C] FIG. 8B is a detailed view of a portion of the NMR system shown in FIG. 8A. [Figure 9A] FIG. 1 illustrates a dual-purpose coil divided into multiple sections, according to one embodiment. [Figure 9B] FIG. 2 illustrates a circuit for connecting coils L1-L4 of a dual-purpose coil according to one embodiment. [Figure 10A] FIG. 13 is a cross-sectional view of a two-dimensional axisymmetric simulation of the unsafe zone for an embodiment of an NMR system without a countercoil and a magnetic structure. [Figure 10B] 1 is a cross-sectional view of a two-dimensional axisymmetric simulation of the unsafe zone for an embodiment of an NMR system having a countercoil and a magnetic structure. [Figure 11] 1 is a schematic diagram of a combination of a passive cooling element 1010 and a magnetic structure 1020 of an NMR system, according to one embodiment. [Figure 12] 11 is a schematic diagram showing a cross section of a dual purpose coil 1110, according to one embodiment. [Figure 13] FIG. 11 illustrates simulated temperature profiles of the bed, magnetic structure, and combined coil 1110 over seven measurement cycles, according to one embodiment. [Figure 14] FIG. 14 shows the temperature distribution of the simulated NMR system shown in FIG. 13 at time=3.725 hr. [Figure 15] FIG. 13 illustrates simulation results of coil sensitivity spatial profiles, according to one embodiment. [Figure 16] 16A and 16B show the results of a simulation of the loss of SNR of an NMR signal using active noise cancellation without a gap and with a gap. [Figure 17] FIG. 2 illustrates a magnetic structure according to one embodiment. [Figure 18] Figure 18A illustrates the eddy current backfield resulting from the casing top of a metal enclosure 1702 without a passive coil 1802, according to one embodiment, and Figure 18B illustrates the eddy current backfield resulting from the casing top of a metal enclosure 1702 with a passive coil 1802, according to one embodiment. [Figure 19] FIG. 13 illustrates the eddy current backfield resulting from eddy currents propagating to the top of the casing of a metal enclosure below a dual-purpose coil when a magnetic structure is present but no passive coil is present, according to one embodiment. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0004] According to one embodiment, a nuclear magnetic resonance coil is provided, configured in a first mode to receive and conduct at a drive port an electric current for generating a static magnetic field in a volume adjacent to the coil, and in a second mode to receive and output at a receive port a nuclear magnetic resonance signal generated in the volume, the first and second modes being consecutive to each other.

[0005] In one embodiment, the coil comprises a plurality of inductors, all of which are used in generating the static magnetic field, but only a subset of the inductors, or only one of the inductors, is used for sensing the NMR signal. The plurality of inductors may be provided discretely or may share the same winding core.

[0006] In one embodiment, only a subset of the inductors of the multiple inductors, or only one inductor of the multiple inductors, that are closest to the patient-contacting surface of the coil and / or closest to the centerline of the coil are used to sense the NMR signal.

[0007] In one embodiment, the coil comprises a plurality of inductors that are electrically connected in series for DC current and are electrically connected such that during signal reception, the signal is not amplified by a plurality of inductors that do not form part of a subset or only one of the plurality of inductors.

[0008] In one embodiment, the coil includes an electrical circuit that electrically isolates the drive port and the receive port from each other.

[0009] In one embodiment, the coil is dimensioned to generate a static magnetic field in a volume of interest that enables acquisition of nuclear magnetic resonance (NMR) signals throughout the depth of the torso or other part of the body of an adult subject positioned prone or supine on the surface of the coil.

[0010] In one embodiment, the NMR signal is a magnetic resonance imaging (MRI) signal.

[0011] In one embodiment, the electrical circuit is a passive circuit.

[0012] According to another embodiment, a nuclear magnetic resonance coil is provided, the coil comprising a ferromagnetic core surrounded by a coil winding. In one embodiment, the nuclear magnetic resonance coil is a nuclear magnetic resonance coil as previously described, i.e., a nuclear magnetic resonance coil configured to operate in the aforementioned first mode and the aforementioned second mode.

[0013] In one embodiment, the NMR signal voltage received by the coil is amplified by a factor of 20 or less, preferably a factor of 5 or less. In other words, in this embodiment, low amplification is applied to the received NMR signal. This is advantageous in situations where the entire combined coil described herein is used for signal reception.

[0014] In another embodiment, where a small portion of the dual-purpose coil is used for signal reception, the NMR signal voltage received by the coil is amplified by a factor of 1000 or more.

[0015] In another embodiment, the coil is non-resonant.

[0016] In another embodiment, a standard amplifier coil is used in MRI.

[0017] In another embodiment, the self-resonant frequency of the coil is selected such that the highest Larmor frequency observed approaches the self-resonant frequency of the coil while maintaining a high input signal sensitivity, hi one embodiment, the self-resonant frequency of the coil is selected such that the highest Larmor frequency observed using the coil is 0.9 times or less, preferably 0.8 times or less, of the self-resonant frequency of the coil.

[0018] According to another embodiment, there is provided a nuclear magnetic resonance coil comprising a patient side adjacent to which a patient is to be positioned during a magnetic resonance examination and a soft ferromagnetic shielding on at least one side of the coil other than the patient side. In one embodiment, the nuclear magnetic resonance coil is as hereinbefore described.

[0019] According to another embodiment there is provided a nuclear magnetic resonance apparatus comprising a static magnetic field driver, a receiving chain and a nuclear magnetic resonance coil as claimed in any of the preceding claims.

[0020] In one embodiment, the magnetic resonance device further comprises a static magnetic field coil configured to generate a static magnetic field substantially perpendicular to the static magnetic field generated by the magnetic resonance coil in the region of interest of the device and driver for driving the static magnetic field coil when energized. In this embodiment, the magnetic resonance device is configured to adiabatically switch between the static magnetic field generated by the magnetic resonance coil and the static magnetic field generated by the static magnetic field coil.

[0021] In this embodiment, the region of interest is located on the patient side of the nuclear magnetic resonance coil and at a distance of 20 cm from the front surface of the nuclear magnetic resonance coil facing the patient.

[0022] In an alternative embodiment, the static magnetic field substantially perpendicular to the static magnetic field generated by the nuclear magnetic resonance coil is not generated by a coil, but instead by a permanent magnet that does not need to be selectively energized. In this embodiment, it will be understood that the prepolarization field is the sum of the static magnetic field generated by the nuclear magnetic resonance coil and the static magnetic field generated by the permanent magnet. In one embodiment, this sum can be dominated by the static magnetic field generated by the nuclear magnetic resonance coil to the extent that the summed field is still substantially perpendicular to the static magnetic field generated by the permanent magnet. For example, the static magnetic field generated by the nuclear magnetic resonance coil can be 200 mT at a given location in the coil's volume of interest, while the static magnetic field generated by the permanent magnet can have a strength of 1 mT. In another embodiment, the static magnetic field generated by the nuclear magnetic resonance coil and the static magnetic field generated by the permanent magnet have a relative strength such that the sum of both fields is no longer substantially perpendicular to the static magnetic field generated by the permanent magnet. For example, the relative strength of the fields can be such that the sum of the fields is inclined 45 to 135 degrees to the direction of the static magnetic field generated by the permanent magnet. In this example, the sum of the fields is much larger than the field generated by the nuclear magnetic resonance coil alone, so that a high degree of prepolarization is achieved. It will be understood that even in this example, the magnetization is preceded around the direction of the static magnetic field generated by the permanent magnet during the acquisition of the nuclear magnetic resonance signal. The shape of the nuclear magnetic resonance coil is not affected by the choice of the permanent magnet for generating the static magnetic field for measurement, but the nuclear magnetic resonance coil is, on the other hand, highly sensitive to this signal.

[0023] According to another embodiment, a nuclear magnetic resonance coil is provided that includes a ferromagnetic core and a coil winding wound around the ferromagnetic core.

[0024] In one embodiment, the ferromagnetic core comprises a plurality of contacting ferromagnetic components that are electrically insulated from one another.

[0025] In one embodiment, the components are electrically insulated from one another by insulation extending in a radial plane that contains the longitudinal axis of the coil.

[0026] In one embodiment, the coil system comprises a first coil as described above having a first diameter and a first longitudinal axis, the system further comprising a second coil having a second diameter and a second longitudinal axis, the first and second longitudinal axes being substantially coincident, the second diameter being greater than the first diameter, and the first and second coils being arranged such that at a distance from the longitudinal axis greater than the second diameter, the second coil generates a field that opposes the field generated by the first coil.

[0027] In one embodiment, the coil system further comprises a flux guide component positioned to guide magnetic flux generated by the second coil away from the longitudinal axis.

[0028] In one embodiment, the flux guide component comprises one or more, or all of the following: a flux guide plate disposed below the second coil and having a footprint adjacent to the second coil or projecting radially inwardly and / or radially outwardly beyond the second coil; a flux guide plate disposed between the first coil and the second coil, preferably adjacent to and radially inward of the second coil; A flux guide plate configured to extend in a radially outward direction and from the first coil toward the region of interest from a point at or beyond a maximum diameter of the second coil in the deployed position.

