Conductive loop with inductive path for magnetic resonance imaging (MRI) receiving coil

Abstract

A magnetic resonance (MR) receive coil (18) includes a conductive loop (20) that resonates at a Larmor frequency of a B0 magnetic field on a design basis of at least 3 Tesla; and a conductive trace (22) configured to detune the conductive loop from the Larmor frequency in response to a DC current flowing through the conductive trace. The conductive trace also includes a receive B1 field uniformity enhancing inductor (L3) positioned at an intermediate point along the conductive trace.

Description

Technical Field

[0001] The following generally relates to magnetic resonance (MR) imaging techniques, MR coil techniques, high magnetic field MR signal acquisition techniques, MR detuning techniques, and related techniques. Background Art

[0002] Magnetic resonance (MR) imaging requires placing an object (e.g., medical patient, veterinary subject, archaeological mummy, etc.) in a static magnetic field (commonly referred to as B 0 The magnetic field is placed in a static B field and nuclear magnetic resonance is excited in the object and the excited magnetic resonance is then detected. For imaging, a magnetic field gradient is superimposed on the static B field during the excitation, during the time interval between the MR excitation and the MR readout, and / or during the MR readout. 0 The magnetic field spatially encodes the excited MR with respect to position, phase and / or frequency. In a typical design, an MR imaging device (sometimes called an MRI scanner) includes a housing having a central bore within which the MR examination region is located. Static B 0 The magnetic field is generated by solenoid magnet windings wrapped around the central bore and housed within the MRI scanner housing. These solenoid magnet windings are typically superconducting windings in modern MRI scanners, and the housing includes a reservoir of liquid helium (LHe) that cools the superconducting windings. Magnetic field gradient coils are also disposed in the housing around the central bore.

[0003] To provide MR excitation in the case of human subjects, a body coil is typically used, which is typically a cylindrical birdcage coil, a TEM coil, or some variation thereof, mounted concentrically around the bore. Alternatively, a local coil positioned near the body anatomy to be imaged is used for excitation. MR readout is typically performed using a local MR receive coil positioned near the anatomy to be imaged. The local MR receive coil and the local MR excitation coil readout (if used) may be the same coil or different coils. For various reasons, the MR receive coil (and MR excitation coil, if used) may include an MR coil that includes one or more coil elements, wherein each coil element is typically configured as a loop coil, but other coil element designs are also known.

[0004] In current coil designs, the B based on designs of at least 3 Tesla 0 A single loop is used in the magnetic field. However, at 3T and above B 0 Such a single loop used in a magnetic field will suffer from receiving |B 1 |Field distortion, resulting in a decrease in the MRI signal-to-noise ratio (SNR).

[0005] Certain improvements are disclosed below to overcome these and other problems. Summary of the invention

[0006] In some embodiments disclosed herein, a magnetic resonance (MR) receiving coil includes: a conductive loop having a B of at least 3 Tesla. 0 The conductive trace further comprises a receiving portion located at a midpoint along the conductive trace. 1 |Field uniformity enhancing inductors.

[0007] In some embodiments disclosed herein, a method for optimizing the reception of an MR receiving coil is provided. 1 | The method of field uniformity includes placing the MR receiving coil at a design base of at least 3 Tesla B 0 In a magnetic field, wherein the MR receiving coil is coupled to a dielectric body. The MR receiving coil comprises: a conductive loop, which is designed based on B 0 resonant at the Larmor frequency of the magnetic field; and a conductive trace configured to detune the conductive loop from the Larmor frequency in response to a DC current flowing through the conductive trace, the conductive trace also comprising a receiving | B 1 |Field uniformity enhancement inductor, the receiving|B 1 | Field uniformity enhancement inductors are positioned in the conductive traces with the design based on the B 0 The method further comprises adjusting the MR receiving coils including at least the receiving coils. 1 | One or more components of the field uniformity enhancing inductor to optimize the B being placed on the design basis 0 The receiving coil of the MR receiving coil in the magnetic field and coupled to the dielectric body|B 1 |Field uniformity.

[0008] One advantage is that it provides a method for providing a receiver having a reduced |B 1 | Field distortion in MRI receive coils.

[0009] Another advantage resides in providing an MRI receive coil having conductive traces positioned on one side of a single loop coil.

[0010] Another advantage resides in improved SNR of the MR scanner.

