How to identify coils and coil geometries

Abstract

A method of identifying a geometry of a coil comprising obtaining first information identifying a flow function associated with the coil, determining the flow function based on the first information, determining a position of a first contour in the flow function, and determining a position of a first coil track based on the position of the first contour and a track width, the track width of the first coil being identified by a value of the flow function, the method further comprising discretizing the first coil track into a plurality of sub-tracks and generating second information identifying a geometry of the coil.

Description

[Technical field]

[0001] SUMMARY OF THE DISCLOSURE The embodiments described herein generally relate to coils, methods for determining the geometry of a coil, and NMR systems including the coils. [Background technology]

[0002] Conventional NMR scanners include a donut-shaped housing that contains multiple coils. In use, the subject to be imaged is positioned within the donut, i.e., surrounded by the housing. The housing therefore undesirably limits the size of the subject that can be imaged.

[0003] Another approach uses a one-sided NMR scanner. In this configuration, the coils that control the magnetic field are located in close proximity to each other. Previous approaches to designing coils and generating arbitrary magnetic fields have proven inadequate when the coils are in close proximity to each other, as required in a one-sided NMR scanner. Given this, there is a need for new methods to design and produce coils that generate user-defined magnetic fields.

[0004] The configurations of the embodiments will be more fully understood and appreciated from the following detailed description, which is given by way of example only and is taken in conjunction with the drawings in which:

[0005] In the drawings, like reference numbers refer to like components. [Brief description of the drawings]

[0006] [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] FIG. 1 illustrates a perspective view of a single-sided NMR scanner, according to an example. [Figure 8B] FIG. 2 is a side view of a single-sided NMR scanner, according to the present example. [Figure 8C] FIG. 2 illustrates an NMR system according to another example. [Figure 9A] FIG. 1 is a perspective view of a coil component according to an example. [Figure 9B] 1 is a cross-sectional view of a coil component according to an example. [Figure 10] 10A and 10B illustrate a method for iteratively designing a coil, according to an example, and a method for using direct computation to design a coil, according to an example. [Figure 11] 1 is a diagram of 2D basis functions, according to an example. [Figure 12] FIG. 13 illustrates a method for generating coil shapes from stream functions, according to an example. [Figure 13A] FIG. 1 is a diagram showing an illustrative example of a stream function. [Figure 13B] FIG. 13 illustrates the effect of setting a constant track width in contour space, according to an illustrative example. [Figure 13C] FIG. 13 is a diagram illustrating an example of a coil shape. [Figure 14]FIG. 14A is a plan view of a track according to an illustrative example; FIG. 14B is a plan view of a track discretized according to a first approach according to an illustrative example; FIG. 14C is a diagram of an example of a Litz wire-like structure according to an example; FIG. 14D is a diagram of first and third sub-tracks of a Litz wire-like configuration according to an example; and FIG. 14E is a diagram of a top and bottom layer of a Printed Circuit Board (PCB) coil according to an example. [Figure 15] Figure 15A shows a top and bottom layer of a coil with connected tracks according to an illustrative example, Figure 15B shows a close-up view of the connections between adjacent coils according to an illustrative example, and Figure 15C shows a perspective view of a coil fabricated with Litz wire sub-tracks according to an illustrative example. [Figure 16] 1 illustrates a method of manufacturing a coil, according to an example. [Figure 17] FIG. 1 illustrates an implementation of an apparatus, according to an example. [Figure 18] FIG. 1 illustrates a method for designing multiple coils capable of performing Magic Angle Field Spinning (MAFS), according to an example. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0007] According to a first aspect, there is provided a method of identifying a geometry of a coil, the method comprising obtaining first information identifying a flow function associated with the coil, determining the flow function based on the first information, determining a position of a first contour of the flow function, and determining a position of a first coil track based on the position of the first contour and a track width, the track width of the first coil being identified by a value of the flow function. The method further comprises discretizing the first coil track into a number of sub-tracks and generating second information identifying a geometry of the coil.

[0008] In one embodiment, the second information is used to produce a coil.

[0009] In one embodiment, the method further includes producing the coil based on second information specifying a shape of the coil.

[0010] In one embodiment, determining the location of the first contour line includes identifying a point in the plane of the coil that is associated with a first value of the stream function.

[0011] In one embodiment, the method further includes determining a first value of the stream function by determining a difference between a minimum value of the stream function and a maximum value of the stream function, determining a contour distance based on the difference divided by a plurality of contour lines, and determining a first value of the stream function based on the contour distance and a second value of the stream function associated with a second contour line.

[0012] In one embodiment, determining the first value based on the contour distance and the second value includes adding the contour distance to the second value or subtracting the contour distance from the second value.

[0013] In one embodiment, the method further includes determining a position of a second contour line based on the stream function, and determining a position of a second coil track based on the position of the second contour line and a track width, the track width being determined by a value of the stream function such that the first coil track and the second coil track occupy the same range of stream function values.

[0014] In one embodiment, the coil comprises a first coil track and a second coil track.

[0015] In one embodiment, the first coil track and the second coil track occupy the same range of stream function values, but also different geometric distances / widths in the plane of the coils.

[0016] In one embodiment, determining the position of the first coil track includes identifying a position in the plane of the coil associated with a value of the stream function that is 1) greater than or equal to the first value of the stream function minus half the track width, and 2) less than or equal to the first value of the stream function plus half the track width.

[0017] In one embodiment, the plurality of sub-tracks comprises a first sub-track having a first width and a second sub-track having a second width, the first sub-track and the second sub-track being coplanar.

[0018] In one embodiment, discretizing the first coil into a plurality of sub-tracks includes determining a first width of the first sub-track such that the first sub-track has an impedance equal to an impedance of the second sub-track.

[0019] In one embodiment, discretizing the first coil into a plurality of sub-tracks includes determining a first width of the first sub-track such that the first sub-track has an impedance similar to an impedance of the second sub-track.

[0020] In one embodiment, discretizing the first coil track into a plurality of sub-tracks includes producing Litz wire in a first layer and a second layer of the coil.

[0021] In one embodiment, the first layer comprises a first sub-track and the second layer comprises a third sub-track associated with the first sub-track.

[0022] In one embodiment, the information identifying the stream function associated with the coil includes a plurality of coefficients, each coefficient of the plurality of coefficients being associated with a basis function in a plurality of basis functions for modeling the stream function associated with the coil.

[0023] In one embodiment, the stream function associated with a coil is represented by a weighted sum of the basis fields, where the weight applied to each basis field is a coefficient associated with the respective basis field.

[0024] In one embodiment, the method further includes a basis set, the basis set including each basis function of the plurality of basis functions, the basis set having the following mathematical properties: 1) each basis function equals 0 at the ends of the design space, 2) divergence of J=0, where J is the current density, and 3) each basis function has a finite value in the design space.

[0025] In one embodiment, the multiple basis functions model a Fourier-Bessel basis set.

[0026] In one embodiment, obtaining information identifying a flow function associated with the coil includes obtaining a first set of coefficient values ​​including a plurality of coefficients, each coefficient of the first set of coefficients being associated with a basis function for modeling a flow function associated with the coil, determining a value of the first coefficient based on the first set of coefficients, determining a magnetic field in a volume of interest generated by the coil having a flow function represented by the first set of coefficient values ​​and the first coefficients, comparing the magnetic field to a target magnetic field and generating a difference metric, and determining a loss metric based on the difference metric. Then, in response to determining that the difference metric satisfies a stopping condition, the method further includes generating information identifying the flow function based on the first set of coefficient values ​​and the first coefficient.

[0027] In one embodiment, the loss metric includes a difference metric and a regularization term.

[0028] In one embodiment, the normalization terms include 1) a power normalization term that is calculated based on the power efficiency of the coil, and 2) an inductance normalization term that is calculated based on the inductance associated with the coil.

[0029] In one embodiment, determining that the difference metric satisfies a stopping condition includes determining that the difference metric is less than a threshold.

[0030] In one embodiment, determining that the difference metric meets a stopping condition includes determining that the loss metric is not improving.

[0031] In one embodiment, obtaining information identifying the stream function further includes adjusting the first set of coefficient values ​​in response to determining that the difference metric does not satisfy a stopping condition, recalculating the first coefficients, determining a second magnetic field in the volume of interest, and comparing the second magnetic field to the target magnetic field.

[0032] In one embodiment, the model of the system includes a coil and a second coil, and determining a value of the first coefficient includes determining a value of the first coefficient such that a voltage induced by the coil in the second coil is zero.

[0033] In one embodiment, the first coefficient is:

number

[0034] In one embodiment, the voltages v0, v i is calculated from first principles.

[0035] In one embodiment, the voltages v0, v i is calculated using an electromagnetic simulator.

[0036] In one embodiment, determining the magnetic field in the volume of interest includes simulating the coils and the current distribution associated with the stream functions using a computational electromagnetics solver. Preferably, determining the magnetic field in the volume of interest includes simulating the magnetic field in the volume of interest for each basis function using a Finite Element Model (FEM) solver and combining the simulated magnetic fields according to a first set of coefficients. In one embodiment, the FEM solver simulates the magnetic field in the volume of interest in the presence of eddy currents in the magnetic field generated by linear magnetic and / or ferromagnetic materials and / or current distributions.

[0037] In one embodiment, the computational electromagnetics solver is a FEM model or an FDTD simulator of Maxwell's equations.

[0038] In one embodiment, the second information specifying the geometry of the coil includes information indicative of the position of the conductors in the plane of the coil.

[0039] In one embodiment, the conductors comprise copper traces / tracks.

[0040] In one embodiment, a method of generating a multiple coil geometry for magic angle magnetic field rotation is provided, the multiple coils comprising a first coil, a second coil, and a third coil, the method including identifying a first coil geometry as described in any of the previous embodiments such that the first coil generates a first magnetic field. The method further includes obtaining a second plurality of coefficients and a third plurality of coefficients, the second plurality of coefficients modeling a second stream function associated with the second coil and the third plurality of coefficients modeling a third stream function associated with the third coil, determining a second magnetic field associated with the second plurality of coefficients and a third magnetic field associated with the third plurality of coefficients, determining a metric based on the first magnetic field, the second magnetic field, and the third magnetic field, and adjusting a value of one of the second plurality of coefficients and / or a second value of one of the second coefficients of the third plurality of coefficients in response to determining that the metric is less than a predetermined threshold.

[0041] In one embodiment, the second plurality of coefficients and the third plurality of coefficients are used to produce a second coil and a third coil.

[0042] In one embodiment, the method further includes creating a second coil and a third coil based on the second plurality of coefficients and the third plurality of coefficients.

[0043] In one embodiment, determining a metric based on the first magnetic field, the second magnetic field, and the third magnetic field includes calculating a second metric based on the first magnetic field, the second magnetic field, and the third magnetic field at each point in space within the volume of interest, and summing the second metrics associated with each point in space to obtain the metric.

[0044] JPEG2025512122000003.jpg43166

[0045] In one embodiment, the sum is a weighted sum.

[0046] In one embodiment, calculating the second metric further includes adding a regularization term.

[0047] In one embodiment, the normalization terms include a power normalization term calculated based on the power efficiency of the first, second, and / or third coils, and / or an inductance normalization term calculated based on the inductance of the first, second, and / or third coils.

[0048] According to a second aspect, there is provided a non-transitory computer readable medium comprising computer program instructions suitable for execution by a processor, the instructions being configured, when executed by the processor, to perform a method according to any of the preceding embodiments.

[0049] According to a third aspect, there is provided a coil obtainable by the method according to any preceding embodiment.