[0029] In one embodiment, the flux guide plate configured to extend in a radially outward direction and in a direction from the first coil towards the region of interest is further configured to be moved to a stowed away position in which the flux guide plate does not extend any further in the radially outward direction from a point at or beyond the maximum diameter of the second coil.

[0030] In one embodiment, the flux guide plate, configured to extend in a radially outward direction and in a direction from the first coil towards the region of interest, is planar and hingedly connected to a point at or beyond the maximum diameter of the second coil.

[0031] In one embodiment, the coil system is further configured to simultaneously activate and / or deactivate the first and second coils.

[0032] In one embodiment, the second coil is more sensitive to noise sources in the far field than the first coil.

[0033] In one embodiment, the ferromagnetic material is ferrite.

[0034] In one embodiment, suitable ferromagnetic materials include 10 -1 Ferromagnetic materials having any possible combination of these properties are suitable for use in the embodiments.

[0035] According to another embodiment, a magnetic resonance method is provided that includes generating a first static magnetic field in a region of interest using a coil by applying a current to the coil, ceasing application of the current to the coil, generating a second static magnetic field in the region of interest, applying a radio frequency magnetic field to the region of interest at a frequency based on the strength of the second static magnetic field in the region of interest, and receiving any nuclear magnetic resonance signals generated in the region of interest.

[0036] In one embodiment, ceasing the application of the current to the coil and generating the second static magnetic field are performed such that an adiabatic switching is performed between the static magnetic field generated by the current through the coil and the second static magnetic field in a region of interest of the magnetic resonance device in which the method is performed. It will be appreciated that in one embodiment, this switching is achieved by at least a partial overlap between the ramping down of the current through the coil and the ramping up of the current in the coil generating the second static magnetic field.

[0037] In one embodiment, nuclear magnetic resonance signals are acquired simultaneously with the application of the radio frequency magnetic field.

[0038] According to one embodiment, there is provided a nuclear magnetic resonance coil system comprising a coil on a ferromagnetic core, the ferromagnetic core having a plurality of spokes, each of which extends beneath the coil from a center beneath the coil to a radial end at a radial distance greater than a diameter of the coil, some or each of the plurality of spokes further extending upward from the radial end to form an upwardly extending portion outside a maximum diameter of the coil, and gaps between the spokes and / or the upwardly extending portions include a material having a thermal conductivity greater than that of the ferromagnetic core.

[0039] In one embodiment, a material having a thermal conductivity greater than that of the ferromagnetic core is further provided over the coil and / or in the center of the coil.

[0040] In one embodiment, the coil comprises a longitudinal axis and a plurality of layers and / or windings stacked on the longitudinal axis and radially adjacent to one another, the coil further comprising a material having a thermal conductivity that exceeds the thermal conductivity of the windings of the coil and / or the thermal conductivity of the ferromagnetic core.

[0041] According to one embodiment, a magnetic resonance system is provided comprising a first coil having a longitudinal axis and sensitive to radio frequency signals emitted from a region of interest and electromagnetic noise emitted outside the region of interest, the system further comprising a noise cancellation coil having a longitudinal axis substantially coincident with the longitudinal axis of the coil, the noise cancellation coil being sensitive to electromagnetic noise emitted outside the region of interest, the system being configured to sense the noise emitted outside the region of interest and to subtract the noise from the signal received by the first coil using a predetermined scaling factor.

[0042] In one embodiment, the ferromagnetic shield is positioned to reduce the sensitivity of the noise cancellation coil to signals emanating from the region of interest.

[0043] In one embodiment, it is desirable for the ratio of the sensitivity to background noise of the noise cancellation coil to the sensitivity to background noise of a coil used to sense NMR signals, such as a dual-purpose coil, to be greater than one.

[0044] In one embodiment, the ratio of the sensitivity to NMR signals of the noise cancellation coil to the sensitivity to NMR signals of a coil used to sense the NMR signals, such as a dual-purpose coil, is preferably less than 1, and more preferably much less than 1.

[0045] In one embodiment, the sensitivity of the noise cancellation coil to noise sources in the far field is greater than the sensitivity of the first coil.

[0046] In one embodiment, the system includes a second noise cancellation coil having a sensitivity profile that enables the second noise cancellation coil to sense noise in spatial regions where the noise cancellation coil is insufficiently sensitive to enable removal of the noise detected by the first coil, and the noise sensed by the second noise cancellation coil is subtracted from the signal received by the first coil using a second predetermined scaling factor.

[0047] In one embodiment, the predetermined scale factor and / or the second predetermined scale factor are scale factors determined using an empirical determination of the relative sensitivity of the first coil and the noise cancellation coil, hi another embodiment, the predetermined scale factor and / or the second predetermined scale factor are scale factors determined using a simulation to determine the relative sensitivity of the first coil and the noise cancellation coil.

[0048] According to one embodiment, an NMR system is provided comprising a coil configured to generate a time-varying magnetic field, a conductive structure and a passive coil positioned between the coil and the conductive structure, the conductive structure being located in the time-varying magnetic field generated by the coil in the absence of the passive coil, the passive coil being positioned in the time-varying magnetic field such that the time-varying magnetic field induces a current in the passive coil, the passive coil comprising a variable resistance.

[0049] In one embodiment, the variable resistor is configured to present a resistance that allows eddy currents to form in the coil, and then present a higher resistance that causes the eddy currents to dissipate.

[0050] In one embodiment, the variable resistor is configured to exhibit an initial resistance that allows eddy currents to form in the coil, and then exhibit a higher resistance that dissipates the eddy currents, hi one embodiment, the variable resistor gradually increases from its initial resistance to the higher resistance.

[0051] According to one embodiment, there is provided a method of operating an NMR system including generating a time-varying magnetic field using a coil, a passive coil positioned between the coil and a conductive structure, the conductive structure positioned in the time-varying magnetic field that would be generated by the coil in the absence of the passive coil, the passive coil positioned in the time-varying magnetic field such that the time-varying magnetic field induces a current in the passive coil, and the method includes changing a resistance of the passive coil to suppress or stop eddy currents flowing in the coil after the eddy currents are generated.

[0052] In one embodiment, the variable resistor is configured to present an initial resistance that allows eddy currents to form in the coil and then present an open circuit that prevents the eddy currents from flowing in the passive coil. Preferably, a voltage limiting circuit is also provided that allows dissipation of energy stored in the passive coil.

[0053] In one embodiment, a method of operating any of the coils or systems described above is provided.

[0054] According to another embodiment, a nuclear magnetic resonance coil or a method of operating a nuclear magnetic resonance coil is provided, the nuclear magnetic resonance coil being configured to alternately generate a static magnetic field and receive nuclear magnetic resonance signals, the coil comprising a conductor and a cooling arrangement configured to flow a cooling fluid through the conductor.

[0055] In one embodiment, the conductor is disposed within the fluid conduit and the coil is positioned to allow liquid to flow within the fluid conduit.

[0056] In one embodiment, the coil includes a pump that pumps liquid into the tube.

[0057] In one embodiment, the conductor is a tube and the liquid flows within the lumen of the tube.

[0058] In one embodiment, at least a portion of the conductor is disposed within a fluid-tight container containing a liquid.

[0059] In one embodiment, the conductors form windings, the windings being spaced apart from one another to allow liquid to circulate between adjacent windings.

[0060] It has been found that in known magnetic resonance devices, multiple coils compete to maximize the field strength generated in a volume of interest or to maximize their sensitivity to magnetic resonance signals generated in the volume of interest, respectively. In the following, a magnetic resonance device is disclosed that mitigates this competition for magnetic or electromagnetic access to a volume of interest by using a single coil for generating a prepolarizing static magnetic field in the volume of interest and for sensing the magnetic resonance signals generated in the volume of interest.

[0061] FIG. 1 shows an NMR system 100 according to one embodiment. The NMR system 100 is configured with a static magnetic field B 0prepolarise 1 as extending vertically, this is not required. As will be explained further below, the dual coil 110 can be energized and de-energized to generate a field B. 0prepolarise can be activated and deactivated accordingly.

[0062] The system 100 further comprises two coils 120 and 130. The coil 120 also generates a static magnetic field B 0measurement As can be seen in Figure 1, this field is the field B 0prepolarise As in the case of the dual-purpose coil 110, the coil 120 can be energized and de-energized to provide a field B 0measurement can be activated and deactivated accordingly.

[0063] Coil 130 is in field B. 0measurement creates a B1 radio frequency (RF) magnetic field at the precession frequency generated by the B 0prepolarise and place B 0measurement As is the case with known NMR RF coils, the B1 field can be activated and deactivated.

[0064] The dual-purpose coil 110 and the two coils 120 and 130 generate a magnetic field B 0prepolarise、 B 0measurement , and B1 are configured to be generated in a space 140 occupied by a subject, such as a patient, that is the subject of the NMR measurements.