[0011] Another advantage is that a device having a receiving device including receiving |B 1 | Field uniformity enhancing inductor conductive traces of an MRI receive coil, the receive | B 1 | Field uniformity enhancing inductors are positioned in the conductive traces and designed based on the B 0 The magnetic field is oriented parallel to the part.

[0012] A given embodiment may provide none, one, two, more, or all of the foregoing advantages, and / or may provide other advantages that will become apparent to one of ordinary skill in the art upon reading and understanding the present disclosure. BRIEF DESCRIPTION OF THE DRAWINGS

[0013] The present disclosure may take the form of various components and arrangements of components, and various steps and arrangements of steps.The drawings are only for the purpose of illustrating the preferred embodiments and are not to be construed as limiting the present disclosure.

[0014] Figure 1 A magnetic resonance (MR) imaging device including an MR coil element according to the present disclosure is schematically shown.

[0015] Figure 2 Shown for Figure 1 The MR coil of the device.

[0016] Figure 3-5 Shown for Figure 2 Experimental data of MR coils.

[0017] Figure 6 Schematically shows the use of Figure 1 An MRI imaging method of the device. DETAILED DESCRIPTION

[0018] In high B 0 Under magnetic field (e.g., 3T and above), |B 1 |Field distortion is a problem due to the short radio frequency (RF) wavelength in the patient. For example, at 3T, the magnetic resonance (MR) resonant frequency is at 128 MHz, which corresponds to an RF wavelength of 2.34 meters in air. Due to dielectric effects, the wavelength is shortened in the patient to a length comparable to the cross-sectional diameter of the patient. The shortened wavelength causes |B 1 |fields, which will cause |B 1 |Field distortion.

[0019] The following discloses the use of conductive copper traces, which are sometimes included in MR receive coils, for performing detuning of the MR receive coil loop during the transmit phase and also for suppressing receive |B 1 |Field distortion to enhance reception|B 1 | Field uniformity. In some embodiments, additional inductors are included in the copper traces. In some embodiments, additional inductors are positioned along the z direction (which is B 0The MR receiving coil may also include end inductors L1 and L2 for DC pass-through to perform detuning of the MR receiving coil loop during the transmit phase.

[0020] In addition to providing for suppressing the reception of |B 1 | In addition to an additional (e.g., symmetrically centered) inductor for field distortion, a calibration coil loop is disclosed below to enhance reception |B 1 |Field uniformity method. In one method, the inductor of the MR receive coil (in addition to the inductor added to suppress the receive |B 1 |Field distortion inductor) has a fixed inductance value (e.g., 1.8 microhenry in one non-limiting illustrative example), and the calibration method includes adjusting the inductance of the added inductor to optimize the reception |B 1 | Field uniformity. A phantom representing the patient's dielectric block can be used in the calibration method, or an electromagnetic simulation of the coil and dielectric block can be used instead. In some embodiments, the disclosed calibration method requires adjustment of a center inductor that is symmetrically positioned on the coil loop along the z direction. Optimization is an empirical process.

[0021] Typically, a calibration process is performed in the factory where the MR receive coils are manufactured to optimize the receive |B 1 | Field uniformity calibration of the coil loop. The calibrated coil loop can then be installed as one coil loop, and the MR coil includes an array of such coil loops. When the MR coil is used for patient imaging, the calibrated MR coil loop should be positioned about B 0 The field and patient are in the same spatial orientation as used in the calibration. In practice, the calibrated coil loop is typically part of a coil array that is sized and shaped to couple to a specific part of the human anatomy (e.g., a head coil array or a torso coil array), thereby naturally achieving proper positioning of the coil loop in clinical use.

[0022] refer to Figure 1 , an illustrative magnetic resonance (MR) imaging system or apparatus 10 includes a magnetic resonance (MR) imaging scanner (e.g., an MR scanner that generates a magnetic field of at least 3 Tesla (3T)). Figure 1 As shown, the MR device 10 includes a housing or a frame 2, which contains Figure 1 The various components shown, for example, as non-limiting illustrative examples, Figure 1 Schematic indication of the generation of static (B 0 ) Superconducting magnet 4, used to superimpose the magnetic field gradient on B 0Magnetic field gradient coils 6 on the magnetic field, whole body radio frequency (RF) coils 8 for applying RF pulses to excite and / or spatially encode magnetic resonance in an imaging patient disposed in an MR bore 12 or other MR examination region, etc. The magnet 4 and gradient coils 6 are arranged concentrically around the bore 12. A robotic patient examination table 14 or other patient support enables a medical patient, a patient undergoing medical screening, or other imaging patient to be loaded into the MR bore 12 for imaging. The magnetic resonance excited in the imaging subject is read out by an MR receiving coil 18. Figure 1 The illustrative MR coil 16 is shown as an array of coil loops 18 (i.e., MR receive coils 18). It will be appreciated that the coil array 16 may generally include any number of coil elements, for example, 16 coil elements, 20 coil elements, 32 coil elements, etc. In some examples, the MR coil 16 is configured to be disposed in an examination region (i.e., the MR bore 12), such as Figure 1 shown.