[0050] According to a fourth aspect, there is provided a coil component comprising a coil according to the third aspect and a second coil, both of which are coplanar and coaxial, and the coil component is integrally formed.

[0051] In one embodiment, the coil component comprises a printed circuit board, the printed circuit board comprising a first layer and a second layer, the first layer comprising the coil and the second layer comprising the second coil.

[0052] In one embodiment, there is provided a nuclear magnetic resonance system comprising a coil component according to any of the above embodiments and a receive coil, the receive coil being coaxial with the coil component.

[0053] According to a fifth aspect, there is provided a coil component comprising a first coil and a second coil, both of which are coplanar and coaxial, and the coil component is integrally formed.

[0054] In one embodiment, the first coil is separate from the second coil.

[0055] In one embodiment, the first coil is detuned from the second coil.

[0056] In one embodiment, the coil components are produced in a single process such that the first coil and the second coil are subject to the same manufacturing variations.

[0057] In one embodiment, the coil component is cylindrical.

[0058] In one embodiment, the coil component comprises a printed circuit board, the printed circuit board comprising a first layer and a second layer, the first layer comprising the first coil, and the second layer comprising the second coil.

[0059] In one embodiment, there is provided an NMR system comprising a coil component according to any of the above embodiments and a receive coil, the receive coil being coaxial with the coil component.

[0060] In one embodiment, the first coil is for generating a B0 field.

[0061] In one embodiment, the coil components include all of the coils for operating the NMR system except for the receive coils.

[0062] In one embodiment, all coils for operating the NMR system are 1,tx , G x , G y , and G z Includes coil.

[0063] In one embodiment, the first coil and the second coil in the coil component are separate from the receive coil.

[0064] According to a sixth aspect, there is provided a method of generating a multiple coil shape for magic angle magnetic field rotation, the multiple coils including a first coil, a second coil, and a third coil, the method including obtaining a first plurality of coefficients associated with the first coil, the first coil generating a first magnetic field; obtaining a second plurality of coefficients and a third plurality of coefficients, the second plurality of coefficients modeling a second flow function associated with the second coil and the third plurality of coefficients modeling a third flow function associated with the third coil; determining a second magnetic field associated with the second plurality of coefficients and a third magnetic field associated with the third plurality of coefficients; determining a metric based on the first magnetic field, the second magnetic field, and the third magnetic field; and adjusting a value of one coefficient of the second plurality of coefficients and / or a second value of a second coefficient of one of the third plurality of coefficients in response to determining that the metric is less than a predetermined threshold.

[0065] In one embodiment, the first set of coefficients, the second plurality of coefficients, and the third plurality of coefficients are used to produce a second coil and a third coil.

[0066] In one embodiment, the method further includes creating the first coil, the second coil, and the third coil based on the first plurality of coefficients, the second plurality of coefficients, and the third plurality of coefficients.

[0067] According to a seventh aspect, there is provided a method for determining a shape of a coil, the method including: obtaining a first set of coefficient values ​​including a plurality of coefficients, each coefficient of the first set of coefficients being associated with a basis function for modeling a stream function associated with the coil; determining a value of the first coefficient based on the first set of coefficients; determining a magnetic field in a volume of interest generated by the coil having a stream function represented by the first set of coefficient values ​​and the first coefficients; comparing the magnetic field with a target magnetic field; and in response to a determination that the magnetic field is similar to the target magnetic field, generating information identifying the stream function based on the first set of coefficient values ​​and the first coefficient.

[0068] According to an eighth aspect, a nuclear magnetic resonance coil is provided, the nuclear magnetic resonance coil being configured in a first mode to receive and conduct current at a drive port for generating 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.

[0069] In one embodiment, the coil includes an electrical circuit that electrically isolates the drive port and the receive port from each other.

[0070] In one embodiment, the coil is dimensioned to generate a static magnetic field in a volume of interest that enables acquisition of NMR signals throughout the torso depth of an adult subject positioned prone or supine on the surface of the coil.

[0071] In one embodiment, the electrical circuit is a passive circuit.

[0072] 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.

[0073] In one embodiment, the NMR signal voltage received by the coil is amplified by a factor of 20 or less, preferably 5 or less.

[0074] In one embodiment, the coil is non-resonant.

[0075] 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.

[0076] According to another embodiment, there is provided a nuclear magnetic resonance coil comprising a patient side adjacent to which a patient is positioned during a magnetic resonance examination and a soft ferromagnetic shield on at least one side of the coil other than the patient side. In one embodiment, the nuclear magnetic resonance coil is as previously described.

[0077] 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.

[0078] In one embodiment, the magnetic resonance device further comprises a static magnetic field coil configured to generate a static magnetic field 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.

[0079] In this embodiment, the region of interest is located on the patient side and at a distance of 20 cm from the front surface of the nuclear magnetic resonance coil facing the patient.

[0080] In an alternative embodiment, the static magnetic field 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 may be 200 mT at a given location in the coil's volume of interest, while the static magnetic field generated by the permanent magnet may 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 may be such that the sum of the fields is inclined 45 to 85 degrees to the direction of the static magnetic field generated by the permanent magnet. In this example, the sum of the fields has a magnitude significantly larger than the field generated by the magnetic resonance coil alone, thus achieving a high degree of prepolarization. It will be understood that even in this example, the magnetization is preceded around the direction of the static magnetic field generated during the acquisition of the magnetic resonance signal. The shape of the magnetic resonance coil is not affected by the choice of the permanent magnet for generating the measurement static magnetic field, but the magnetic resonance coil is, on the other hand, highly sensitive to this signal.

[0081] 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.

[0082] In one embodiment, the ferromagnetic material is ferrite.

[0083] In one embodiment, suitable ferromagnetic materials include 10 -7Ferromagnetic materials having any possible combination of these properties are suitable for use in the embodiments.

[0084] 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.

[0085] In one embodiment, the cessation of application of the current to the coil and the generation of the second static magnetic field are performed such that an adiabatic switch 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 switch 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.

[0086] In one embodiment, nuclear magnetic resonance signals are acquired simultaneously with the application of the radio frequency magnetic field.

[0087] 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.

[0088] 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.

[0089] In one embodiment, the coil includes a pump that pumps liquid into the tube.

[0090] In one embodiment, the conductor is a tube and the liquid flows within the lumen of the tube.

[0091] In one embodiment, at least a portion of the conductor is disposed within a fluid-tight container containing a liquid.

[0092] In one embodiment, the conductors form windings, the windings being spaced apart from one another to allow liquid to circulate between adjacent windings.

[0093] 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.

[0094] 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.

[0095] Coil 130 is in field B. 0measurement The B1 field creates a B1 RF magnetic field at the precession frequency generated by 0prepolarise and place B 0measurement As is the case with known NMR RF coils, the B1 field can be activated and deactivated.

[0096] 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 142 occupied by a subject 140, such as a patient, that is the subject of the NMR measurements.

[0097] Although a particular configuration of system 100 is shown in Figure 1, the spatial arrangement of magnets and coils shown in Figure 1 is not required so long as the magnetic fields generated are substantially mutually perpendicular to one another. Although dual-purpose coil 110 and coils 120 and 130 are shown in Figure 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 system 100 shown, and that dual-purpose coil 110 and some or all of coils 120 and 130 may instead be provided in a single unit. In one example, coils 120 and 130 may form part of a single PCB.

[0098] 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 0prepolarise The dual-purpose coil 110 is energized so that a field is generated as shown in FIG. 0prepolarise The 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 of the spin species, but longer than the T1 relaxation time of one of the spin species. As is well known, a static magnetic field applied to the 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 are0prepolarise and B. 0measurement extend at right angles 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.

[0099] 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. 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 0prepolarise Direction and field B 0measurement 3C. The magnetization vectors are tilted in a plane perpendicular to both the directions of the magnetization vectors. Because the dual-purpose coil 110 created fields in the directions shown in FIG. 3A, the dual-purpose coil 110 is sensitive to the precession magnetization vectors shown in FIG. 3C. Therefore, the dual-purpose coil 110 can be used to detect the magnetic resonance signals thus generated. 0prepolariseBy 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.

[0100] 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.

[0101] 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

[0102] 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 0prepolarise The difference between 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

[0103] B, as described above in connection with step 220 0prepolarise From B 0measurement The change in the difference in the spin population states that inevitably occurs when switching the static magnetic field between 0prepolarise The advantage realized in the net magnetization in field strength is that, as described above in connection with step S220, 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 After obtaining a higher signal intensity than would normally be obtained from spins exposed to 0measurement As a result, NMR measurements are performed at the B 0measurement From B 0measurement For a period of time following the switch to B 0measurement This can be accomplished using a B1 field at a resonant frequency determined by 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 perform the adiabatic switching so as to complete it as quickly as possible without violating the regulation of stimulation to the peripheral nerve. In one embodiment, B 0prepolarise 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 0prepolarise is completely ramped down, and B 0measurement The shortest longitudinal relaxation time T1 of all spin species for which NMR signals are acquired from the time when In another embodiment, NMR signals are acquired for a shorter period of time after the shortest longitudinal relaxation time T1 of all spin species for which NMR signals are acquired has passed, until the end of a longer or longest longitudinal relaxation time T1 of another of the spin species for which NMR signals are acquired. Then, the field B 0prepolarise A further measurement cycle can be started by reactivating , again prepolarizing the spins to be examined.

[0104] In an alternative embodiment, the switching of the fields is not adiabatic. Instead, field B 0prepolarise is B 0prepolarise to a non-zero value in a time window that does not allow for redistribution of the spin population that was generated at the full strength of field B. 0prepolarise When is small enough, say 10% of its full strength, B 0measurement is rapidly activated, while B 0prepolarise is also rapidly deactivated. 0prepolarise The magnitude of magnetization generated by is not only not maintained by the magnetization, but also by the B without the need to apply a B excitation pulse. 0prepolarise This will initiate procedures for:

[0105] 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.

[0106] 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, and prevents 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, the dual-purpose coil 110 can be provided with direct current 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 resistance 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.

[0107] 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 a further embodiment, further cross-coupled diodes may be provided to ground the terminals of each of the two capacitors 560 connected to the receive ports. In this way, low amplitude magnetic resonance signals received using the dual-purpose coil 110 reach the receive port, while any higher amplitude signal spikes that may be caused by the prepolarization current applied to the dual-purpose coil 110 or by current switching are driven to ground.

[0108] 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.

[0109] 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.

[0110] 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).

[0111] 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.

[0112] In another embodiment, the diodes 730 and 740 may be replaced by simple active switches that can interrupt the connection from the coil 110 to the driver 710 .

[0113] [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 the simulation are also shown in FIG. 6. The high B value (about 0.5T) shows the rough outline of the ferrite material. The current density in the coil windings is shown in the inset image, highlighting the cross-section of the individual copper windings. In one embodiment, the individual copper windings are themselves made of Litz cross-section / Litz wire.

[0114] The magnetic core 620 acts as a flux concentrator to concentrate the magnetic flux generated by the solenoid 610 in a region 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 flux concentrator that concentrates the magnetic flux generated by the solenoid 610 in a region 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. 0prepolariseand the high frequency magnetic field generated in the field of interest 630 during use of the dual coil 110. The frequency of the magnetic resonance signal generated in the field of interest 630 is 0measurement It will be understood that the intensity of the field B depends on the intensity of the field B. 0measurement 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 10 kHz may be used.

[0115] JPEG2025512122000005.jpg75166

[0116] 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 field 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 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).

[0117] 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 potential sources 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.