[0065] In one embodiment, B 0prepolaris may vary from about 50 mT to about 300 mT within the space 140. In one embodiment, B 0measurement may be about 1 mT in the space 140. In one embodiment, B 0measuremen may be about 2 mT within the space 140.

[0066] Although a particular configuration of the system 100 is shown in FIG. 1, the spatial arrangement of the magnets and coils shown in FIG. 1 is not required. In an exemplary embodiment, the generated magnetic fields are substantially mutually perpendicular to one another. In another embodiment, the directions of the two fields are parallel. Although the dual coil 110 and the coils 120 and 130 are shown in FIG. 1 as being spaced apart from one another by gaps, it will be understood that these gaps are shown for illustrative purposes only, and that some or all of the gaps shown may be omitted in the physical implementation of the illustrated system 100, and that the dual coil 110 and some or all of the coils 120 and 130 may instead be provided in a single unit. In one example, the coils 120 and 130 may form part of a single printed circuit board (PCB).

[0067] FIG. 2 shows an activation sequence 200 for the dual-purpose coil 110 and the coils 120 and 130. FIGS. 3A-3C show the change in state of the polarization vector resulting from various conditions of the applied field. In a first step 210, B 0prepolaris The dual-purpose coil 110 is energized so that a field is generated as shown in FIG. 0prepolariseThe duration of application of field B may be selected to provide the desired imaging parameters. For example, if the generation of T1 contrast is to be avoided, then field B 0prepolarise may be applied for a period exceeding the longest T1 expected in the sample being investigated. If T1 contrast between the various spin species is to be generated, B 0prepolarise is applied for a period of time shorter than the T1 relaxation time of one spin species but longer than the T1 relaxation time of the other spin species. As is well known, a static magnetic field applied to a spin species generates a magnetization as indicated by the arrows shown in FIG. 3A. In step 220, dual-purpose coil 110 is deactivated and coil 120 is activated. As described above, the respective static magnetic fields B produced by dual-purpose coil 110 and coil 120 0prepolarise and B. 0measurement extend substantially perpendicular to each other. The deactivation of the dual-purpose coil 110 and the activation of the coil 120 occur in a short time frame, specifically, time t ≪ T1, where T1 is the time 0prepolarise Or B 0measurement (since T1 is a function of the field strength), and the shortest longitudinal relaxation time associated with either 0prepolarise The polarization generated by B 0measurement The plane is now aligned with the horizontal plane. 0prepolarise From the consistency with field B 0measurement The switching of the magnetization to a matching with is shown in FIG. 3B and occurs without precession, and the components producing the net magnetization remain in phase with each other. This is known as adiabatic pulsing / transfer / switching.

[0068] Because the dual-purpose coil 110 is de-energized at the end of step 220, it may be switched to a receive mode in step 230. In one embodiment, once the dual-purpose coil 110 is in the receive mode, an RFB1 field may be applied in step 240. This creates a field B, as shown in FIG. 0measurement Due to the magnetization precession around the direction of the field B 0measuremenThe magnetization vectors are tilted in a plane substantially perpendicular to the direction of the magnetization vectors in ... 0prepolarise By using electromagnets to generate the field, the strength, direction, and duration of the prepolarization field can be varied by changing the current applied to the dual coil. This allows the measurement of various T1-weighted MR signals in combination with gradients to provide images with intensity changes corresponding to the longitudinal relaxation time T1 of the tissue that gives rise to the magnetic resonance signal. In one embodiment, the prepolarization field is applied for a time that exceeds the expected T1 in the tissue, maximizing the acquired signal. Conversely, in another embodiment, the prepolarization field may be applied for a shorter period of time. This reduces the time required for longitudinal relaxation, at the expense of signal strength, and subsequent prepolarization and associated imaging steps can be performed sooner than if the prepolarization field was activated for longer than T1.

[0069] In another embodiment, images are acquired for different current amplitudes used in generating the prepolarization field, and the resulting variation in prepolarization field strength from image to image causes image contrast to vary from image to image as a function of T1.

[0070] n for a nucleus of spin 1 / 2 at a given field strength B0 and a given temperature T - and n + The difference in the occupation of each of the spin states can be expressed as:

number

[0071] The coil of the dual-purpose coil 110 can carry a much higher current than the coil 120. As a result, the prepolarization B 0prepolarise The field is measurement field B. 0measurement As a result, the spin is in the field B 0prepolariseThe difference in the state of the spin occupancy while the stationary state at 0measurement is larger than the difference in the spin occupancy between the stationary state at B 0prepolarise The sample according to B 0measurement The net magnetization is larger than that of the sample that follows

[0072] B, as described above in connection with step 220 0prepolarise From B 0measurement The change in the difference in the spin population states, which inevitably occurs when switching the static magnetic field to B, is not instantaneous, but is instead characterized by the longitudinal relaxation time T1. As a result, 0prepolarise The advantage realized in the net magnetization in field strength is that, as described above in connection with step 220, the static magnetic field B 0measurement It will be understood that after being switched to, the current is maintained for a while. Due to this advantage, in the steady state, 0measurement The spins are then exposed to a B 0measurement As a result, NMR measurements are controlled by the T1 relaxation time of the spins, 0prepolarise From B 0measurement For a period of time following the switch to B 0measurement This can be done using a B1 field with a resonant frequency determined by 0measurement Due to the small norm of B 0measurement The magnetic field can have low absolute inhomogeneity and large relative inhomogeneity. As a result, losses through dephasing of signals in the projected field configurations disclosed herein can be avoided. This allows the B 0measurement For a frequency of precession / signal readout chosen to be proportional to , the norm of the longitudinal magnetization can be arbitrarily small, B, as long as adiabatic switching is achieved. 0prepolarise And B 0measurement It is desirable to complete the adiabatic switching as quickly as possible, but without violating the regulation of stimulation to the peripheral nerve. In one embodiment, B0prepolarise From B 0measurement The adiabatic switching to B is completed within a time window less than the shortest longitudinal relaxation time T1 of all spin species for which NMR signals are acquired. 0measurement Case B with a gradual increase in 0prepolarise In one embodiment, the NMR signal may be accompanied by a gradual reduction in B 0measurement is completely ramped down, and B 0measurement is fully ramped up, for a period of time shorter than the shortest longitudinal relaxation time T1 of all the spin species for which NMR signals are acquired. In another embodiment, NMR signals are additionally acquired after the shortest longitudinal relaxation time T1 of all the spin species for which NMR signals are acquired, until the end of the longer or longest longitudinal relaxation time T1 of another of the spin species for which NMR signals are acquired. Then, when the field B 0prepolarise A further measurement cycle can be started by reactivating , again prepolarizing the spins to be examined.

[0073] In an alternative embodiment, the switching of the fields is not adiabatic. Instead, field B 0prepolarise is B 0measurement The field B is reduced to a non-zero value in a time frame that does not allow for redistribution of the spin distribution that was generated at the full strength of 0prepolarise When is small enough, say 10% of its full strength, B 0measurement is rapidly activated, while B 0prepolarise is also rapidly deactivated. 0prepolarise Not only is the magnitude of the magnetization created by B maintained, but the magnetization is also increased by 100% without the need to apply a B excitation pulse. 0measurement The precession about also begins.

[0074] FIG. 4 illustrates an embodiment of a dual-purpose coil 110 in a prepolarization field B 0prepolarise 5 shows an example network 500 that may be used to generate a prepolarization driver 570 for alternately switching the dual-purpose coil 110 to a receive mode. As can be seen, the network 500 does not include any active components, but instead is a passive network.

[0075] The connection 510 to the prepolarization driver comprises two sets of cross-coupled diodes 520 connected between each terminal of the dual-purpose coil 110 and a respective port to the prepolarization driver. A capacitor 540 is also provided across the terminals leading to the ports for the prepolarization driver. The capacitor 540 forms a low-pass filter with a cutoff frequency below the frequency of the magnetic resonance signals that the system 100 is designed to generate or receive, preventing higher frequency signals generated by the driver 570 from passing to the coil 110. In one embodiment, this low-pass may be omitted if high frequencies are not expected to come from the driver 570 and the input impedance of the port is high enough to avoid changing the resonance behavior of the coil 110. In this way, a direct current may be provided to the dual-purpose coil 110 from the port connectable to the prepolarization driver via the diode 520, while the magnetic resonance signals are also prevented from leaking into the prepolarization driver. The cross-coupled diodes 520 also block lower amplitude signals, such as received magnetic resonance signals, and create a very high impedance path / filter when the coil is in receive mode, eliminating or at least reducing noise generated by the driver 570 while allowing signals with amplitudes above the threshold voltage of the diodes (i.e., B 0prepolarise (the signal that produces the signal) to pass.

[0076] On the receive side 550, two capacitors 560 prevent direct current and large DC voltages applied to the dual-purpose coil via connection / network 510 from being applied to the receive chain 590. In further embodiments, additional cross-coupled diodes may be provided to ground each of the terminals of the two capacitors 560 connected to the receive ports.