[0023] Typically, the MR coil 16 will be at a position relative to the B generated by the magnet 4. 0 The illustrative coil 16 is a torso coil placed on the patient's torso in a fixed orientation. Therefore, each MR receive coil 18 making up the coil array 16 is also typically placed in a fixed orientation relative to B. 0 The fixed spatial orientation of the magnetic field is positioned on the patient. It should also be noted that the MR coil array 16 (and therefore its component, the MR receive coil loop 18) is placed on or in close proximity to the patient. This advantageously provides a strong electromagnetic coupling between the patient and the MR receive coil loop 18. However, at high magnetic fields (e.g., in some embodiments, |B 0 |=3 Tesla or higher), |B may occur due to the dielectric effect of the patient 1 |Field distortion. As previously mentioned, this is a result of the shortened wavelength in a patient at high fields becoming comparable to the cross-sectional diameter of the patient.

[0024] Reference now Figure 2 , showing Figure 1 An illustrative MR receiving coil 18 of the MR coil (array) 16 is shown. The MR receiving coil 18 includes, for example, a B based on a design of at least 3T. 0 The conductive loop 20 is a single loop of copper, copper alloy, or another electrically conductive material, such as formed as a copper layer deposited on a circuit board, plastic sheet, plastic former, or other electrically insulating substrate; or alternatively formed as a free-standing metal loop. The conductive loop 20 includes at least one capacitor (four of which are labeled C1-C4, which are Figure 2 ), and in the illustrative example also includes an inductor (in Figure 24. The illustrative conductive loop 20 is a circuit diagram of an MR receive coil 18 that is electrically conductive and resonant. The illustrative conductive loop 20 is an inductor (labeled L4 in FIG. 1 ) that is placed in parallel across capacitor C4 to form an LC detuning circuit for detuning the MR receive coil 18 during the transmit phase of the whole body coil 8. The illustrative conductive loop 20 is rectangular, but other geometric shapes (e.g., circular, etc.) are also contemplated. Note that the copper layer or the like that forms the body of the conductive loop 20 may have gaps at the locations of the capacitors C1-C4, such that the connection of the capacitors across these gaps collectively forms the conductive loop 20. It will also be appreciated that the conductive loop 20 is conductive (and resonant) at the Larmor frequency; however, at DC, the conductive loop 20 is non-conductive due to the presence of the capacitors C1-C4 (which act as blocking capacitors to block the flow of DC current around the conductive loop 20).

[0025] The MR coil 18 also includes a conductive trace 22 that is separate from the conductive loop 20 and located close enough to it to enable inductive coupling between the conductive loop 20 and the conductive trace 22 at (and near) the Larmor frequency. The conductive trace 22 is configured to detune the conductive loop 20 from the Larmor frequency in response to a DC current flowing through the conductive trace 22. Figure 2 As shown, conductive trace 22 extends parallel to conductive loop 20 along a portion of conductive loop 20. In the illustrative example, conductive trace 22 extends parallel to conductive loop 20 along half of conductive loop 20, although deviations from fractions of half are also contemplated.