[0118] 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 amplification of 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 current / sensitivity realized by the coil and core combination in the volume of interest.

[0119] 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 be made 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 very little field in a certain location also means that dipoles at that location cannot induce a voltage in the coil.

[0120] In another embodiment, an additional, thinner shield is provided on the sides and underneath, surrounding the shield 650 shown in FIG. 6. This additional shield further reduces stray magnetic fields outside the volume of interest, improving safety of the system. In one embodiment, this additional shield extends vertically higher than the patient-facing surface of the coil 110, so that the stay 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.

[0121] [Coil cooling] 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 0prepolarise Given that a higher field strength results in a larger available difference in the occupancy of 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 obtain a B 0prepolarise As a result, it is desirable to generate a DC field by using a high current that leads to an increase in temperature of the dual purpose coil 110. In some embodiments, the dual purpose coil 110 can benefit from active cooling, further increasing the DC current that can be applied to the dual purpose coil 110.

[0122] Various 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, preferably coaxially with, a tube. A cooling fluid is pumped through the tube to remove excess heat. In another embodiment, the windings of the coil are comprised of a conductive tube, such as a copper tube, through whose lumen a cooling fluid may be pumped / flowed. In yet another embodiment, the dual-purpose coil 110 may be submerged in a fluid-tight container through which a cooling fluid is circulated. The windings of the dual-purpose coil 110 may be spaced apart from one another to allow for 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.

[0123] 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.

[0124] 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.

[0125] 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.

[0126] In known Nuclear Magnetic Resonance (NMR) scanners, coils responsible for generating the magnetic fields used in the imaging process surround a Volume of Interest (VOI). In known examples, a toroidal / donut shaped housing is provided that contains various coils, within which the Volume of Interest (e.g., a patient) is positioned for imaging. These NMR scanners have various disadvantages. For example, in these machines, the patient is surrounded by the housing during the imaging process. This can be uncomfortable for the patient being imaged. Furthermore, the openings in the machine housing also impose physical limits on the size of the object that can be imaged by the NMR scanner. To address these issues, single-sided NMR scanners are provided. Single-sided NMR scanners include Magnetic Resonance Imaging (MRI) scanners that do not require hardware that occupies a space surrounding the desired imaging volume, but instead, and preferably, only require the scanner hardware to be located on one side of the imaging volume / patient.

[0127] FIG. 8A shows a perspective view of a single-sided NMR scanner according to one example. FIG. 8B shows a side view of a single-sided NMR scanner according to this example. In particular, in FIGS. 8A and 8B, the single-sided NMR scanner 150 comprises a surface 151 on which a volume of interest (e.g., a human / patient 152 is located) is placed for imaging. The single-sided NMR scanner 150 also comprises an enclosure 153 that, during use, is located directly below the surface 151. The enclosure 153 comprises an NMR system that is used to image the volume of interest (VOI).

[0128] 8C shows an NMR system according to another example. The NMR system 800 includes a receive, Rx, coil 810 for observing a volume of interest (VOI).

[0129] The system 800 further comprises two coils, the first coil 120 and the second coil 130. As mentioned above, the first coil 120 generates a static magnetic field, B 0measurement The first coil 120 is energized and de-energized to generate a field B 0measurement can be activated and deactivated accordingly.

[0130] The second coil 130 is in field B. 0measurement The B1 field creates a B1 RF magnetic field at the precession frequency generated by 0measurement As is the case with known NMR RF coils, the B1 field can be activated and deactivated.

[0131] Two coils, a first coil 120 and a second coil 130, generate a magnetic field B 0measurement and B1 are configured to be generated in a space occupied by a subject 140, such as a patient 152, that is the subject of an NMR measurement.

[0132] In one example, the NMR system 800 is operated in a manner similar to the NMR system described above in connection with Figures 1-7. However, it will be understood that the Rx coil 810 is not dual purpose, and therefore the Rx coil 810 in this example is not operated in a manner that would have generated a prepolarization field.

[0133] In one example, the coils that influence the evolution of the spin state / net magnetization vector are provided in a single integrally formed component / unit, optionally on a single Printed Circuit Board (PCB).

[0134] 8C, the coils that affect the evolution of the spin state / net magnetization vector include a first coil 120 and a second coil 130. In one example, the conductors forming the first coil 120 are coplanar, the conductors forming the second coil 130 are coplanar, the first coil 120 and the second coil 130 are coaxial, and the first coil 120 and the second coil 130 are implemented as a single unit (i.e., as part of a single integrally formed structure).

[0135] In the illustrative example of Figure 8C, the first coil 120 and the second coil 130 are shown as separate objects spaced apart by a gap, however, it is emphasized that in embodiments described herein, this gap is omitted.

[0136] FIG. 9A shows a perspective view of a coil component according to an example. In the example of FIG. 9A, the coil component 370 comprises a first coil 120 and a second coil 130. In the example of FIG. 9A, the first coil 120 is a planar coil (i.e., the copper tracks forming the coil lie in a plane) and the second coil 130 is also a planar coil. The first coil 120 and the second coil 130 are circular (i.e., the shape of the surface including the tracks of the first coil 120 and the second coil 130 is circular). Optionally, the first coil 120 and the second coil 130 have the same radius and the coil component 370 has a cylindrical shape. In the coil component 370 shown in FIG. 9A, the axis of the first coil 120 is aligned with the axis of the second coil 130. However, the second coil 130 is offset along this axis with respect to the first coil. This arrangement can be seen in FIG. 9A where the second coil 130 is located directly below the first coil 120 in the coil component 370.

[0137] During use, the coils in the coil component 370 are energized and de-energized to affect the evolution of the spin state / net magnetization vector of the sample being analyzed.

[0138] In one example, the coil component 370 is implemented using a printed circuit board (PCB), optionally a multi-layer PCB. As is known in the art, a printed circuit board (PCB) is a laminated sandwich structure that includes conductive and insulating layers.

[0139] FIG. 9B shows a cross section of a coil component according to an example. In the example of FIG. 9B, the coil component 370 is a two-layer structure. The coil component 370 comprises a first solder mask layer 381, a first copper layer 382 forming the first coil 120, and a core layer 380. The first copper layer 382 is disposed on a first surface of the first core layer 380 between the core layer 380 and the first solder mask layer 381. The coil component 370 further comprises a second copper layer 384 forming the second coil 130, and a second solder mask layer 385. The second copper layer 384 is disposed on a surface opposite the first copper layer 382 between the core layer 380 and the second solder mask layer 385.

[0140] As is known in the art, a solder mask layer is a thin lacquer-like layer of polymer that is applied to the copper traces of a PCB for, among other things, protection against oxidation. Similarly, a core layer is a rigid substrate that is laminated with copper on one or both sides. The core layer is made of an insulating material. In one example, the core layer is an FR4 core layer or a ceramic layer.

[0141] In the example shown in Figures 9A and 9B, the coil component 370 comprises two coils, the first coil 120 and the second coil 130. In other examples, the coil component comprises at least one coil. In other examples, the coil component comprises three or more coils. In a particular example, the coil component 370 comprises four coils, a first coil for B0 field generation, a second coil for B1 RF excitation, a third coil for G1 gradient field generation, and a fourth coil for G2 gradient field generation. In other examples, the coil component 370 comprises any number of gradient coils, but at least three substantially gradient coils.

[0142] 9B illustrates a particular implementation of the coil component 370 using a two-layer PCB, but it is emphasized that other configurations may also be used. For example, the coil component 370 may instead be implemented using a multi-layer PCB that further comprises one or more "prepreg" layers. As is known in the art, a "prepreg" layer is a dielectric material sandwiched between two core layers that is used to provide insulation between the copper tracks located on the two core layers.

[0143] In these examples, the coil component 370, which comprises one or more coils used to control the spin state of the sample being imaged, comprises coaxial coils in an integrally formed component, each coil having a coplanar track. In conventional MRI devices, the coils used to control the spin state of the sample (i.e., B0, B 1,tx , G x , G y , and G z Coils) are positioned around the volume of interest. In conventional MRI scanners, these coils are not integrally formed in a single component and do not have a common axis like coil component 370 shown in Figures 9A and 9B.

[0144] Integrating the coils that control the spin state of the sample during an experiment into a single multi-layer structure makes the coils more compact, reducing production costs. Furthermore, a large distance between the pre-polarization / receive coils and the volume of interest (VOI) reduces the performance of the MRI device. As a result, integrating all the coils that control the spin state into a single component (e.g., a multi-layer PCB) reduces the height and form factor, allowing the pre-polarization / receive coils to be placed closer to the volume of interest (VOI).

[0145] Furthermore, the use of an integrally formed component (e.g., a multi-layer PCB) with the coils allows for a geometrically consistent interaction. As a result, calibration and design considerations are repeatable between coils, and the response obtained from a volume at a fixed location relative to the coil component 370 is consistent. In particular, in conventional NMR systems, the alignment and positioning of the coils relative to each other are mechanically adjusted (and often electrically adjusted) to compensate for manufacturing challenges. On the other hand, providing the spin steering coils in a single integrally formed component (e.g., in a single PCB) means that any manufacturing challenges (e.g., deflections due to mechanical inaccuracies) are imposed equally on all layers / coils. As a result, the decoupling between the coils in the PCB is substantially maintained, along with the alignment (e.g., potential perpendicularity of the coil field, if necessary) and uniformity (e.g., shim alignment relative to B0) of the coils.

[0146] In one example, the coil component 370 includes a receive, Rx, coil and a B 0measurement Coils for field generation and any other coils used in the NMR system. For example, coil component 370 includes, in one embodiment, a B 0measurement It comprises a coil for field generation, a coil for transmitting the B1 field, and a series of B0 gradient coils. Coupling between coils in close proximity is unavoidable.

[0147] In one example, the individual coils in the coil component 370 are untuned coils, i.e., coils that have a self-resonant frequency that does not match any of the operating frequencies of the coils in the coil component 370. In this way, although residual coupling between the coils is unavoidable, such residual coupling does not affect the performance of any active coil or the performance of other coils in the coil component 370 in a manner that would prevent them from functioning in the intended manner. In one embodiment, crosstalk between the coils can be electronically removed using known techniques such as closed loop control, signal pre-distortion, etc. Advantageously, this allows for transmission of Tx of either the gradient / B0 coil.

[0148] In one example, one coil in the coil component 370, and optionally each coil in the coil component 370, is separate from another coil (e.g., an Rx coil in an NMR system). In a further example, each coil in the coil component 370 is combined with the other coils in the coil component 370.

[0149] As mentioned above, one purpose of the coils in the coil component 370 is to control the spin state of a sample in a volume of interest (VOI). As such, the coils generate a magnetic field that interacts with the sample in the volume of interest (VOI). The integration of the coils in a single component provides numerous advantages as discussed above. However, previous approaches to coil generation have not considered placing multiple spin manipulation coils, much less all of the spin manipulation coils, in a single, integrally formed structure due to the need for large B0 fields and the sharing of electromagnets with cores. This has created a need for new ways to design coils that can generate arbitrary, user-defined, magnetic field shapes / profiles in 3D space.

[0150] As discussed above, the integrally formed coil component 370 in the NMR system of Figure 8C enables a single-sided MRI system that has many advantages over conventional MRI system configurations, however, this also introduces challenges and considerations that must be overcome when designing a coil to achieve a particular magnetic field.