[0077] FIG. 5 illustrates an embodiment of a dual-purpose coil 110 in a prepolarization field B 0prepolarise5 shows another example network 700 that may be used to generate a prepolarization driver 710 for alternately switching the prepolarization driver 710 to a receive mode and allowing the prepolarization driver 110 to transmit the received NMR signal to a low noise amplifier 720. The circuit shown in FIG. 5 also includes only passive components. As shown in FIG. 5, the circuit includes diodes 730. Although a single diode is shown connected to the terminals of the prepolarization driver 710, in an alternative embodiment, a pair of cross-coupled diodes per terminal may instead be provided in the manner shown in FIG. 4. Two further diodes 740 each include a parasitic capacitance. The total capacitance presented to the prepolarization driver 110, together with the inductance of the coil 110, determines its resonant frequency. In this embodiment, it is desirable for the resonant frequency of the coil 110 to occur above the observed NMR frequency, but not at or near the observed NMR frequency. Capacitor C2 is in series with the parasitic capacitance of the diode 740, thereby lowering the overall capacitance presented to the prepolarization driver 710. This increases the resonant frequency of the coil 110.

[0078] Diodes 730 and 740 present a high impedance to the NMR signal received by the dual-purpose coil 110 and represent signal leakage to the prepolarization coil driver 710. In an alternative embodiment, diode 730 may be replaced by an inductor depending on the desired cutoff frequency of the network connecting the prepolarization driver 710 with the dual-purpose coil 110. In an alternative embodiment, the diode is replaced by a parallel LC tank circuit that is tuned to the operating frequency.

[0079] As can be seen from the above, passive isolation of the prepolarized driver from the receive chain and vice versa is made possible by the operating frequency difference of the two branches and the various signal amplitudes used (essentially switching diodes on and off).

[0080] Use of the coil 110 under receive mode can be matched with any conventional receive electronics desired, depending on the frequency, past the passive switches (i.e., DC blocking capacitors and diodes that protect the receive stage), since in one embodiment where the coil nodes are in contact only with cross-coupled diodes and DC blocking capacitors, the coil functions essentially as any other receive coil in NMR / MRI, and can therefore be tuned, matched and resonated at one or more frequencies, or simply connected directly to an amplifier, for example.

[0081] In another embodiment, diodes 730 and 740 may be replaced by simple active switches that can interrupt the connection from coil 110 to driver 710 .

[0082] [Dual-purpose coil magnetic core] FIG. 6 shows an axisymmetric simulation cross-section of the dual-purpose coil 110 of one embodiment. The dual-purpose coil 110 of this embodiment is rotationally symmetric about an axis coinciding with the ordinate of FIG. 6. As a result, it will be appreciated that only the right-most half of the cross-section of the dual-purpose coil 110 is shown in FIG. 6. The dual-purpose coil 110 comprises windings 610 forming a solenoid coil. Furthermore, a magnetic core 620 is provided at the center of the solenoid. In this embodiment, the magnetic core is cylindrical. The static magnetic field B generated by the dual-purpose coil 110 is 0prepolarise The results of this simulation are also shown in FIG. 6. In this simulation, the magnetic core 620 is made of a ferrite compound. The high B value (about 0.5T) indicates the rough contours of the ferrite material. The current density in the coil windings is shown in the inset image, highlighting the cross-sections of the individual coil windings. In one embodiment, the individual copper windings are themselves made of Litz cross-section / Litz wire.

[0083] The magnetic core 620 acts as a flux concentrator to concentrate the magnetic flux generated by the solenoid 610 in an area 630 occupied by a patient during use, and in particular in a field of interest 640 up to about 20 cm above the top surface of the dual-purpose coil 110. The magnetic core 620 is a quasi-static magnetic field B 0prepolarise and the high frequency magnetic field generated in the region of interest 630 during use of the dual-purpose coil 110. The frequency of the magnetic resonance signal generated in the region of interest 630 is 0prepolarise It will be understood that the intensity of the field B depends on the intensity of the field B. 0prepolarise The strength of, in turn, depends on the geometry of the coil 120 and the current applied to the coil 120. In one configuration, the frequency of the magnetic resonance signal may be as high as 200 kHz, although various center frequencies may be envisaged. In one embodiment, a center frequency of about 40 kHz may be used.

[0084] JPEG2025512134000003.jpg84166

[0085] The use of magnetic core 620 supports strong field amplification of the magnetic field used for polarization. As will be appreciated from above, the resulting B 0prepolarise An increase in the magnetic resonance signal linearly increased the available magnetic resonance signal. The presence of the magnetic core 620 also increases the receiver sensitivity of the coil 110. This in turn improves signal reception without appreciably increasing the noise of the signal. In one embodiment, this technique is used with other peripherals required for MRI (e.g., gradient coils) that have the effect of the core on the gradient fields accounted for in the design or post-processing, or have their effect countered with a correction procedure (e.g., a shim-like procedure to the gradient fields).

[0086] By using the core, the magnetic field (or its sensitivity to magnetic resonance signals) generated by the coil can be directed to the desired volume of interest. Conversely, the use of the core can prevent fields from being generated in areas outside the coil that are not of interest to the magnetic resonance measurement or that may even be a potential source of interference. The core may therefore be used to shape the magnetic field / sensitivity of the coil and may be used as a magnetic shield or may be extended to act as a magnetic shield. In this way, the prepolarization field is directed only to the designated measurement volume of the patient / coil, thereby reducing potentially harmful or at least undesirable fringe fields remaining in the examination room.

[0087] Although particular shapes of the coil 110 and magnetic core 620 are shown in FIG. 6, it will be understood that these shapes are not required and other coil and core shapes may be selected instead. More generally, but without wishing to be bound by theory, the magnetic core 620 is advantageously located at a distance from the field of interest 640 that is less than the maximum dimension of the receiving coil (i.e., the coil diameter in the example of FIG. 6). In this way, an increase in RF sensitivity may be realized in the field of interest when compared to a coil of equal shape but omitting the ferromagnetic core. In other embodiments, the shape of the magnetic core may be non-cylindrical. For example, in one embodiment, the core may have a frustum shape with the smaller of the two circular faces of the frustum facing the volume of interest. In other embodiments, the core shape is not symmetrical or rotationally symmetrical. In one embodiment, the shape of the core is not uniform but is the result of a numerical design optimization process of the core and / or coil shapes, which may maximize the magnetic field strength per unit sqrt watt realized by the coil and core combination in the volume of interest.

[0088] In addition to the magnetic core 620 shown in FIG. 6, in one embodiment, the dual-purpose coil 110 further comprises a shield 650. In the embodiment shown in FIG. 6, the shield is provided to surround the solenoid 610 and the magnetic core 620 on all sides that do not face the region of interest 630, i.e., the area where the patient is positioned for the nuclear magnetic resonance examination. In this way, leakage of fringe magnetic fields can be suppressed. The shield 650 can be made of a soft ferromagnetic material. Although FIG. 6 shows a continuous structure of the shield surrounding the solenoid 610 on three sides, it will be understood that this structure is not required. Instead, the shield 650 can be provided on fewer sides, for example only on the side of the solenoid 610 opposite the region of interest 630, or only on one or more sides surrounding the solenoid 610. In another embodiment, the shield 650 may consist of multiple parts that are bonded together or that do not directly fix and join the individual components of the shield 650 to each other, but are held in a fixed relationship to each other by fixing elements. At the same time, the use of magnetic cores below and to the sides of the receiving coil also creates directional selectivity of the signal, projecting a field in a given volume of interest, where only the field lines coming from dipoles approximately above / within the coil create a magnetic flux change at the center of the coil, thereby inducing a voltage in the coil. This can be understood through the reciprocal field of the coil. A coil that creates only a small field at a certain location also means that a dipole at that location cannot induce a significant voltage in the coil for the same dipole amplitude.

[0089] In another embodiment, an additional, thinner shield is provided that surrounds the shield 650 shown in Figure 6 on the sides and underneath. This additional shield further reduces stray magnetic fields outside the volume of interest, improving the safety of the system. In one embodiment, this additional shield extends vertically higher than the patient-facing surface of the coil 110, so that stray magnetic fields on the sides of the volume of interest are also shielded. In one embodiment, the additional shield is removable from the coil 110 and / or the shield 650.

[0090] Magnetic shielding is especially important in portable NMR systems because there is less control over the surrounding environment. Many jurisdictions impose a non-safety zone around operating NMR systems. This zone is typically defined by the contour where the fringe fields fall off to 5 Gauss (0.5 mT). Laws aside, magnetic field strengths greater than 5 Gauss can adversely affect surgical implants such as pacemakers, which is, of course, undesirable.

[0091] Embodiments of NMR systems that use electromagnets, as opposed to permanent superconducting or permanent magnets, are advantageous because they have no fringe fields in the deactivated state, which makes them easier to transport.