[0026] Conductive trace 22 includes a plurality of inductors (at Figure 2 2. The plurality of (illustratively, two) inductors (e.g., L1 and L2) can be coupled inductors that connect opposite ends of the conductive trace 22 to the conductive loop 20. At least one of the inductors (illustratively, inductor L3) includes a receiving |B 1 |Field uniformity enhancement inductor. In the following cases, tuned reception |B 1 |Field uniformity enhancement inductor L3 to optimize reception of MR receive coil 18 |B 1 |Field uniformity: The MR receiving coil 18 is placed on the B 0 In the magnetic field, the branch of the conductive trace 20 includes a B 0 Reception of magnetic field parallel orientation|B 1 |Field uniformity enhancement inductor. Figure 2 In the illustrative example of FIG. 1 , the MR receive coil loops 18 are positioned throughout the coil array 16 (see Figure 1 ), where, in this illustrative example, B 0 The magnetic field is oriented parallel to the branch of conductive trace 20 that includes inductor L3. Figure 2The illustrative XZ coordinate system shown corresponds to B oriented in the Z direction. 0 Magnetic field, and similarly, the MR receiving coil loop 18 includes a receiving coil | B 1 | The branch of the field uniformity enhancement inductor L3 also follows the B 0 The magnetic field is oriented parallel to the Z direction.

[0027] Receive | B 1 |Field uniformity enhancement inductor L3 is positioned at a midpoint along conductive trace 22. In one example, |B 1 |The field uniformity enhancement inductor L3 is positioned at the midpoint of the conductive loop 20. In another example, |B 1 | Field uniformity enhancement inductor L3 and one of the capacitors of the conductive loop 20 (ie, as Figure 2 The capacitor C4) is shown located in one place.

[0028] Figure 3-5 Example experimental data for the MR coil 18 is shown. Figure 3 A simulated arrangement of an MR coil 18 comprising a single conductive loop 20 with a conductive trace 22 on the left half of the loop 20 is shown, the simulated arrangement being arranged on a simulated dielectric or dielectric block 24 simulating the effects of a patient. The simulated coil 18 is oriented so that the branch containing capacitor C4 is immediately adjacent to B 0 Field parallel orientation (using an illustrative XYZ coordinate system, along the Z direction) reception | B 1 |Field uniformity enhancement inductor L3. Figure 4 and Figure 5 The dimensions of the loop 20 simulated in the example are as follows: 0 The conductive width of the loop 20 is 6 mm, and the conductive width of the trace 22 is 2 mm. In this embodiment, the conductive trace 22 is placed inside the loop 20, in one location with the branch containing the capacitor C4, wherein the conductive trace 22 is positioned 2 mm away from the loop 20. In another embodiment, the trace 22 can be placed below or above the single loop conductor 20. In order to reduce the occupied area of ​​the additional trace 22 in practice, the conductive width of the trace 22 can be smaller than the loop 20. In addition, the trace 22 can be shared with a portion of the single loop conductor 20 to form a closed DC path for detuning the single loop 20 during the transmit phase of the whole body coil 8. In one embodiment, the trace 22 can be part of the DC path. In another embodiment, the trace 22 can be a post along the conductive trace. The DC voltage source (not shown) is appropriately placed in the L4 position because the L4 value is important for receiving |B 1 |Correction of field distortion has minimal sensitivity.

[0029] To calibrate the MR receive coil 18 using this simulation (or alternatively, using a physical setup, e.g., a physical instantiation of the coil 18 and a physical body 24 made of a biological or synthetic material having dielectric properties close to those of the patient), the single loop 20 is tuned to a 128 MHz resonance (MRI at 3 T, or more generally to a B based on the design) using four capacitors C1-C4 of 18 pF (in this particular example). 0 Next, four inductors L1-L4 are placed along the DC path: L1 (1.8 μH) and L2 (1.8 μH) are located at both ends of trace 22, and L4 (1.8 μH) is placed in a single loop 20 in parallel with C4. Inductor L3 is at the midpoint of trace 22 and is then tuned to optimally suppress the received |B 1 | Field distortion. In a specific example, an inductance value of inductor L3 = 1.69 μH was found to be suitable for this purpose.

[0030] In operation, during the transmit phase of the whole body coil 8, the trace 22 and the four inductors L1-L4 provide a closed DC path to control the bias diodes for detuning the single loop 20. During the receive phase of the receive coil 18, the DC path serves as a diode for providing receive |B 1 | Additional conductive traces for field uniformity correction. In practice, a resistor R is optionally used in the DC path to adjust the bias DC current. 1 |The resistance of resistor R is adjusted for both field uniformity correction and a satisfactory DC bias current. In this example, R is selected to have a value of 40Ω and is located in conductive trace 22 next to inductor L3, but resistor R may also be placed elsewhere in conductive trace 22.