[0151] For example, when applying a magnetic field in close proximity to a receive coil (as in the example configuration used in the single-sided NMR system of FIG. 8C), it is important that the spin steering coil is completely isolated (i.e., completely geometrically separated) from other coils in the system (e.g., Rx coil 810 in FIG. 8C). This consideration is less important in standard MRI scanners because the coils are located further apart in these scanners. Therefore, there is also a need for a way to design the field steering coils in a way that reliably produces a particular magnetic field shape / profile in 3D space when the coils are in close proximity to other coils, thereby enabling a single-sided MRI scanner.

[0152] 10A illustrates a method for iteratively designing a coil, according to an example. The method starts at step 301 with obtaining first coefficient values: c1, c2, c3, etc. In the method for designing a coil, the coil shape is (indirectly) related to multiple coefficients. Defining the coil by multiple coefficients allows for more efficient calculations and makes each iteration of the design process numerically easier to compute.

[0153] As will become apparent from the following description, the coefficients indirectly define the geometry of the coil. In the method of Figures 10A and 10B, the coefficients model the contribution of various "assistance functions" to the stream function associated with the coil. The stream function is based on the integral of the current density. The current on the surface of the coil is determined based on partial derivatives of the stream function with respect to the x and y directions. In one example, the current density at a point in the xy plane is obtained from the stream function according to:

number

[0154] Basis functions are selected to model the stream functions associated with the coil. As is well known in the art, a basis function is an element of a particular basis for a function space, and any function in a function space can be represented as a linear combination of basis functions. For example, a function can be approximated by a weighted sum of cosine harmonics, where each cosine harmonic is a different basis function.

[0155] There are several different types of basis that can be used to model the function. However, proper selection of the basis can reduce the number of coefficients required to model the coil. Furthermore, improper selection of a basis for modeling the stream function can lead to stream functions associated with coils that are complex and physically impossible to realize.

[0156] In one example, an arbitrary basis set (including basis functions) is used, provided that the basis set (that models the stream function) has the following requirements: 1) stream function=0 at the boundary of the design space (in this context, the design space is a collection of locations that form the coil, e.g., a circle in the xy plane as presented in the example above, or even a surface like a square or a donut), 2) continuity of J (i.e., divergence of J=0), where J is the current density, and 3) the stream function is limited / finite within the design space.

[0157] In a particular example (used for illustration in the following examples), a Fourier-Bessel basis set is used.

[0158] In this case, each coefficient is associated with a different basis function in the basis set and acts as a weight for that particular basis function. The stream function associated with a coil is a weighted sum of the basis functions, and the weights are the coefficients associated with each basis function.

[0159] It has been found that using a Fourier-Bessel basis set to model the stream function usually results in a solution that can be realized with a circular coil (e.g., as shown in the circular coil component 370 of FIG. 9A). As a result, the solutions explored in the optimization process (where various values ​​of coefficients are tested to design a coil that will generate a particular magnetic field in use) can usually be implemented in practice. It has also been found that modeling the stream function using a Fourier-Bessel basis set requires a reduced number of coefficients (compared to other basis functions) to arrive at a coil shape that achieves the desired magnetic field. Achieving the desired field by controlling the values ​​of a smaller number of coefficients is advantageous because it reduces the number of variables that need to be adjusted to achieve the desired result.

[0160] The Fourier-Bessel basis set (in other words, the basis functions used to model the stream function associated with the coil in this example) is defined by:

number

[0161] The Fourier-Bessel basis set is also referred to herein as the vibration modes (or modes of oscillation) of the circular membrane.

[0162] FIG. 11 shows a diagram of 2D basis functions, according to one example. In particular, FIG. 11 shows plots of 2D, spherical harmonic, basis functions defined on a single circle. Each plot is a different basis function in the basis set. Each basis function is uniquely associated with a different coefficient value.

[0163] For example, the first plot 401 is associated with a first expansion order. As described above, the basis functions are parameterized by a first parameter, m, and a second parameter, n. In one example, the first parameter, m, ranges from 0 to a first predetermined value, m, greater than zero. max The second parameter, n, takes an integer value between 1 and a second predetermined value n. maxEach combination of the first and second parameters is uniquely associated with a different basis function, as may be shown in FIG. 11. Each basis function is also uniquely associated with a coefficient. For example, the first basis function (e.g., m=0 and n=1, i.e., U 01 ) is uniquely associated with the first coefficient (c0) and the second basis function (e.g., m=0 and n=2, i.e., U 02 ) is uniquely associated with the second coefficient (c1). It is emphasized that this association is merely one example, and that the first coefficient (c0) may be associated with different basis functions in different examples. As described in more detail below, in one example, the first coefficient (c0) is associated with a basis function that has a non-zero coupling with nearby coils, such that the first coefficient (c0) cancels the contribution to the flux in the nearby coils, thereby obtaining a coil that is separated from the nearby coils.

[0164] Returning to FIG. 10A, in step 301, a first set of coefficients is obtained. The first set of coefficients (e.g., c1, c2, c3, ... c i ) includes at least one coefficient, and optionally two or more coefficients. The stream function associated with the coil is represented by a weighted sum of basis functions, and the weight applied to each basis function is the coefficient associated with the basis function. As a result, the coefficients used to design the coil in the method of FIG. 10A represent the contribution of each basis function to the stream function. In this manner, each basis function may be referred to as a "support function," in the sense that each basis function provides a contribution (according to its associated coefficient / weight) to the stream function associated with the coil.

[0165] In one example, a first set of coefficients obtained in step 301 is provided to the method (e.g., from user input, by reading coefficients from a file stored in computer memory, by random initialization). After obtaining the first set of coefficients, the method proceeds to step 302.

[0166] As mentioned above, the new architecture of the NMR system (i.e., stack of coils) that allows for a one-sided MRI device also makes the coil design challenging. In particular, it has been found to be advantageous to completely separate the spin steering coil from the other coils in the system (e.g., Rx coil 810 in FIG. 8C).

[0167] Having a spin manipulation coil that is completely separate from the Rx coil can be advantageous for several reasons: When the frequency of the spin manipulation magnetic field is not the receive frequency of the Rx coil (i.e., when the spin manipulation field is out of band), a completely separate spin manipulation coil has the advantage of avoiding loss of coil power due to the spin manipulation coil coupling to the Rx coil and the coupled power being dissipated in the electronics at the output of the Rx coil that prevent power from reaching the Rx port.

[0168] A second advantage is that the separated coils avoid coupling of any in-band noise (from the spin manipulation coil) into the receive chain: in this case, there may not be any desired signal (close to NMR frequencies) transmitted in-band by the spin manipulation coil, but all amplifiers / drivers (including those used to drive the spin manipulation coil) generate some degree of white / broadband noise that may be present in the receive chain unless the spin manipulation coil is separated from the Rx coil.

[0169] In other cases where the spin manipulation field is at the receive frequency / band of the Rx coil (i.e., when the transmission from the spin manipulation coil is in band, such as an on-resonance excitation B1, or any gradient coil that also has a B1-like excitation), separating the coils reduces the amount of B1 coupling with the Rx coil, thereby reducing the amount of interference reaching the Rx port (of the Rx coil) to a level that can be digitally and / or electrically corrected / compensated for.

[0170] Furthermore, even if perfect separation is not achieved, imperfect separation can improve performance and is not easy to obtain.

[0171] Previous approaches to reducing the effects of coupling between the coils have been unsatisfactory. For example, one approach is to use an ultra-low noise power amplifier, which reduces the amount of in-band noise. However, this comes at an increased cost. Another approach is to turn off the receive chain when transmitting. However, this approach does not allow for simultaneous transmission and reception. Furthermore, this approach does not isolate the receive chain from the gradient field and must be turned on when receiving. Another approach is to filter the noise in the received signal. However, high levels of filtering can be complex to implement and are not easy, especially when the undesired frequency bands are close to the desired frequency bands. As a result, new approaches are desirable.

[0172] As will be understood more fully below, the coil design method shown in FIG. 10A provides a method for designing coils that isolates coils from nearby coils. In particular, the method of FIG. 10A involves modifying the stream function to reduce coupling to nearby coils, thereby allowing spin steering coils to be placed in close proximity to other coils, a key requirement for implementing a unilateral MRI system.

[0173] As mentioned above, the coil has multiple coefficients (c0, c1, c2, c3, ... c i ) In order to maintain zero coupling (i.e., non-coupling) to the nearby receive coils, the magnetic flux through the surface perpendicular to the B field of the Rx coil needs to be zero net (i.e., no magnetic flux). In one example, this is achieved by designing the spin-manipulated coil such that its positive contribution to the magnetic flux at the Rx coil (from the spin-manipulated coil being generated) is cancelled by its negative contribution to the magnetic flux at the Rx coil (from the spin-manipulated coil), resulting in zero net magnetic flux and non-coupling. In other words, this is achieved by having the spin-manipulated coil additionally have a canceling contribution to the magnetic flux seen by the receive coil.

[0174] This condition is met when the sum of all induced voltages in the Rx coils is zero. The voltage induced in the Rx coil by the ith fundamental field is the voltage, v i In this case, to achieve non-coupling, the following must be satisfied:

number

[0175] In the method of FIG. 10A, the multiple coefficients used to represent the contribution of each basis function to the stream function associated with the coil being designed are divided into two parts: a first coefficient (c0) and a first set of coefficients (c1, c2, c3, ... c i ) in the method described herein, one arbitrary coefficient (the first coefficient c0) is used to ensure that all induced voltages generated by the spin steering coils sum to zero. In this case, the first set of coefficients (c1, c2, c3, ... c i ) is free in the sense that it can be optimized by the method of FIG. 10A or can be user defined, while the first coefficient (c0) is calculated based on the values ​​of the other coefficients. In a specific example, the first coefficient (c0) is calculated according to:

number

[0176] This solution is mathematically equivalent to taking any general current expansion and showing a rotated basis set in which the coupling to the Rx coil is net zero except for one element. This is done through a change of basis that produces m*n-1 basis elements of null coupling (i.e., for the m,n expansion) and one element of non-zero coupling, whose coefficient can now be set to 0. This consists of a rotation in the space of basis elements, but only in the dimension of the non-zero coupling.

[0177] Setting the first coefficient in this manner provides little or no coupling between the spin steering coil and the Rx coil, allowing the coils to be located very close to each other, which is necessary to realize a unilateral MRI system.

[0178] Following this approach, there is one coefficient that is not available for optimization in the method of FIG. 10A, but instead is derived from the coefficients that are available for optimization. As a result, in step 301, a first set of coefficients (e.g., c1, c2, c3, ... c i After obtaining c), the method proceeds to step 302 where a first coefficient (c) is determined such that the induced voltage in the nearby receive coil is zero. For the avoidance of doubt, it is emphasized that the first coefficient (c) may be associated with any basis function / expansion order. In one example, step 302 includes calculating the first coefficient (c) according to Equation 3 above.

[0179] To calculate the first coefficient (c0) according to equation 3 above, we calculate the voltage induced in the Rx coil by each current distribution in the plane of the coil associated with the support function (or contribution to the stream function, represented by the basis fields) (i.e., v i ) is required.

[0180] In one example, the voltages (e.g., v0, v1, v2, v3, v4, etc.) are pre-calculated prior to performing the method of FIG. 10A. In one example, the voltages are determined using an electromagnetic simulator by constructing a model of an NMR scanner (comprising at least an Rx coil and a spin steering coil in relative positions) and simulating the currents associated with basis fields in the spin steering coil. In one example, the currents on the surface of the coil associated with each respective basis function (modeling the assisting function) are calculated, and a voltage for each basis field is determined based on the respective currents on the surface of the coil. In this example, the simulation is performed multiple times, each time with a different current distribution corresponding to a different basis field, thereby obtaining values ​​for multiple voltages (e.g., v0, v1, v2, v3, v4, etc.) associated with each basis field.