[0092] FIG. 8A is a schematic diagram showing a half cross-sectional view of an axisymmetric NMR system according to one embodiment. The NMR system 700 includes a dual-purpose coil 710 with a magnetic core 770, a counter coil 720, an active noise cancellation (ANC) coil 730, and a magnetic structure 760. In the illustrated embodiment, the NMR system 700 further includes a bed 740 within the imaging volume 750 on which a patient undergoing an examination can lie. FIG. 8B shows a three-dimensional isometric view of an NMR system according to one embodiment similar to that shown in FIG. 8A, but with a flap 760b extending along a portion of the length of the patient's bed, with the patient's head positioned in the central field of the dual-purpose coil 710. FIG. 8C shows a detailed view of a portion of the NMR system shown in FIG. 8A. As shown in FIG. 8A, the center of the magnetic core 770 is hollow approximately halfway down the thickness of the core. Filling this region with a high magnetic permeability material would provide little or no benefit to the operation of the dual-purpose coil and would add undesirable weight. In one embodiment, this space is filled with a material that has a thermal conductivity that exceeds that of the material of the magnetic core 770 .

[0093] [Dual-purpose coil 710] As already detailed above, the combination coil 710 is operated under DC conditions for pre-polarization and under AC conditions for receiving NMR signals. However, because the optimal parameters for the coil for pre-polarization and for receiving NMR signals may vary, in one embodiment, the combination coil 710 is divided into multiple sections (L1, L2, L3, L4, and L5), as shown in FIG.

[0094] FIG. 9B shows a circuit for connecting the portions L1-L4 of another dual-purpose coil 710 to a power supply and a preamplifier, respectively. The dual-purpose coil of FIG. 9B does not include L5. However, the inductor L5 of the dual-purpose coil shown in FIG. 9A is connected in the same manner as inductors L1-L3 of FIG. 9B. The circuits of FIGS. 9A and 9B are connected to a power supply (not shown) through a noise reduction circuit. Noise reduction circuits are known to the skilled reader, for example, from FIG. 8 of U.S. Pat. No. 4,906,931 and the accompanying description, which are incorporated herein by reference in their entirety.

[0095] In some embodiments, each coil (L1, L2, L3, L4) comprising the dual-purpose coil 710 has comparable parameters (e.g., radius, number of turns, material, etc.). It will be understood that the invention is not limited thereto and that alternatively, some or all of the coils may have different parameters, such as different radii, different thicknesses along the axis of rotational symmetry, different number of turns, different materials, etc. Additionally, the height / wire thickness of the individual windings may vary within a coil. Furthermore, while a dual-purpose coil having four or five sub-coils is shown in the embodiments, it is emphasized that the invention is not limited thereto and that, alternatively, a different number of sub-coils may be used to form the dual-purpose coil.

[0096] [DC operation] As shown in the circuit of FIG. 9B, each coil (L1, L2, L3, L4) is connected in series via a respective diode (D1, D2, D3) so that DC current can flow through each of the coils. The coils (L1, L2, L3, L4) can therefore function to generate a static prepolarization field. The mutual inductive coupling between the coils (L1, L2, L3, L4) is typically strong, depending on the coil design. For DC operation, the capacitors C1-C8 can be considered to have high impedance or even represent an open circuit.

[0097] [AC operation] For AC operation, the impedance of the diodes for very small signals received by the inductors is so large that they can be considered to be substantially non-conductive. As will be apparent to those skilled in the art, for AC operation, the circuit shown in FIG. 9B is modified to one in which each inductor (e.g., L4) is connected in series with two capacitors on either side of it (C4 and C5 for conductor L4), and the resulting four series CLC circuits are connected in parallel with each other. This parallel connection allows the effective inductance of the dual-purpose coil 710 during AC reception to be much lower than its effective inductance during DC operation. This increases the self-resonant frequency of the coil, but does not reduce the sensitivity of the receiving section, even when the diodes are open.

[0098] Part or all of one of the coils (e.g., L4) is used to receive the NMR signal. A capacitor (not shown) is connected in parallel with this coil to make it resonate at the operating frequency. This adjustment requires the self-resonance of the dual-purpose coil 710 to be greater than the operating frequency. A skilled reader will be able to meet this requirement.

[0099] The NMR signal received by the coil (i.e., L4) is then sent through a DC blocking capacitor to a preamplifier. Optionally, the preamplifier is noise matched and the NMR signal is further sent through a filtering network and / or a blanking switch. In one embodiment, a switched attenuation or detuning circuit is connected between the common node of the capacitors (C1, C2, C3, C4) and the common node of the capacitors (C5, C6, C7, C8) to shorten the ring-down of the current induced in the coils (L1, L2, L3, L4) during the excitation pulse.

[0100] Part or all of one of the coils (e.g., L3) may also be used for excitation. If the same part of the coil is used for excitation and sensing (at L4 in the illustrated example), a T / R switch is connected to the same port of the coil (i.e., L4) via a DC blocking capacitor, and a controller is used to control the reverse bias applied to the diode so that conduction during the excitation pulse is minimized. In another embodiment, a different part of the coil than that used for receiving is used for excitation, and circuitry is provided that isolates the transmit chain from the coil in receive mode and the preamplifier from the coil in transmit mode. In an alternative embodiment, circuitry that actively cancels the transmit signal bleed-through to the receive chain is used to isolate the transmit chain and the preamplifier.

[0101] [Counter coil 720] In some embodiments, the NMR system includes a counter coil 720 to reduce the magnetic footprint of the dual-purpose coil 710 by supplementing the fringe fields generated by the dual-purpose coil 710. In the embodiment shown in Figure 8A, the counter coil 720 is wrapped concentrically around the dual-purpose coil 710.

[0102] Preferably, the magnetic dipole moment of the counter coil 720 is configured to be the same or similar, even if of opposite sign, to that of the dual-purpose coil 710. This can be achieved through control of the radius and number of turns. In some use cases, however, the upper limit of this radius is limited by physical constraints (e.g., available space). The counter coil 720 can then, at least partially, cancel the fringe magnetic field generated by the dual-purpose coil 710 when equal and opposite currents are applied to the coils 710, 720.

[0103] In some implementations, the magnetic field profile of the dual-purpose coil 710 varies with time. Preferably, the magnetic field generated by the counter coil 720 exhibits the same time-varying profile to effectively complement the fringe fields. In one embodiment, the dual-purpose coil 710 and the counter coil 720 are connected in series such that the current amplitude supplied to each coil is the same at any time. In such an implementation, the input power is divided between the coils 710, 720 according to the relative resistance of the coils 710, 720. Since greater magnetization improves signal quality, the resistance of the counter coil 720 is preferably minimized so that the power drawn by the dual-purpose coil 710 can be maximized. In one embodiment, the wire used for the windings of the counter coil 720 is larger in cross section than that of the dual-purpose coil 710, thereby reducing the resistance of the counter coil. In an alternative embodiment, the counter coil 720 has fewer windings than the dual-purpose coil 710, thereby reducing its resistance. In another embodiment, the counter coil 720 has fewer windings than the dual purpose coil 710 and a larger wire cross section than the dual purpose coil 710 .

[0104] In one embodiment, the counter coil 720 and the dual-purpose coil 710 are concentric with each other such that their dipole vectors coincide. The coils 710, 720 may have the same or different shapes in plan view. For example, the coils 710, 720 may be circular, polygonal, etc.

[0105] It is further desirable that the region of interest 750 is shielded from the magnetic field generated by the countercoil 720 so as not to diminish the prepolarization field for the measurement. An embodiment of the NMR system comprises a magnetic structure 760 configured to shape the magnetic field profile of the countercoil 720 in a manner that directs magnetic flux away from the region of interest 750. The magnetic structure 760 is described in further detail below.

[0106] [Magnetic Structure 760] The magnetic structure 760 comprises various components that are magnetically coupled to the magnetic core 770 of the dual-purpose coil 710. As will be appreciated, the magnetic structure 760 concentrates the magnetic flux within its volume, thereby affecting the path of the magnetic flux lines, the field of view, or the free space in another part of the NMR device. As mentioned above, the counter coil 720 is configured to generate a static magnetic field substantially equal and opposite to the static magnetic field generated by the dual-purpose coil 710 at the fringes of the static magnetic field generated by the dual-purpose coil 710, thereby canceling or at least reducing the fringes of the static magnetic field generated by the dual-purpose coil 710. By providing a low resistance magnetic flux path, in particular through the plate 760a shown below the counter coil 720 and the flap 760b shown outside the counter coil 720 when viewed relative to the region of interest 750, the magnetic flux lines generated by the counter coil 720 are focused towards and outside the counter coil 720 when viewed relative to the region of interest 750 (ROI) / imaging volume. In this way, the counter coil 720 can reduce / create a field that counters the fringe fields created by the dual-purpose coil 710, while the negative / catastrophic effects of the fields created by the counter coil 720 in the ROI are reduced to an acceptable level.

[0107] In one embodiment, the magnetic structure 760 further comprises an annular inner magnetic structure 760e provided inside the upper casing 760c, one or more casing sides 760d, and / or ANC coil 730 to help reduce the generation of static magnetic fields therebelow. The use of any of the components 760a-760e individually without any of the other components 760a-760e is also expressly contemplated. It is also expressly contemplated that components 760a and 760b may be used in combination with each other as described above without components 760c-760e. In one embodiment, component 760c connects the magnetic core 770 to the other components of the magnetic structure 760.