[0031] In order to demonstrate that the disclosed MR coil 18 receives |B by tuning 1 |Field uniformity enhancement inductor L3 provides for receiving |B 1 | Advantageous suppression of field distortion, using Figure 3 The same simulation setup of was used to model a conventional single loop coil without additional conductive traces. The conventional loop was also tuned to the same resonance of 128 MHz with the same four capacitors of 18 pF. Both MR receive coil loop designs were simulated by placing the MR receive coil loop 10 mm above a cylindrical uniform phantom 24 with a diameter of 150 mm having a conductivity of 0.6 S / m and a relative permittivity of 78 to simulate the dielectric properties of the patient. The phantom 24 was placed along the Figure 3The Y direction is shown to have a length of 120 cm. A radio frequency (RF) voltage source V (not shown) was placed parallel to capacitor C3 and a simulated signal-to-noise ratio (SNR) was calculated on the central transverse slice of the phantom 24 for comparison.

[0032] Figure 4 A conventional single loop (without the receiver disclosed herein) is shown. 1 | Field uniformity correction; upper plot) and the SNR calculated by the disclosed MR coil 18, the MR coil 18 includes a loop 20 and a conductive trace 22, wherein, as previously described, the receiving | B 1 |Field uniformity enhancement inductor L3 is tuned to 1.69μH (lower plot). Figure 4 The SNR line at a depth of y = 12 mm from the phantom surface along the x-axis is plotted in FIG. 1 |The field uniformity corrected MR coil 18 (lower plot) has improved left-right SNR uniformity compared to the conventional loop (upper plot).

[0033] Figure 5 A table showing the SNR ratio values ​​of the MR coil 18 compared to the conventional loop coil in the central transverse slice (Z=0 plane) of the phantom 24 shows that the SNR is improved (>1) in most areas. Each box represents a spatial location in the XY plane, and the value marking the box is the value with the received |B 1 | SNR of the MR coil 18 with field uniformity correction and without receiving |B 1 |Ratio of the SNR of the field uniformity corrected MR coil 18. Note that the bolded SNR value highlights |B 1 |Field uniformity correction improves SNR uniformity.

[0034] refer to Figure 6 And continue to refer to Figure 1 and Figure 2 , will be used to optimize the reception of the MR coil 18 |B 1 | The illustrative method 100 of field uniformity is schematically shown as a flow chart. To begin the method 100, at operation 102, the MR receive coil 18 to be tuned is placed on the B of the MR scanner or design basis. 0 Other sources of magnetic field include placing the coil 18 on a dielectric phantom 24 (see Figure 3 ) and the coil 18 is about the B 0 Magnetic field orientation, as the coils 18 would be in the case of the coil array 16 (see Figure 1 At operation 102, the MR coil 18 is placed on the phantom 24 so that the MR coil 18 is placed at a design level of at least 3 Tesla.0 The MR receiving coil 18 is coupled to the dielectric body 24 (ie, the phantom). At operation 104, one or more components of the MR receiving coil 18 are adjusted. For example, at least the receiving coil 18 is adjusted. 1 | Field uniformity enhancement inductor L3 is placed on the basis of the design to optimize the B 0 The receiving coil 18 of the MR receiver in the magnetic field and coupled to the dielectric body receives the 1 |Field uniformity. For another example, adjusting the resistor R in the trace 22 to optimize the reception of the MR receiving coil 18 |B 1 | Field uniformity and satisfying the DC bias current for detuning the MR receive coil 18. The impedance of the additional inductors L1, L2 can be fixed. The adjustment operation 104 can be performed on a phantom, or by electromagnetic simulation.

[0035] The present disclosure has been described with reference to the preferred embodiments. Modifications and alterations may occur to others upon reading and understanding the preceding detailed description. The exemplary embodiments are intended to be interpreted as including all such modifications and alterations as long as they fall within the scope of the appended claims or their equivalents.

Claims

1. A magnetic resonance (MR) receiving coil (18), comprising: A conductive loop (20) resonating at a Larmor frequency of a B0 magnetic field based on a design of at least 3 Tesla; as well as A conductive trace (22) configured to detune the conductive loop from the Larmor frequency in response to a DC current flowing through the conductive trace, the conductive trace also including a receive |B1| field uniformity enhancement inductor (L3) positioned at an intermediate point along the conductive trace.

2. The MR receiving coil (18) according to claim 1, wherein: The receive |B1| field uniformity enhancing inductor (L3) is tuned to optimize the receive |B1| field uniformity of the MR receive coil when the MR receive coil is placed in the B0 magnetic field on the design basis, wherein the branch of the conductive trace (22) includes the receive |B1| field uniformity enhancing inductor oriented parallel to the B0 magnetic field on the design basis.