[0181] In an alternative example, the voltage induced in the Rx coil for each basis field is determined analytically from first principles.

[0182] To determine the first coefficient (c0) using Equation 3, first set of coefficients (i.e., c1, c2, c3, … c i ) are also required. These values ​​are obtained in step 301 or in step 307 (described in more detail below). After determining the first coefficient (c0), the method proceeds to step 303.

[0183] Step 303 includes determining the magnetic field in a volume of interest (VOI). In one example, the volume of interest (VOI) includes the space occupied by the subject 140 (e.g., a human) being imaged.

[0184] In a first example, determining the magnetic field in a volume of interest (VOI) includes determining the magnetic field generated in the volume of interest for each basis field current distribution in the coil. Determining the magnetic field in the volume of interest includes modeling the NMR system (including the coil being designed) and the volume of interest in an electromagnetic simulator (also called a computational electromagnetics solver, or electromagnetics solver), applying the basis fields and associated current distributions to the surface of the coil being designed, and calculating the magnetic field in the volume of interest (e.g., by running the electromagnetic solver).

[0185] A variety of computational electromagnetic techniques can be used to determine the magnetic field in a volume of interest based on a set of initial conditions. In one example, a Finite Element Method (FEM) solver is used to solve Maxwell's equations. In another example, a Finite Difference Time Domain (FDTD) solver is used to solve Maxwell's equations.

[0186] JPEG2025512122000010.jpg30166

[0187] JPEG2025512122000011.jpg39166

[0188] Total magnetic field in the volume of interest, B total To determine C, a weighted sum is performed based on the magnetic fields obtained for each evolution order. The weights are determined by the coefficients obtained in steps 302 and 303 (i.e., C 0、 C 2、 C 3、 In one example, after obtaining the fields associated with the different expansion orders, the total field, B total is calculated according to:

number

[0189] In one example, the magnetic field in the volume of interest includes magnetic field vectors (ie, magnitude and direction) calculated at multiple points.

[0190] In previous approaches to calculate the total magnetic field, the current density is used with the Biot-Savart equation to calculate the magnetic field originating from the microcurrent elements, which is then integrated over the entire surface to calculate the total field. However, this approach assumes that there are no other sources of magnetic field activated by creating the current itself, and therefore the total field comes only from the coil. As a result, extra effects caused by conductive materials (e.g., eddy currents in the time-varying magnetic field), soft magnetic materials (e.g., fields from material magnetization), or any other effects cannot be modeled within the optimization method. In contrast to these approaches, by using an electromagnetic solver (e.g., FEM model) that simulates a model of the NMR system (including the coil being designed and any surrounding coils, such as the Rx coil), these effects can be modeled, improving the accuracy of the optimized design in implementation. In one embodiment, the total magnetic field is simulated by simulating the magnetic field obtained from each basis function and combining the resulting partial fields using coefficients.

[0191] Total magnetic field in the volume of interest, B total After determining, the method proceeds to step 304.

[0192] In step 304, a number of coefficients (e.g., C 0、 C 2、 C 3、 C4) and the total magnetic field generated by the associated coil, B total is the target magnetic field, B targetIn one example, the target magnetic field is defined by a user. In one example, the total magnetic field, B total is the target magnetic field, B target The comparison with the total magnetic field, B total The magnetic field vectors associated with the target magnetic field, B, at multiple points in the volume of interest (VOI) are target and comparing it with the magnetic field vector associated with it.

[0193] In step 304, the degree to which the vector field generated based on the current coil configuration matches / approximates the target field is determined. In one example, the total magnetic field, B total is the target magnetic field, B target To compare it with the total magnetic field, B total and the target magnetic field, B target In one example, the difference metric is based on a sum of errors. In a particular example, the difference metric is based on the sum of errors. target and (expected) total magnetic field, B total The method proceeds to step 304A.

[0194] At step 304A, a loss metric is determined. In one example, the loss metric is equal to the difference metric. In a further example, the loss metric includes a loss metric and an additional normalization term. In one example, the additional normalization term includes, but is not limited to, a power normalization term (e.g., calculated based on a power efficiency associated with the coil design) and an inductance normalization term (e.g., calculated based on an inductance associated with the coil design).

[0195] In step 305, it is determined whether the loss metric satisfies a stopping condition. In one example, the stopping condition includes a loss metric below a certain threshold or a loss metric that is not improved (i.e., the historical loss metric of X is similar). If the loss metric satisfies the stopping condition, the method proceeds to step 306.

[0196] In step 306, the method ends. Upon reaching step 306, a set of coefficients (i.e., C 0、 C 2、 C 3、 C4, etc.) has been obtained. How this stream function is used to generate the coil is described below with reference to FIG. 12 and may be performed once step 306 is reached.

[0197] If, in step 305 , it is determined that the loss metric does not satisfy the stopping condition, the method proceeds to step 307 .

[0198] In step 307, a first set of coefficients in the plurality of coefficients (e.g., c1, c2, c3, ... c i ) are adjusted. As mentioned above, the method is free to select the values ​​of the coefficients in the first series of coefficients, but the first coefficient (i.e., c0) is calculated based on the values ​​of the coefficients in the first series of coefficients. In one example, the values ​​of the coefficients in the first series of coefficients (e.g., c1, c2, c3, ... c i ) are adjusted based on the difference metric. In a further example, the coefficient values ​​are adjusted to minimize the difference metric (i.e., the total magnetic field, B total is the target magnetic field, B target (to more closely resemble a linear regression model, in other words, to minimize the sum of errors). In one example, the first set of coefficients is updated using a gradient-based method (e.g., analytical, stochastic, or finite difference). In another example, the set of coefficients is updated using a gradient-free method (e.g., a genetic or annealing algorithm).

[0199] After adjusting the coefficients in step 307, the method proceeds to step 302 where the first coefficient, c0, is calculated as described above. total is the target magnetic field, B target Continue multiple iterations until it resembles

[0200] The method described in Figure 10A represents an iterative approach to determining the current density in the plane of the coil acquiring the target magnetic field, however, it is emphasized that the coefficients used to determine each expansion order into the stream function may also be determined using direct methods (i.e., from direct calculation).

[0201] FIG. 10B illustrates a method for designing a coil using direct computation, according to an example. In the example of FIG. 10B, the stream functions associated with the coil being designed are modeled using a Fourier-Bessel basis set, as described above in connection with FIG. 10A. However, as described above, the basis set could use other basis functions, provided the basis functions satisfy requirements 1-3, as described above in connection with FIG. 10A. The method begins at step 351, where a target magnetic field, B, in a volume of interest is calculated. target and the magnetic field in the volume of interest associated with each basis function. The method proceeds to step 352.

[0202] In step 352, a number of coefficients (i.e., c0, c1, c2, c3, ... c) associated with each respective basis function are i ) is calculated based on the target magnetic field and the magnetic fields associated with each basis function.

[0203] In one example, multiple coefficients (i.e., c0, c1, c2, c3, ... c i) is calculated according to the approach provided in equations 3.3a-3.19 in “Implementation of biplanar coils for magnetic field generation, by Miguel Maria Abreu Condessor, Thesis to obtain the Master of Science Degree in Engineering Physics, Tecnico Lisboa,” available at https: / / fenix.tecnico.ulisboa.pt / downloadFile / 281870113705157 / MScThesis.pdf, which is incorporated herein by reference. Following this approach produces a matrix that can be inverted to determine the coefficients.

[0204] In one example, the matrix is ​​inverted using a matrix inversion approach described in M. Janosek, M. Dressler, V. Petrucha and A. Chirtsov, "Magnetic Calibration System With Interference Compensation," in IEEE Transactions on Magnetics, vol. 55, no. 1, pp. 1-4, Jan. 2019, Art no. 6000104, doi: 10.1109 / TMAG.2018.2874169, which is incorporated herein by reference.

[0205] In one example, the coefficients are determined using Turner's Target Field method as described in R. Turner. A target field approach to optimal coil design. Journal of Physics D: Applied Physics, 19 (8), 1986. URL https: / / doi.org / 10.1088 / 0022-3727 / 19 / 8 / 001, which is incorporated herein by reference.

[0206] JPEG2025512122000013.jpg30166

[0207] The stream function discretization process is a non-optimally defined process, the effects of which can significantly change the expected magnetic field in its 3D vector field distribution, its response to currents at various frequencies, and its interference with adjacent coils via eddy currents. In light of this, a new method for determining coil geometry based on a target stream function associated with a coil is also provided.

[0208] As mentioned above, and as emphasized again here, the stream function (also called the stream function representation of the surface current) is the vector potential of the current stream density and is used to plot contours, also called streamlines, that represent the trajectories of steady current flow. The stream function is a scalar field and is related to the surface current density of the coil. In particular, the stream function is related to the integral of the surface current density.

[0209] One approach to generating the coil geometry is to use a stream function to generate the current tracks. In one approach, the stream function is obtained and the streamlines (i.e., the contours / level sets of the stream function) are determined.

[0210] In previous approaches, the optimal solution is defined as tracks of very thin wire placed on the contours of the stream function. As the width of the tracks increases past a negligible size (e.g., because it is impossible to manufacture very thin wires), the current paths deviate from the desired path (e.g., wide curved tracks have a preferred path through the interior of the curve), and the error between the magnetic field generated by the physically realized coil geometry and the target magnetic field increases. Similar errors occur when the number of tracks is too small to adequately approximate a continuous current distribution, as would be necessary due to excessively large inductance.

[0211] Furthermore, this approach has shown that the coil tracks are subject to skin and proximity effects as higher frequency signals are carried by the tracks. For example, higher frequency signals may arise during the continuous evolution of the field (e.g., by applying a sawtooth wave with successively high slew rates to the track) or during transitions between states. This has the effect of increasing the AC resistance and increasing the current distribution at the ends of the track. In conventional MRI devices, these effects do not usually need to be considered because the distance between the coil and the volume of interest (VOI) is usually large enough that these field distribution effects have negligible impact on the field in the volume of interest (VOI).

[0212] As mentioned above, the method described herein is suitable for designing and generating coil geometries for use in unilateral MRI systems. In unilateral MRI systems, such as the example shown in Figures 8A and 8B, a volume of interest (VOI) is located near the coil. As a result, these field distribution effects have a significant impact on the magnetic field generated in the volume of interest (VOI). In light of this, a method is provided for generating coil geometries that are resilient to these field distribution effects, so that the coil geometries can reliably generate a magnetic field in the volume of interest that is similar to the target magnetic field.

[0213] 12 illustrates a method for generating a coil shape from a stream function, according to one example. The method begins at step 502 with obtaining a stream function associated with the coil being generated.

[0214] In one example, obtaining a stream function associated with the coil includes obtaining a plurality of coefficients (e.g., c0, c1, c2, c3, c4, etc.) representing weights associated with each basis function and determining the stream function by performing a weighted sum of each basis function, where the weights are determined by the values ​​of the coefficients associated with the respective basis function. The method proceeds to step 503.

[0215] Step 503 involves determining the contours (also called streamlines) of the flow function. As will become apparent from the following description, the number of contours is related to the number of tracks in the coil. In one example, the number of contours is pre-defined (e.g., specified by a user). The contours are generated by plotting the trajectory of a constant flow of current at equal intervals.