[0108] Figure 8C shows a further beneficial modification to the magnetic structure 760 of one embodiment. As shown in Figure 8C, an additional annular core 760f is provided inside the counter coil 720. This additional annular core 760f serves to further reduce the field strength generated by the counter coil 720 in the ROI.

[0109] 10A and 10B show cross-sectional views of 2D axisymmetric simulations of the unsafe zone (as defined by the 5 G or 0.5 mT contours) for an embodiment of an NMR system with (FIG. 10B) and without (FIG. 10A) a counter coil 720. In FIG. 10A, the magnetic structure 760 consists only of a magnetic core 770 with a dual-purpose coil, while in FIG. 10B, the magnetic structure 760 further comprises components 760a-760e and a counter coil 720.

[0110] As can be seen by comparing Figures 10A and 10B, the approximate radius of the unsafe zone is reduced from about 90 cm to 70 cm. In both simulations, the total power consumption was fixed at 4000 kW, with the counter coil at less than 300 W. An embodiment of the NMR system comprising the counter coil 720 and magnetic structure 760 can therefore significantly reduce the fringe field footprint, at a small cost to power consumption.

[0111] 8B, it can be seen that flap portions 760b of the magnetic structure extend along at least a portion of each long edge of the patient's bed 740. Each flap portion 760b may extend along the entire long edge of the bed, or along only a portion of it. In a preferred embodiment, the flap portions 760b that do not extend along the entire length of the patient's bed 740 extend the same length on either side of the vertical centerline of the dual purpose coil 710.

[0112] In some embodiments, each flap portion 760b is foldable between an extended position for use (as shown in FIG. 8B) and a stowed position for transport to provide patient access to the patient's bed. As can be inferred from FIG. 8A, the width of the system is reduced from about 1.2 m to about 1 m in the stowed position. This improves maneuverability of the system through doorways, hallways, elevators, etc., which is important for a portable NMR system. In the extended position, the flap portions 760b extend upward and outward from the patient's bed 740.

[0113] In some embodiments, each flap portion 760b weighs less than 10 kg, and more preferably less than 5 kg, such that a single operator can move the flap 760b between the extended and stowed positions.

[0114] Embodiments of an NMR system that include a magnetic structure 760 with a foldable flap portion 760b may require a gap between the flap portion 760b and the remainder of the magnetic structure 760 to allow for folding movement. In embodiments with a gap width of less than 1 cm, the gap has an acceptable impact on performance. In some embodiments, there may be gaps between other portions of the magnetic structure 760. In some embodiments, these gaps are less than 1 cm wide and the gaps have an acceptable impact on performance.

[0115] In a particular embodiment, the magnetic structure 760 is made of a ferrite compound. However, other materials with high magnetic permeability are possible. Parts of the magnetic structure, for example the magnetic core 770 with the dual coil 710 and the magnetic structure 760, are made of the same or different magnetic materials. For example, in one embodiment, different parts 760a-760e of the magnetic core 770 may be made of different ferrite materials. More generally, the parts 760a-760e of the magnetic core 770 may include any combination of the same or different high magnetic permeability materials. In this way, the choice of material, and in particular the magnetic permeability of the material, can be matched to the technical requirements of the component in which the material is used. In the selection of materials, therefore, more cost-effective materials can be used for components with less stringent requirements for magnetic permeability.

[0116] In some embodiments, the magnetic structure 760 is a collection of parts that are bonded together for use. This reduces manufacturing costs and increases ease of assembly. Each part may in turn be made up of one or more subparts. In one embodiment, the core 770 with the dual-purpose coil 710 is made up of multiple components that are, for example, electrically insulated from one another. In one embodiment, the insulating layers between the individual components extend in a radially extending plane that includes the longitudinal axis of the dual-purpose coil 710. Optionally, parts (or subparts) of the magnetic structure 760 and / or the core 770 are coated with an electrically resistive layer to limit conduction between adjacent components, thereby limiting the size of possible eddy current loops.

[0117] [Coil cooling and temperature control] Place B 0prepolarise The strength of the prepolarization field B 0prepolarise increases with the amplitude of the current used in the dual-purpose coil 110 when used to generate field B. 0prepolarise Assuming that a higher strength of β results in a larger available difference in the occupancy of the spin states and, consequently, a larger available magnetic resonance signal, it is advisable to use the highest possible current in the dual-purpose coil 110 and to obtain a B 0prepolarise As a result, it is desirable to generate a field using a high current. The use of high current leads to a temperature rise in the dual purpose coil 110 due to resistive heating. In some embodiments, the dual purpose coil 110 can benefit from active cooling, further increasing the DC that can be applied to the dual purpose coil 110. Other coils, such as the counter coil 720, the ANC coil 730, and the PCB used for the imaging pulses presented in some embodiments, can also benefit from thermal management.

[0118] In particular, if resistive heating is left uncontrolled over an imaging measurement, the temperature of the coils can increase to a point where it adversely affects performance. Other components of the system, such as the magnetic structure 760 or bed 740, can also be adversely affected by and / or even exacerbate the heating.

[0119] In some embodiments, the permeability and power loss of the magnetic structure 760 and core 770 are temperature dependent. In embodiments having a ferrite compound based magnetic structure 760 or core 770, the relatively low thermal conductivity of the ferrite compound limits heat flow to the extent that the magnetic structure 760 and core 770 act as a heat flow barrier. Of course, it is paramount that the temperature felt by the patient on the bed 740 be maintained at a safe and comfortable level. As such, a temperature management system that uses both passive and / or active cooling elements is desirable.

[0120] In some embodiments, the NMR system includes one or more heat exchangers as active electrical and / or mechanical elements to remove heat away from the coils and bed. Such active elements may be positioned such that EMI noise or eddy current backfields generated by the extractors are at least partially shielded by the magnetic structure 760. In one embodiment, the heat extractors are surrounded or partially surrounded by the magnetic structure 760. In one embodiment, the heat extractors are located directly below the bed, inside the electronics box, optionally on an aluminum plate.

[0121] Preferably, the active elements are placed far enough away from the coil and bed that they have minimal impact on the NMR output signal. However, the dimensions of the NMR device may make this impossible. In such circumstances, it is advantageous to place the active elements in a magnetically shielded enclosure, such as the enclosure formed by the magnetic components 760c and 760d, which will be described in more detail below. One or more passive cooling elements with high thermal conductivity are then arranged to direct heat towards the active elements. Examples of passive cooling elements are provided below, but materials with high thermal conductivity but negligible electrical conductivity and magnetic permeability may be used.

[0122] FIG. 11 is a schematic diagram of an interdigitated configuration of a passive cooling element 1010 and a magnetic structure 1020 of an NMR system, according to one embodiment. The magnetic structure 1020 comprises a number of radial extensions 1022 extending from a star-shaped magnetic core 1024, and a tail portion 1026 extending upwardly, each of which is defined by a radially distal end of the portion 1022. The passive cooling element 1010 is positioned to occupy the space between the portions 1022 and 1026. The interdigitated configuration promotes efficient heat transfer away from the bed (i.e., downward). Although an active cooling element is not shown in FIGS. 10A and 10B, it will be understood that the passive cooling element extends from a coil (e.g., a dual-purpose coil) to the active element and away from the bed. Although a configuration having eight extensions 1022 is shown in FIG. 11, it will be understood that a different number of extensions and associated tail portions 1026 are possible and envisioned as well.

[0123] Although the discrete nature of the magnetic structure 1020 at its periphery concentrates magnetic flux in the region directly above the portion 1026, it has been found that this localized flux concentration does not result in prepolarization field non-uniformities in the ROI or imaging volume 750 that adversely affect performance. The performance of the combined receive coil 710 is similarly unaffected.

[0124] Also, the interdigitated configuration shown in FIG. 11 makes more efficient use of thermally conductive materials in terms of volume, weight, and cost than an approach in which the magnetic structure and combined coil are entirely enclosed.

[0125] FIG. 12 is a schematic diagram showing a cross section of a dual-purpose coil 1110 according to one embodiment. As already mentioned, the dual-purpose coil 1110 comprises a number of separate coils 1120. As shown, these coils are spaced apart from each other by respective passive cooling layers 1130. These layers help to draw heat away from sections of the coil that would otherwise require thermal conduction through the coil itself. The thickness of these layers is selected according to the required heat drain and the volume allocated to the dual-purpose coil 1110 in the NMR system. When used in combination with the embodiment shown in FIG. 11, the layer 1130 is connected in a highly thermally conductive manner to the thermally conductive structure of FIG. 11.

[0126] In some embodiments, the passive cooling element 1010 of FIG. 11 and / or layer 1130 of FIG. 12 are comprised of a thermally conductive ceramic, such as aluminum nitride, boron nitride, alumina, or a combination thereof.