3. The MR receiving coil (18) according to any one of claims 1 and 2, wherein: The receive |B1| field uniformity enhancing inductor (L3) is positioned at the midpoint of the conductive loop (20).

4. The MR receiving coil (18) according to any one of claims 1 to 3, wherein: The conductive trace (22) extends along a portion of the conductive loop (20) and parallel to the conductive loop.

5. The MR receiving coil (18) according to any one of claims 1 to 3, wherein: The conductive trace (22) extends along one half of the conductive loop (20) and parallel to the conductive loop.

6. The MR receiving coil (18) according to any one of claims 1 to 5, wherein: In addition to the receive |B1| field uniformity enhancing inductor (L3), the conductive trace (22) includes at least two additional inductors (L1, L2).

7. The MR receiving coil (18) according to any one of claims 1 to 6, wherein: The at least two additional inductors (L1, L2) include: Two coupled inductors (L1, L2) connect opposite ends of the conductive trace (22) to the conductive loop (20).

8. The MR coil (18) according to any one of claims 1 to 7, wherein: The conductive loop (20) includes at least one capacitor (C1-C4).

9. The MR coil (18) according to claim 8, wherein: The at least one capacitor (C1-C4) includes a capacitor (C4) co-located in the conductive loop with the receive |B1| field uniformity enhancing inductor (L3).

10. A method (100) for optimizing receive |B1| field uniformity of a magnetic resonance (MR) receive coil (18), the method comprising: placing the MR receiving coil in a design-based B0 magnetic field of at least 3 Tesla, wherein the MR receiving coil is coupled to a dielectric body, the MR receiving coil comprising: A conductive loop (20) resonates at the Larmor frequency of the B0 magnetic field on a design basis, and a conductive trace (22) configured to detune the conductive loop from the Larmor frequency in response to a DC current flowing through the conductive trace, the conductive trace further comprising a receive |B1| field uniformity enhancing inductor (L3) positioned on a portion of the conductive trace oriented parallel to the B0 magnetic field on the design basis; and One or more components of the MR receive coil, including at least the receive |B1| field uniformity enhancing inductor (L3), are adjusted to optimize the receive |B1| field uniformity of the MR receive coil placed in the B0 magnetic field based on the design and coupled to the dielectric body.

11. The method (100) according to claim 10, further comprising: The impedance of the receive |B1| field uniformity enhancing inductor (L3) is adjusted while the impedance of the additional inductors (L1, L2) of the conductive trace (22) is fixed.

12. The method (100) according to any one of claims 10 and 11, wherein: Adjustments are performed on a phantom or by electromagnetic simulation.

13. The method (100) according to any one of claims 10-12, further comprising: The receive |B1| field uniformity enhancing inductor (L3) is tuned to optimize the receive |B1| field uniformity of the MR receive coil when the MR receive coil is placed in the B0 magnetic field on the design basis, wherein the branch of the conductive trace (22) includes the receive |B1| field uniformity enhancing inductor oriented parallel to the B0 magnetic field on the design basis.

14. The method (100) according to any one of claims 10-13, further comprising: The receive |B1| field uniformity enhancing inductor (L3) is positioned at the midpoint of the conductive loop (20).

15. The method (100) according to any one of claims 10-13, further comprising: The conductive trace (22) is positioned along a portion of the conductive loop (20) extending parallel to the conductive loop.

16. The method (100) according to any one of claims 10-13, further comprising: The conductive trace (22) is positioned to extend parallel to the conductive loop along one half of the conductive loop (20).

17. The method (100) according to any one of claims 10 to 16, wherein: In addition to the receive |B1| field uniformity enhancing inductor (L3), the conductive trace (22) includes at least two additional inductors (L1, L2).

18. The method (100) according to any one of claims 10 to 17, wherein: The at least two additional inductors (L1, L2) include: Two coupled inductors (L1, L2) connect opposite ends of the conductive trace to the conductive loop (20).

19. The method (100) according to any one of claims 10-18, wherein: The conductive loop (20) includes at least one capacitor (C1-C4).

20. The method (100) of claim 19, wherein: The at least one capacitor (C1-C4) includes a capacitor (C4) co-located in the conductive loop with the receive |B1| field uniformity enhancing inductor (L3).

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