[0216] FIG. 13A shows an illustrative example of a stream function. In FIG. 13A, the shading indicates the value of the stream function and the arrows indicate the direction and magnitude of the associated current stream density. The stream function shown in FIG. 13A is parameterized by a Fourier-Bessel basis set as shown in FIG. 11, with coefficients associated with the vibrational modes in the second row of the first column (i.e., m=1 and n=1, or u 11 ) is 1, and the coefficients associated with the vibration modes in the third row of the first column (i.e., m=2 and n=1, or u 21 ) is 0.5 and all other coefficients have a value of 0.

[0217] In step 503, the method determines contours (also called streamlines) of the stream function by determining a number of equally spaced values ​​between a maximum of the stream function and a minimum of the stream function, the number of the plurality of equally spaced values ​​being equal to the number of predefined streamlines. In an illustrative example, the number of contours is 4. In one example where the maximum of the stream function is +1 and the minimum of the stream function is -1, and the number of streamlines is equal to 4, the contours are located at -0.75, -0.25, 0.25, and 0.75 in the contour space.

[0218] After determining the locations of the contour lines (ie, their location in the 2D plot that also includes the stream function), the method proceeds to step 504 .

[0219] Step 504 involves positioning the coil tracks. In contrast to previous approaches, the tracks are not initially defined by a physical width (e.g., 250 μm). Instead, in the method of FIG. 12, the tracks are defined by a width "in contour space".

[0220] A stream function can be represented by a function with two input variables (e.g., x and y) that represent the position in the plane of the coil, and the output of the function is the value of the stream function. The value of the stream function is not necessarily a straight line with respect to the input values. For example, the distance in the xy plane between a first series of contour lines (e.g., between contour lines having stream functions of -0.75 and -0.25) and the distance in the xy plane between a second series of contour lines (e.g., between -0.25 and +0.25) may be different even though these contour lines are equally spaced in "contour space."

[0221] Contour space is used herein to describe the (mathematical) space containing the values ​​of the stream function (e.g. values ​​between -1 and +1), while geometric space refers to the positions in the xy plane that correspond to the surface of the coil.

[0222] 12, the contours are equally spaced in "contour space", but the contours are not necessarily equally spaced in "geometric space" because the slope of the stream function is different at different values ​​of the input variables (i.e., at different points in geometric space).

[0223] FIG. 13B shows the effect of setting a constant track width in contour space according to an illustrative example. As mentioned above, the flow function has a value for each point in the plane and is also a function of the x and y positions in the plane. FIG. 13B shows an example contour function as a function of x values, where the y values ​​are fixed. In FIG. 13B, it can be seen how the constant track width changes to a variable distance in geometric space (i.e., in the xy plane of the coil) when the track width is specified in contour space (i.e., as part of the contour function).

[0224] Returning to FIG. 12, positioning the coil track in step 504 includes obtaining a track width value. In one example, the track width value is predefined. In one example, the predetermined width of the track is specified by a user and is specified in "contour space" (i.e., not specified as a width measurement relative to a "geometric space" such as 200 μm, but instead in contour space, e.g., a value of 0.05). The predetermined width of the track is less than the distance, in "contour space", between the contours generated in step 503. For example, in the case where there are four tracks / contours and the flow function takes values ​​between +1 and -1 such that the distance between the contours is 0.5 in "contour space", the maximum width of the track is less than 0.5 (e.g., 0.499 or less).

[0225] The location of the track in "contour space" is then determined. Determining the location of the track in contour space involves determining a point in the plane associated with a stream function value that is less than or equal to half the predetermined track width from the stream function values ​​associated with the contours. In other words, a track in contour space is a point in the plane associated with a stream function value,

number

[0226] For example, if a first contour is associated with a stream function value of 0.5 and the width of the track is specified (in contour space) as 0.2, then determining the location of the track in contour space involves identifying a point in the plane of the coil that is associated with a stream function value that is greater than or equal to 0.4 and less than or equal to 0.6. This process is repeated for each of the contours in the stream function.

[0227] After determining the position of the track in "contour space", a point in geometric space (eg, x and y values ​​in the plane of the coil) that is associated with the position of the track in contour space is then determined.

[0228] By using the areas between the contour lines to define current tracks / coils of varying thickness, it is possible to form the tracks (e.g. copper) and increase the coverage of the coil with conductive material which improves heat dissipation.

[0229] Figure 13C shows an illustrative example of a coil shape. In particular, Figure 13C shows the positions of the tracks generated from the stream functions shown in Figure 13A. As can be seen from Figure 13C, the tracks form a closed loop at this stage in the method. After obtaining the positions of the tracks (in "geometric space"), the method proceeds to step 505.

[0230] In step 505, the track is discretized. After completing step 504, information including the position of the track in geometric space (e.g., in the xy plane corresponding to the surface of the coil) is obtained. The coil track is a closed loop (as shown in FIG. 13C) and may have a variable width in "geometric space" (i.e., in the xy plane of the coil). In step 505, the track is subdivided (i.e., discretized) into multiple sub-tracks that have equal impedance at the frequency of use (which can be obtained by slightly varying the width in contour space to enforce the exact impedance) or have equal width in contour space and Litz wire configuration.

[0231] Advantageously, the subdivision into multiple sub-tracks helps to achieve equal current distribution through the tracks when applying DC, AC, or AC+DC currents, which has the effect of reducing eddy currents and minimizing skin and proximity effects, creating frequency dependent resistance and moving the generated magnetic field away from the target magnetic field (especially in close proximity to coils associated with unilateral MRI machines).

[0232] In a first approach, each track is subdivided into a plurality of concentric and coplanar sub-tracks. After implementing the first approach and discretizing the track, the plurality of concentric and coplanar sub-tracks each have equal impedance. In one example, an iterative method is applied to determine a width for each sub-track in the plurality of sub-tracks that achieves equal impedance for each sub-track in the plurality of sub-tracks. In one example, the number of sub-tracks in the plurality of sub-tracks is predetermined.

[0233] In one example, the first track is divided into multiple portions (i.e., multiple sub-tracks) with gaps between each portion (i.e., between each sub-track). In one example, the width of the gaps is predetermined and specified in "contour space", thus resulting in an amount of available space that can be occupied by the multiple sub-tracks.

[0234] The amount of available space is initially divided equally among the multiple sub-tracks so that each sub-track has an equal width in the "contour space". The length and width of each sub-track are then used to calculate the impedance of each sub-track. If the impedances associated with each sub-track are not similar (i.e., within a predetermined margin of each other), the proportion of available contour width / space allocated to each sub-track is adjusted so that the width of each of the multiple sub-tracks in the "contour space" varies.

[0235] In an illustrative example, a track of width w in "contour space" is divided into two sub-tracks, resulting in a gap of width g between the two sub-tracks. In this case, the available width, u, that can be occupied by each sub-track is wg. Initially, each sub-track is assigned a width that corresponds to an equal share of the available width (i.e., 50% of the available width).

[0236] The impedance of each sub-track is calculated based on the width and length of the sub-track. In response to determining that the impedances of each sub-track are not similar (including equal), the percentage of available space is adjusted (e.g., 45% of the available width is allocated to the width of a first sub-track and 55% of the available width is allocated to the width of a second sub-track). The impedance of each sub-track is then recalculated and the percentage of available width is iteratively applied until similar impedances are obtained for each sub-track.

[0237] During the iterative process, after determining the updated widths of the sub-tracks in "contour space", it is determined whether the width values ​​(specified in "contour space") generate sub-tracks in "geometric space" that are less than a minimum width. In one example, determining whether a sub-track width in "contour space" is associated with a sub-track in "geometric space" that is less than a minimum width includes determining points of a stream function (i.e., in "contour space") that are associated with the sub-track, determining points in an xy-plane that are associated with those points of the stream function, and calculating a geometric width based on the points in the xy-plane.

[0238] In one example, the sub-tracks each have a minimum width equal to the lithographic limits of the equipment used to produce the printed circuit board with the coil, hi one example, the minimum width of each sub-track is 100 μm.

[0239] Similarly, during the first approach it is also determined that the given gap size (specified in "contour space") is greater than the minimum width of the gap in "geometric space".

[0240] In the first approach described above, the width of the sub-tracks is adjusted to obtain similar DC resistance for each sub-track in the plurality of sub-tracks. In another example, the width of each sub-track in the plurality of sub-tracks is adjusted as described above to obtain similar (e.g., identical) impedance in each sub-track at a desired frequency.

[0241] In a further example, the width of each sub-track is determined using the method described in BJ Varghese, T. Smith, A. Azad and Z. Pantic, "Design and optimization of decoupled concentric and coplanar coils for WPT systems," 2017 IEEE Wireless Power Transfer Conference (WPTC), 2017, pp. 1-4, doi: 10.1109 / WPT.2017.7953838, which is incorporated herein by reference, where design constraints (e.g., outer dimensions of the tracks) are provided from the information generated in step 505.

[0242] Fig. 14A shows a plan view of a track according to an illustrative example. Fig. 14B shows a plan view of a track discretized according to the first approach according to an illustrative example. In Fig. 14A, the track 701 (e.g., generated after performing step 504) has a constant track width in "contour space" and a track width in "geometric space" that is determined based on the shape of the stream function as described above.

[0243] 14B, on the other hand, shows a track discretized according to the first approach. After discretizing the track according to the first approach, multiple parallel sub-tracks 702 are obtained, each having a width that achieves equal (or at least similar) impedance in each sub-track.

[0244] In the second approach, each track is discretized into multiple sub-tracks and subsequently transformed into a Litz wire-like structure.

[0245] In one example, each track is discretized into multiple sub-tracks by dividing the width of the "contour space" for the track into multiple portions, each portion associated with a sub-track and with a gap between the portions. In one example, the available space (i.e., the width of the "contour space" of the track minus the contour width of the gap) is divided equally among the multiple sub-tracks. For example, in the case of two sub-tracks, the width of each sub-track is 50% of the available space. The division of the track into multiple sub-tracks, each having equal width in the "contour space", depends on the sub-track having a minimum width, and there is a minimum track size in the "geometric space" as described above in connection with the first approach. In one example, if the minimum width in the "geometric space" is not obtained, the number of sub-tracks into which the track is discretized is reduced.

[0246] As is known in the art, Litz wire is a multi-strand wire or cable that reduces skin and proximity effect losses in the conductor. Typically, Litz wire consists of many thin wire strands twisted or woven together.

[0247] In a Litz-like construction, a first subtrack is provided on a first layer of a printed circuit board and a second subtrack is provided on a second layer of the printed circuit board, the first and second subtracks being arranged to avoid eddy currents and minimize skin and proximity effects which create frequency dependent resistance. A Litz-like construction comprises a number of subtracks which are "logically" equivalent to a group of cables twisted together and flattened.

[0248] Figure 14C illustrates an example of a Litz wire-like structure, according to one example. In particular, Figure 14C illustrates a first sub-track 703 in a first layer and a second sub-track 704 in a second layer.

[0249] The Litz wire-like structure comprises two layers of coplanar subtracks. In one example, a first plurality of subtracks is provided on a first layer of a printed circuit board (PCB) and a second plurality of subtracks is provided on a second layer of the printed circuit board (PCB). The first plurality of subtracks includes the first subtracks and the second plurality of subtracks includes the second subtracks.

[0250] The first sub-track is "staircase" shaped and includes a plurality of coaxial sections, each radially offset relative to the axis of the coil, and a plurality of transitions between the sections, thereby creating a staircase shaped conductive sub-track. Each sub-track in the first plurality of sub-tracks has a similar shape and is disposed parallel to the first sub-track.