[0127] FIG. 13 shows the simulated temperature profile of the bed, magnetic structure, and dual-purpose coil 1110 over seven measurement cycles for an NMR system operating with a 3860 W polarizing coil, a 140 W counter coil, and the combined configuration shown in FIG. 11 and further comprising a passive cooling layer 1130 shown in FIG. 12. The simulation also considered the duty cycle of the system since the NMR system is assumed to operate with downtime between measurements. As shown, after three measurement cycles the system reaches equilibrium and the temperature reaches the same maximum value in each subsequent measurement cycle. The room temperature is assumed to be 24° C. The temperature of the surface of the insulation covering the bed reaches a maximum of 38° C, which is within the safety limits, while the temperature inside the coil 1110 and magnetic structure remains within an acceptable range. FIG. 14 shows the temperature distribution of the simulated NMR system shown in FIG. 13 at time=3.725 hr, which coincides with the maximum temperature of the last simulated measurement cycle. Although a hot spot appears to have developed inside the dual-purpose coil 1110, the temperature has dropped, demonstrating the effectiveness of the configurations in Figures 11 and 12 in removing heat from the system.

[0128] Other methods of actively cooling the dual-purpose coil are contemplated. In one embodiment, the conductors forming the windings of the dual-purpose coil 110 are located inside a tube, preferably coaxially. A cooling fluid is pumped through the tube to remove excess heat. In another embodiment, the windings of the coil are made of conductive tubing, such as copper tubing, through whose lumen the cooling fluid can be pumped / flowed. In yet another embodiment, the dual-purpose coil 110 may be submerged in a fluid-tight container through which the cooling fluid circulates. The windings of the dual-purpose coil 110 may be spaced apart from one another to allow the cooling fluid to permeate between the windings. In one embodiment, any coolant that evaporates is captured, cooled / condensed, and returned to a liquid state before being resupplied to the dual-purpose coil 110 and the container holding the coolant.

[0129] [Active Noise Cancellation (ANC) Coil 730] In some embodiments, the NMR system includes one or more ANC coils 730. In the embodiment shown in FIG. 8A, the system includes one ANC coil 730. The ANC coil 730 is configured to be minimally sensitive to the region of interest 750 while being maximally sensitive to distant sources of noise (i.e., sources in the far field). It will be appreciated that since the ANC coil 730 shown in FIG. 8A surrounds the dual-purpose coil 710, the ANC coil 730 can detect unwanted external noise emitted from any direction around the dual-purpose coil 710 from which the dual-purpose coil 710 is prone to pick up noise. The background noise measured by the ANC coil 730 can then be subtracted from the measured NMR signal taking into account a sensitivity factor to more accurately determine the actual NMR signal.

[0130] Noise in this context will be understood to include any coherent electromagnetic interference (EMI) source detectable by the ANC coil 730 and the detection portion of the dual-purpose coil 710. However, it does not include incoherent noise sources, such as Johnson noise from the coil itself.

[0131] The relative sensitivity of a coil to signal and noise sources can be simulated or measured as a function of the location of the source. The background noise scaling factor α is a function of the noise source (S b The scaling factor β of the signal from the patient is given by the ratio of the mutual fields of the ANC coils 730 at the position of the patient (S p ) is given by the relative sensitivity of the ANC coil to the NMR signal generated by the

[0132] That is, for the simplified case of three coils, two detection coils (in this example designated main and head, which may be, for example, coils L3 and L4 in FIG. 9B, but any available detection coil may be used by itself or in combination with other detection coils, such as dual-purpose coils) and one ANC coil 730:

number

[0133] For effective noise removal from the measured NMR signal, the system is ideally designed such that α>1 and β<1. That is, the ANC coil 730 is configured to be more sensitive to background noise (far-field sources) than the detection coil 710, but less sensitive to the NMR signal from the patient. In the above equations, the ratios α and β are expressed with respect to the sum of the sensitivities in the denominator, but it will be understood that the sensitivity of the dual-purpose coil 710 or other detection coils may be expressed differently in this context according to the electrical configuration of the dual-purpose coil or any other detection coil.

[0134] In the case of a single background noise source, the background noise is removed by subtracting the signal detected by the ANC coil 730 from the signals detected by the main and head detection coils.

number

[0135] JPEG2025512134000006.jpg39166

[0136] The fractional increase is therefore minimized by configuring the ANC coil 730 to have a much smaller resistance than the sensing coil of the dual-purpose coil 710. A passive cooling element 1010, described in more detail below, may also be used to keep the temperature of the ANC coil 730 to a minimum. Furthermore, note that the fractional increase in incoherent noise when having an ANC coil 730 in a system is reduced as the operating temperature of the dual-purpose coil 710 increases (all else remaining constant).

[0137] When comparing an NMR system with the ANC coil 730 to an ideal system with no background noise and limited only by Johnson noise (of the detection coil), the signal-to-noise ratio (SNR) of the ideal system is

number

[0138] However, because the background EMI signal power is expected to be much larger than the additive Johnson noise power of the ANC coil, removal of the coherent background noise leads to an overall improvement in the SNR of the NMR signal. In most cases, the 1-10% loss in SNR caused by the Johnson noise of the ANC coil 730 will be less than the loss caused by the background EMI noise.

[0139] Similar to the counter coil 720, the magnetic core 770 with the dual-purpose coil at least partially shields the ANC coil 730 from the region of interest 750, and the flap portion 760b helps direct magnetic flux from the imaging volume 750, away from the ANC coil 730. In one embodiment, the ANC coil 730 is positioned directly below the flap portion 760b and outside the magnetic structure 760. Optionally, the ANC coil 730 is wrapped concentrically around the dual-purpose coil 710. The far-field sensitivity of the ANC coil 730 then has a similar spatial profile to the dual-purpose coil 710, although preferably by a factor α>1.

[0140] FIG. 15 shows the results of a simulation of the coil sensitivity spatial profile according to an embodiment with two detection coils (a dual-purpose coil and an additional coil placed in the simulated patient's head) and one ANC coil 730. The ratio of the SNR from these two detection coils and from the ANC coil, i.e., 1 / α, is plotted. As desired, in the region of interest 750, the ratio is much greater than 1, but outside the volume, the ratio drops sharply to a value much less than 1. This shows that placing the ANC coil 730 outside the magnetic structure and directly under the flap portion 760b is a very effective configuration. This is also shown by the fact that in the regions shielded from the ANC coil 730 with the shielded housing under the core 770 and element 760c, the ratio is much greater than 1, indicating that the ANC coil 730 is much less sensitive to the fields in these regions.

[0141] In some embodiments, a gap exists adjacent to the flap portion 760b of the magnetic structure. This gap acts as a channel through which magnetic flux from the region of interest 750 impinges on the ANC coil 730, thereby increasing the sensitivity ratio, β. Despite this increase in β caused by the gap, it has been found that the overall reduction in SNR of the NMR signal is tolerable, at least for gaps of 1 cm or less. Figures 16A and 16B show the sensitivity ratio of an NMR signal according to one embodiment without (Figure 16A) and with (Figure 16B) such a gap.

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[0142] In some embodiments, the ANC coil 730 comprises its own magnetic core that is distinct from the magnetic structure, while in other embodiments, the annular portion 760e of the magnetic structure forms part of the magnetic core associated with the ANC coil 730, as described further herein.

[0143] In some embodiments, the NMR system includes multiple ANC coils (not shown). This helps improve EMI noise rejection since each ANC coil can then be configured to detect different coherent sources (e.g., noise sources at different locations). For example, it is envisioned that such ANC coils are positioned at different locations within the NMR system and / or at different inclinations relative to the primary axis of the dual-purpose coil 710 to improve their collective sensitivity to far-field noise sources. In some embodiments, the position and inclination of each of the multiple ANC coils may be adjustable to optimize their collective sensitivity to observed far-field noise sources.

[0144] [Reduction of eddy currents] As already detailed above, the operating sequence of the NMR system includes activating and deactivating (ramping down) the dual-purpose coil 710 and switching the prepolarization field on and off. Preferably, the ramping down of the prepolarization field to zero occurs fast enough that NMR measurements can be performed with minimal loss of spin polarization. However, rapid changes in the magnetic field (above about 10 T / s) can cause peripheral nerve stimulation, causing pain or discomfort to the patient. The maximum ramp rate for a given maximum strength of the magnetic field in the patient is known to the skilled reader. In any case, intermediate ramp-down rates still cause significant backaction fields due to eddy currents induced in conductive objects close to the coil 710 during switching. The presence of conductive objects in an NMR system is unavoidable, since the system is electronically controlled. It is preferable to keep the electronics away from the coil, but this is at least to some extent incompatible with a compact and portable NMR system.

[0145] The backaction field, or backfield, of these eddy currents is (1) the amplitude of the backfield in the region of interest, which is the field that carries, B 0measurement and / or (2) the gradient of the backfield in any voxel of the region of interest is comparable to or larger than the surrounding or shimmed inhomogeneity of the holding field with the relevant gradient applied, can adversely affect NMR measurements because (1) the backfield modifies the magnetic resonance frequency in an unpredictable and time-varying manner, thereby reducing the accuracy of the NMR results, and (2) the backfield gradient can cause spin dephasing and loss of magnetization.

[0146] In some embodiments, the NMR system further comprises a passive coil disposed between the dual coil 710 and the conductive object in which eddy currents are expected to occur. The passive coil may simply be a conductive loop positioned between the source of the magnetic field and the object that would conduct the generated eddy currents in the absence of the passive coil. Due to the backfield generated by the current induced in the passive coil, the object that would conduct the generated eddy currents in the absence of the passive coil undergoes a change in which the field strength decreases. Thus, the eddy currents induced in this object are less when compared to the scenario in which the passive coil is not present. From the perspective of the conductive object, the passive coil therefore slows down the change in field amplitude from the dual coil caused by switching, and the amplitude of the eddy currents induced in the conductive object is thereby reduced. Instead, it is preferable to induce eddy currents in a passive coil, since eddy currents can be more easily controlled. Various methods of controlling eddy currents are envisioned. As one example, eddy currents can be more quickly attenuated up to and including an open circuit configuration through control of the resistance of the passive coil. In some embodiments, the system includes two or more of these passive coils.

[0147] In one embodiment, the coil is a metal strip, as opposed to a thinner wire construction, to increase the cross section of the coil and therefore the effectiveness of its inductive response. Examples of metallic materials include copper, aluminum, or alloys thereof.

[0148] In some embodiments, the NMR system further comprises an additional active counter coil to mitigate eddy currents. A configuration in which the additional active counter coil and the dual-purpose coil 710 are connected in series is particularly advantageous because the time variations of the magnetic field are well cancelled out. Eddy currents induced in conductive objects in these reduced field regions are subsequently reduced. In embodiments having a magnetic structure 760 and an active counter coil, no additional passive coil is required. However, a combination of a counter coil 720 and a passive coil is particularly effective in mitigating the effects of backaction fields from eddy currents and is employed in one embodiment. In one embodiment, the counter coil 720 described above is used to additionally provide the functionality of an active counter coil.

[0149] In some embodiments, the metal enclosure 1702 is used to shield magnetic fields and noise within the detection bandwidth of the NMR system between the electronics and the receiver of the dual coil. In one embodiment, the metal enclosure 1702 is therefore located directly below the dual coil 710 and houses the system electronics therein. In one embodiment, the metal enclosure 1702 has at least one opening to provide a partial enclosure for the system electronics. Backaction fields are expected from eddy currents that arise on the surface of this metal enclosure. The metal enclosure may be in contact with the magnetic structure 760 or may be spaced from the magnetic structure by a gap. In some embodiments, as shown in FIG. 17, the magnetic structure 760 at least partially covers the metal enclosure 1702 and directs magnetic flux around the enclosure through the casing side 760d and the casing bottom 760g, but not into the enclosure. This helps to suppress induction of eddy currents in the enclosure 1702.

[0150] 18A and 18B show the eddy current backfield arising from the top casing of the metal enclosure 1702, respectively, without and with the passive coil 1802. As can be seen by comparing the scale bars, the maximum backfield in the region of interest 750 without the passive coil is about 2 mT, and with the passive coil it is about 50 μT, representing about a 20-fold reduction.

[0151] FIG. 19 shows a simulated eddy current backfield in an NMR system without counter coil 720 or passive coil 1802. As shown, in one embodiment, the magnetic structure 760 alone can reduce the backfield in the region of interest to less than 4 μT. The magnetic structure 760 (including the core 770) shields the region of interest 750 from the backfield and redirects the field elsewhere. It can be seen that a properly configured magnetic structure 760 can be more effective in reducing the backfield in the region of interest compared to using the counter coil 720 or passive coil 1802. By comparing FIG. 19 with FIG. 18A and FIG. 18B, it can be seen that the magnetic structure component 760e, which is present in the embodiment shown in the former but not in the latter, is particularly impactful in achieving this reduction.

[0152] As mentioned above, in some embodiments, the windings of the dual-purpose coil 110 are spaced apart from one another. While this is advantageous from a cooling standpoint (as discussed above), such spacing between the windings is also used to advantage in the non-cooling dual-purpose coil 110 because spacing the windings apart reduces parasitic capacitance within the windings of the coil. This increases the self-resonant frequency of the dual-purpose coil 110, allowing a larger number of windings to be used while still maintaining the self-resonant frequency of the dual-purpose coil 110 above the frequency of the nuclear magnetic resonance signal.

[0153] In a further embodiment, the windings of the coil are embedded in a solid temperature conducting material. One side of this material, for example the side facing away from the volume of interest, may be connected to a heat sink, preferably an actively cooled heat sink. In one embodiment, solid state cooling is used to provide such active cooling.

[0154] The self-resonant frequency of a coil is affected by the parasitic capacitance of the coil. Reducing ε' of the dielectric between the turns of the coil reduces the parasitic capacitance of the coil. Reducing ε" also further reduces the electrical losses in the dielectric medium and their associated noise. Thus, in one embodiment, a high quality cooling medium filling the coil or the solid material in which the coil is embedded is selected to reduce the electrical losses in the coil.

[0155] Although specific configurations have been described, these configurations are presented by way of example only and are not intended to limit the scope of protection. The inventive concepts described herein may also be implemented in various other forms. In addition, various omissions, substitutions and modifications can be made to the specific implementations described herein without departing from the scope of protection defined in the following claims.

Claims

1. A first coil that generates a first static magnetic field, A second coil that generates a second static magnetic field perpendicular to the first static magnetic field, A third coil that generates a radio wave magnetic field at a precession frequency determined by the second static magnetic field, Equipped with, The first coil is, A ferromagnetic core and The coil windings are wrapped around the ferromagnetic core, Equipped with, The ferromagnetic core is configured to concentrate the magnetic flux generated by the winding of the coil in a region adjacent to the winding of the coil. Nuclear magnetic resonance system.

2. The ferromagnetic core comprises a plurality of ferromagnetic components that are electrically insulated from each other. The system according to claim 1.

3. The first coil has a first diameter and a first vertical axis, The aforementioned system, The fourth coil further comprises having a fourth diameter and a fourth vertical axis. The first vertical axis and the fourth vertical axis substantially coincide, and the fourth diameter is larger than the first diameter. The system according to claim 1 or 2.

4. The coil is an active noise cancellation coil. The system according to claim 3.

5. The first coil and the fourth coil are arranged such that, at a distance from the vertical axis greater than the diameter of the fourth coil, the fourth coil generates a field that cancels out the field generated by the first coil. The system according to claim 3.

6. The first coil and the fourth coil are further configured to activate and / or deactivate simultaneously, The system according to claim 5.

7. Further comprising a magnetic flux guide component positioned away from the vertical axis to guide the magnetic flux generated by the fourth coil, The system according to claim 5.

8. A receiving coil and A coil having superior sensitivity to noise sources in long-range fields compared to the aforementioned receiving coil, Furthermore, The system according to claim 1.

9. The ferromagnetic material is a soft ferromagnetic material. The system according to claim 1.

10. A system comprising a coil in a ferromagnetic core, wherein the ferromagnetic core has a plurality of spokes, each spoke extending below the coil, from the center below the coil to a radial end at a radial distance exceeding the diameter of the coil, and some or each of the plurality of spokes further extending upward from the radial end, forming a portion that extends upward outside the maximum diameter of the coil. The gaps between the spokes and / or the upward-extending portions include a material having a thermal conductivity exceeding that of the ferromagnetic core. The system according to claim 1.

11. The coil comprises a vertical axis and a plurality of layers stacked on the vertical axis and / or a plurality of windings adjacent to each other in the radial direction, and the coil further comprises a material having a thermal conductivity exceeding the thermal conductivity of the windings of the coil and / or the thermal conductivity of the ferromagnetic core. The system according to claim 1.

12. The invention further comprises an active or passive coil between the first coil and a conductive component capable of conducting eddy currents, the active coil being used to reduce eddy currents within the conductive component. The system according to claim 1.

13. The system comprises a receiving coil, The receiving coil has a vertical axis and is highly sensitive to radio wave signals emitted from the region of interest and electromagnetic noise emitted outside the region of interest. The system further comprises a noise cancellation coil having a vertical axis substantially coinciding with the vertical axis of the receiving coil, The noise cancellation coil is highly sensitive to electromagnetic noise emitted outside the region of interest, and the system is configured to sense the noise emitted outside the region of interest and remove the noise from the signal received by the receiving coil using a predetermined magnification. The system according to claim 1.

14. Further comprising a magnetic shield positioned to reduce the sensitivity of the noise cancellation coil to signals emitted from a region of interest, The system according to claim 11.

15. The sensitivity of the noise cancellation coil to a noise source emanating from outside the region of interest is higher than the sensitivity of the receiving coil. The system according to claim 11 or 12.

16. The first coil comprises a plurality of conductors, all of the plurality of conductors are used when generating the first static magnetic field, and only a subset of the conductors of the plurality of inductors, or only one conductor, is used to receive the nuclear magnetic resonance signal. The system according to claim 1.

17. The system is a single-sided system, and the region adjacent to the coil winding is the patient side of the system. The system according to claim 1.

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