[0251] FIG 14D illustrates a first sub-track and a third sub-track in a Litz wire configuration according to one example. In particular, FIG 14D illustrates a first sub-track 703 and a third sub-track 705 in a first layer of a printed circuit board arranged.

[0252] Similarly, the second sub-track (in the second layer) has a "staircase" shape and includes a number of concentric sections, each radially offset with respect to the axis of the coil, and a number of transitions between the sections, but with an opposite slope to the sub-tracks in the first plurality of sub-tracks.

[0253] In one example, the portion of the first sub-track in the first layer has an increasing radius from the center of the coil in a clockwise direction (in other words, the first sub-track takes the form of a "step" that ascends in a clockwise direction), while the portion of the second sub-track in the second layer has a decreasing radius from the center of the coil in a clockwise direction (in other words, the second sub-track takes the form of a step that descends in a clockwise direction).

[0254] The transitions between the constant radius portions of the subtracks in the first layer are positioned to coordinate with the transitions between the constant radius portions of the subtracks in the second layer, thereby creating an overlapping arrangement that avoids eddy currents and minimizes skin and proximity effects.

[0255] In a Litz wire-like structure, a first layer (also called the top layer) is electrically coupled to a second layer (also called the bottom layer). In one example, the first layer is electrically coupled to the second layer using "vias" (i.e., electrical connections between copper layers of a PCB).

[0256] Figure 14E shows a plan view of the top and bottom layers of a printed circuit board (PCB) coil, according to one example. In the illustrative example of Figure 14E, the track is discretized into a Litz wire-like structure, with multiple sub-tracks provided in the top layer 751 and bottom layer 752 that cross each other. At least one electrical connection in the form of a "via" 753 is also provided between the sub-tracks in the top layer 751 and the sub-tracks in the bottom layer 752.

[0257] In the example shown in Figure 14E, there is a first via coupled to the start of a first subtrack in the top layer. There is a second via coupled to the end of the first subtrack, coupling the first subtrack to the start of a second subtrack in the bottom layer. A third via is also provided, coupled to the end of the second subtrack in the bottom layer, coupling the second subtrack in the second layer to the third subtrack in the top layer.

[0258] Using the second approach (i.e., discretizing the tracks using a Litz wire configuration) is particularly advantageous when developing coils for AC use or combined AC and DC. Using the second approach allows for uniform current distribution not only at DC, but also over the entire operating frequency, provided that the skin depth is not much smaller than the sub-track width. This is a weak constraint for low-field MRI devices operating with current signal frequencies of several hundred kilohertz.

[0259] The first and second track discretization approaches described above allow for near-perfect packing of the copper surface and elimination of skin depth effects across each track, resulting in much higher power efficiency per PCB layer / thickness than previous approaches. At the same time, the first and second approaches ultimately allow for a controllable design that ensures a desired DC to high frequency current distribution in space according to a desired current stream density distribution that ensures that the generated DC and AC magnetic fields match the desired fields. Advantageously, the second approach also allows for a much lower inductance, since it distributes the current equally over many subtracks with Litz wire. This allows for the effective connection of many parallel contours (fewer contours of larger width in the contour space).

[0260] In one example, the approach used to discretize the tracks in step 505 depends on the current used to energize the coil in use: if the coil current is due to DC current, the first approach (i.e. single layer multi-track configuration) is used, whereas if AC and DC are applied to the current in use, the second approach (i.e. multi-layer Litz wire like track configuration) is used.

[0261] Returning to Figure 12, after discretizing the tracks in step 505, the method proceeds to step 506. As mentioned above, the tracks generated in step 504 and the sub-tracks generated from them in step 505 are closed loops (i.e., each track is not connected to another track). In step 506, the coil tracks are connected to each other to form a continuous track.

[0262] In one example, connecting the coil tracks includes forming a break (i.e., interrupting the conductive path) in the first track (in other words, all the sub-tracks into which the first track is discretized), thereby creating a first end of the first track and a second end of the first track. The conductive paths of the adjacent tracks are split in a similar manner, creating a third end and a fourth end of the adjacent tracks. The second end of the first track is coupled (i.e., electrically coupled) to the third end of the adjacent track, thereby creating a continuous conductive path from the first end of the first track to the fourth end of the adjacent track. If there is a track adjacent to the first track in a different direction from the "adjacent track", the first end is connected to the end of that track. Similarly, if there is a track adjacent to the "adjacent track" in a different direction from the first track, the fourth end is connected to the end of that track. Each of the tracks in the coil is connected in this way, and a continuous conductive path is created between all of the tracks.

[0263] Figure 15A shows top and bottom layers of coils with connected tracks according to an illustrative example, Figure 15B shows a close-up view of the connections between adjacent coils according to an illustrative example, and Figure 15C shows a perspective view of a coil fabricated with Litz wire sub-tracks according to an illustrative example.

[0264] Returning to Figure 12, after connecting the tracks in step 506, the method proceeds to step 507. In step 507, information identifying the coil is generated. In one example, the information identifying the coil includes position information in the plane of the coil where the conductive elements (e.g. copper tracks) will be located. In this way, the coil shape is defined. As will be explained in more detail below, this information can then be used to control a machine that manufactures / creates the coil.

[0265] 16 illustrates a method of manufacturing a coil, according to an example. Manufacturing includes creating and / or assembling the coil. In step 901, information identifying the coil is obtained. In one example, the information identifying the coil includes position information identifying the location of conductive elements (e.g., copper tracks) in the plane of a printed circuit board (PCB). The method proceeds to step 902.

[0266] In step 902, a coil is manufactured according to the information obtained in step 901. In one example, manufacturing the coil includes manufacturing a layer of a PCB that includes the coil. There are various techniques for manufacturing a layer of conductive material on a surface of a PCB. As a result, a detailed description of the manufacturing process is omitted for the sake of brevity.

[0267] In one example, the method of Figure 10A (iteratively determining the coil flow function to obtain the target field) or the method of Figure 10B (determining the flow function for the target field by direct calculation), the method of Figure 12 (determining the coil shape based on the current distribution), and the method of Figure 16 (manufacturing the coil based on the determined coil shape) are performed sequentially as part of a single method, resulting in a layer of a PCB having a coil that obtains a target magnetic field in a volume of interest.

[0268] In one example, the methods of Figures 10A, 10B, and 12 are computer-implemented (ie, implemented by a computer processor).

[0269] FIG. 17 illustrates an implementation of an apparatus according to an example. In particular, FIG. 17 illustrates an apparatus 1000 comprising an input / output interface 1010. The input / output interface 1010 is configured to receive input data (e.g., from a user) and transmit output data (e.g., to a user or to a PCB manufacturing machine). The apparatus 1000 further comprises a processor 1020 (e.g., a Central Processing Unit) coupled to a non-volatile memory 1030. The processor 1020 is configured to execute instructions stored in the non-volatile memory 1030. Execution of these instructions causes the apparatus 1000 to perform some of the method steps described herein, including, but not limited to, the methods of FIG. 10A, FIG. 10B and / or FIG. 12.

[0270] The above examples are described in the context of a single coil design. However, as noted above, in some examples, the coil component 370 comprises multiple coils that are offset vertically within the coil component 370, e.g., by being implemented on different layers of a multi-layer printed circuit board. It is emphasized that while the techniques have been described in the context of a single coil design, the techniques may also be used to design multiple coils. In this case, the method (e.g., the method of FIG. 10A or FIG. 10B, FIG. 12, and FIG. 16) is performed multiple times, each time for a different coil in the coil component 370.

[0271] Further, in the above, the methods are applied to coils that control spin manipulation conditions in a sample during an NMR experiment, however, it is emphasized that the techniques described herein may also be used to design other coils in NMR system 800 or NMR system 100 (e.g., Rx coils and / or prepolarization coils).

[0272] Additionally, in other examples, the techniques described herein are used to design coils for use in use cases other than NMR / MRI systems. For example, the methods described herein can be used to design any coil required to generate any magnetic field precisely. In one example, the methods described herein are used to obtain coils for quantum computing use cases where ion traps such as Penning traps are used to control qubits. In another example, the methods described herein are used to obtain coils for use in fundamental research where ion traps store exotic particles such as antimatter or antiprotons in magnetic fields. In another example, the methods described herein are used to obtain coils for use in fusion power reactors where precise magnetic fields are used to confine high energy plasmas. In another example, the methods described herein are used to design coils for use in chemical analysis, particularly mass spectrometry, where charged ions are accelerated through a magnetic field and the resulting deflection from their trajectories provides insight into the mass per charge ratio. In this use case, precise B fields are important to obtain results.

[0273] As described above, the method described herein fully automates the design of a printed circuit board (PCB) that, when connected to an AC and / or DC current source, generates a desired magnetic field that closely resembles the target magnetic field in the target volume, even in the vicinity of soft magnetic materials, independent of the frequency of the driving current, and without coupling to specific (Rx) coils close to the PCB, thereby enabling a single-sided NMR scanner.

[0274] In a further example, a method is provided for designing and producing a coil component comprising at least three coils such that the coils are capable of performing magic angle field rotation (MAFS).

[0275] The use of magic angle field spinning (MAFS) in magnetic resonance imaging (MRI) allows for a controllable increase in the effective T2 resulting from the anisotropic contribution to the binding interaction of the species under observation, which can be significantly increased, assuming that the rate of rotation frequency exceeds the decay rate of the highest binding to the spin species.

[0276] In conventional magic angle spinning, the sample is rotated at the magic angle, MA, with respect to the direction of the magnetic field, ω MA The human rotates at a magic angle frequency of 54.74 degrees. As is known in the art, the magic angle is approximately equal to 54.74 degrees. Of course, rotating a human at high frequencies is impractical.

[0277] At the magic angle to the B0 field, ω MA The geometrical counterpart to rotating the sample at the magic frequency of is to create a conically rotating B0 field, which can be expressed as:

number

[0278] In the above example, the direction vector is aligned with the x, y, and z axes, but it is emphasized that other directions may be used. Furthermore, in other examples, the direction vector may vary along the space within the volume of interest (VOI). For example, the (time-invariant) DC component in the above equation may point in the X direction at a point in the volume of interest (VOI). In one example, the magic angle frequency, ω MA is determined based on the frequency at which the field needs to oscillate to obtain the desired behavior.

[0279] FIG. 18 illustrates a method for designing multiple coils capable of performing magic angle magnetic field rotation (MAFS) according to an example. In this example, the multiple coils include at least three coils for generating a magnetic field in a volume of interest (VOI). In one example, the coil component comprises multiple coils. In the example of FIG. 18, at least three coils are provided for generating a B0 field.

[0280] Each coil generates an orthogonal basis for the magnetic field. In one example, the magnetic field is modeled by an orthogonal basis set that contains multiple orthogonal bases. An orthogonal basis set is a set of perpendicular vectors that together define every direction (e.g., {1,0,0}, {0,1,0}, {0,0,1}). As the MAFS field changes, the direction of Z changes along with the original B0 / b0 direction. Therefore, the other two (AC) coils also change direction since they are all perpendicular.

[0281] The method begins at step 1101. Step 1101 includes generating a first plurality of coefficients (e.g., c 00 , c 10 , c 20 , c 30 , c 40 , c i0 10A or 10B, as described above. The method proceeds to step 1102.

[0282] In step 1102, a second plurality of coefficients (e.g., c 01 , c 11 , c 21 , c 31 , c 41 , c i1 etc.) is obtained and a third plurality of coefficients (e.g., c 02 , c 12 , c 22 , c32 , c 42 , c i2 In one example, a second plurality of coefficients (e.g., c 01 , c 11 , c 21 , c 31 , c 41 , c i1 ) and a third plurality of coefficients (e.g., c 02 , c 12 , c 22 , c 32 , c 42 , c i2 , etc.) may be obtained from a file stored in computer memory or based on a random initialization.

[0283] The field in the volume of interest generated by the second coil can be expressed as:

number

[0284] Similarly, the field in the volume of interest generated by the third coil can be expressed as:

number

[0285] In use, application of sinusoidal currents to the second and third coils results in a field in the volume of interest equal to:

number

[0286] The field generated in the volume of interest is called the magic angle rotation field, B 0_MAFS To be identical to 01 and b 02 Vector Field The following conditions must be met:

number

[0287] The above conditions are that the fields are all perpendicular at each point in space, and 01 and b 02 It specifies that the field has a precise norm (which depends on the size of b0 at the point in space).

[0288] JPEG2025512122000020.jpg41166

[0289] The method proceeds to step 1104 where a metric at each point in space is determined. In one example, a composite metric at each point in space is calculated based on a sum of five metrics, the five metrics being given below:

number

[0290] In one example, a composite metric is calculated based on a weighted sum of metrics 1 through 5. The composite metric is calculated for each point in space in the volume of interest.

[0291] The zero value for the composite metric at each point in space indicates that the three coils are in a magic angle rotation field B since all the equations are satisfied. 0_MAFS On the other hand, if the value of the composite metric at a point in space within the volume of interest is greater than zero, then the magic angle rotation field B 0_MAFS The equation for obtaining {overscore (x)} is not satisfied and the coefficients representing the stream function are not associated with a coil capable of generating a magic angle rotation field, in use. After determining the metrics for each point in space, the method proceeds to step 1105.

[0292] In step 1105, a metric is obtained over the volume of interest. In one example, step 1105 includes summing the composite metrics at each spatial point within the volume of interest to obtain a single value called a total metric. In a further example, the total metric also includes a normalization term that is based on the power efficiency of the coil and the coil inductance. The method proceeds to step 1106.

[0293] The sum metric is compared to a threshold in step 1106. If the sum metric is less than the threshold (indicating that the field associated with the first, second, and third pluralities of coefficients approximates a magic angle rotation field), the method proceeds to step 1107.

[0294] In step 1107, the iterative design method ends and outputs information identifying a stream function associated with each of the at least three coils (i.e., the first, second, and third plurality of coefficients). In one example, the first, second, and third plurality of coefficients are used as part of a production process to produce the coils. The first, second, and third coils are generated and produced based on the first, second, and third plurality of coefficients according to the methods of Figures 12 and 16.

[0295] On the other hand, if the sum metric is greater than the threshold (indicating that the field associated with the first, second, and third pluralities of coefficients does not approximate the magic angle rotation field), the method proceeds to step 1108 .

[0296] In step 1108, a second set of coefficients (e.g., c 01 , c 11 , c 21 , c 31 , c 41 , c i1 ) and a third set of coefficients (e.g., c 02 , c 12 , c 22 , c 32 , c 42 , c i2The values ​​of the coefficients in the set of coefficients are then adjusted. The method then proceeds to step 1103 where the field associated with each set of coefficients is determined.

[0297] In this manner, the method of FIG. 18 iteratively changes the values ​​of the second and third series of coefficients until the sum metric is minimized / converged (i.e., as close to zero as possible), thereby obtaining the magic angle rotation field.

[0298] Advantageously, by using the above multiple metrics, the gradient of the function being minimized (i.e., the gradient of the composite metric) has a well-defined gradient and global optimum is guaranteed (i.e., a convex problem).

[0299] Although the above examples have been described in the context of a B field, it will be appreciated that the techniques are not limited to B fields. In other examples, the above techniques are applied to other fields in an NMR system.

[0300] In one example, the method of Figure 18 is computer-implemented, for example, by using the apparatus described in connection with Figure 17.

[0301] In a further example, a first plurality of coefficients (for a first coil) is obtained using the method of either Figure 10A or Figure 10B, and a second plurality of coefficients (for a second coil) and a third plurality of coefficients (for a third coil) are obtained using the method of Figure 11. Coil shapes for the first coil, second coil, and third coil are determined by repeating the method of Figure 11 for each of the first plurality of coefficients, the second plurality of coefficients, and the third plurality of coefficients. The coil components are then produced in a multi-layer PCB according to the method of Figure 16.

[0302] 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. 1. A method for identifying a coil shape, comprising: obtaining first information identifying a stream function associated with the coil; determining the stream function based on the first information; and determining a position of a first contour of the stream function; determining a position of a first coil track based on the position of the first contour line and a track width, the track width of the first coil being determined by a value of the stream function; discretizing the first coil track into a plurality of sub-tracks; generating second information identifying the shape of the coil; The method includes:

2. The method of claim 1 , further comprising producing the coil based on the second information specifying the shape of the coil.

3. The method of claim 1 or 2, wherein determining the location of the first contour line comprises identifying a point in the plane of the coil that is associated with a first value of the stream function.

4. determining a difference between a minimum of the stream function and a maximum of the stream function; determining a contour distance based on the difference divided by a plurality of contour lines; determining the first value of the stream function based on the contour distance and a second value of the stream function associated with a second contour; The method of claim 3 further comprising determining the first value of the stream function by:

5. determining a location of a second contour line based on the stream function; and determining a position of a second coil track based on the position of the second contour line and the track width, the track width being specified by values ​​of the stream function such that the first coil track and the second coil track occupy the same range of stream function values. The method of any one of claims 1 to 4, further comprising:

6. Determining the position of the first coil track includes: 1) greater than or equal to the first value of the flow function minus half the track width, and 2) equal to or less than the first value of the flow function plus half the track width 5. A method according to claim 3 or 4, comprising identifying a position in the plane of the coil which is associated with a value of the stream function where

7. The plurality of sub-tracks include a first sub-track having a first width; a second sub-track having a second width; Equipped with The method of any one of claims 1 to 6, wherein the first sub-track and the second sub-track are coplanar.

8. discretizing the first coil into the plurality of sub-tracks includes:

8. The method of claim 7, comprising determining the first width of the first sub-track such that the first sub-track has an impedance equal to an impedance of the second sub-track.

9. 9. The method of claim 7 or 8, wherein discretizing the first coil track into a plurality of sub-tracks comprises producing a Litz wire in a first layer and a second layer of the coil.

10. The information identifying the stream function associated with the coil may include:

10. The method of claim 1, comprising a plurality of coefficients, each coefficient of the plurality of coefficients being associated with a basis function in a plurality of basis functions for modeling the stream function associated with the coil.

11. The method of claim 10 , wherein the plurality of basis functions models a Fourier-Bessel basis set.

12. Obtaining the information identifying the stream function associated with the coil comprises: obtaining a first set of coefficient values ​​including a plurality of coefficients, each coefficient of the first set of coefficients being associated with a basis function for modeling the stream function associated with the coil; determining a value of a first coefficient based on the first set of coefficients; determining a magnetic field in a volume of interest generated by a coil having a stream function represented by the first set of coefficient values ​​and the first coefficients; comparing said magnetic field to a target magnetic field to generate a difference metric; determining a loss metric based on the difference metric; In response to determining that the difference metric satisfies a stopping condition, generating the information identifying the stream function based on the first set of coefficient values ​​and the first coefficients; The method according to any one of claims 1 to 11, comprising:

13. Obtaining the information identifying the stream function includes: In response to determining that the difference metric does not satisfy the stopping condition, adjusting the first set of coefficient values; recalculating the first coefficient; determining a second magnetic field in the volume of interest; comparing the second magnetic field to the target magnetic field; The method of claim 12 further comprising:

14. The model of the system includes the coil and a second coil; 14. The method of claim 12 or 13, wherein determining a value of the first coefficient comprises determining a value of the first coefficient such that a voltage induced by the coil in a second coil is zero.

15. The first coefficient is [0010] where c is calculated according to ０ is the first coefficient, and v ｏ is the voltage induced in the second coil due to the current distribution in the plane of the coil corresponding to the 0th basis field, and c ｉ is the coefficient associated with the i-th basis field, and v ｉ 15. The method of claim 14, wherein i i =i i th basis field, ij i =i i th , where i i i =i i th basis field, and j i =i i th , where j i i =i i th basis field.

16. 16. The method of claim 12, wherein determining the magnetic field in the volume of interest comprises: simulating the magnetic field in the volume of interest for each basis function using a Finite Element Model (FEM) solver; and combining the simulated magnetic fields according to the first set of coefficients.

17. The method of claim 16 , wherein the FEM solver simulates the magnetic field in the volume of interest in the presence of linear magnetic materials and / or eddy currents in the magnetic field produced by the current distribution.

18. 1. A method of generating a multiple coil geometry for magic angle magnetic field rotation, the multiple coils comprising a first coil, a second coil, and a third coil, the method comprising: - determining the shape of the first coil according to any one of claims 1 to 17 such that the first coil generates a first magnetic field; obtaining a second plurality of coefficients and a third plurality of coefficients, the second plurality of coefficients models a second stream function associated with the second coil; the third plurality of coefficients modeling a third stream function associated with the third coil; and determining a second magnetic field associated with the second plurality of coefficients and a third magnetic field associated with the third plurality of coefficients; determining a metric based on the first magnetic field, the second magnetic field, and the third magnetic field; In response to determining that the metric is less than a predetermined threshold, adjusting a value of a coefficient of the second plurality of coefficients and / or a second value of a second coefficient of the third plurality of coefficients; The method includes:

19. Determining a metric based on the first magnetic field, the second magnetic field, and the third magnetic field includes: calculating a second metric based on the first magnetic field, the second magnetic field, and the third magnetic field at each point in space within a volume of interest; summing the second metrics associated with each point in space to obtain the metric; 20. The method of claim 18, comprising:

20.

21. A non-transitory computer readable medium comprising computer program instructions suitable for execution by a processor, the instructions being configured, when executed by the processor, to perform a method according to any one of claims 1 to 20.

22. A coil obtainable by the method according to claim 2.

23. 23. A coil component comprising the coil of claim 22 and a second coil, wherein the coil and the second coil are both coplanar and coaxial, and the coil component is integrally formed.

24. 24. The coil component of claim 23, wherein the coil component comprises a printed circuit board, the printed circuit board comprising a first layer and a second layer, the first layer comprising the coil, and the second layer comprising the second coil.

25. A coil component according to claim 23 or 24; Receiving coil and Equipped with The nuclear magnetic resonance system, wherein the receive coil is coaxial with the coil component.

26. 1. A coil component comprising a first coil and a second coil, wherein the first coil and the second coil are both coplanar and coaxial, and the coil component is integrally formed.

27. The coil component of claim 26 , wherein the coil component is cylindrical.

28. 28. The coil component of claim 26 or 27, wherein the coil component comprises a printed circuit board, the printed circuit board comprising a first layer and a second layer, the first layer comprising the first coil, and the second layer comprising the second coil.

29. A coil component according to any one of claims 26 to 28; Receiving coil and Equipped with The NMR system, wherein the receive coil is coaxial with the coil component.

30. The first coil is B ０ 30. The NMR system of claim 29 for generating a field.

31. 31. The NMR system of claim 29 or 30, wherein the coil components comprise all coils for operating the NMR system other than the receive coil.

32. All coils for operating the NMR system are B ０ , B １，ｔｘ , G ｘ , G ｙ , and G ｚ 32. The NMR system of claim 31 comprising a coil.

33. 33. The NMR system of claim 29, wherein the first coil and the second coil in the coil component are separated from the receive coil.

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