Conductive loop having an inductive path for a magnetic resonance imaging (MRI) receive coil

Abstract

The magnetic resonance (MR) receiving coil 18 has a conductive loop 20 that resonates at the Larmor frequency of a B0 magnetic field with a design criterion of at least 3 teslas, 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 has a receiving |B1| magnetic field uniformity improvement inductor L3 disposed at an intermediate point along the conductive trace.

Description

Technical Field

[0001] The present invention generally relates to magnetic resonance (MR) imaging technology, MR coil technology, high magnetic field MR signal acquisition technology, MR decoupling technology, and related technologies.

Background Art

[0002] In magnetic resonance (MR) imaging, a subject (such as a medical patient, a veterinary subject, an archaeological mummy, etc.) is placed within a static magnetic field (often called the B0 magnetic field), nuclear magnetic resonance is excited within the subject, and the excited magnetic resonance is detected. For imaging, during excitation, during the time interval between MR excitation and MR readout, and / or during MR readout, a magnetic field gradient is superimposed on the static B0 magnetic field, so that the excited MR signal is spatially encoded with respect to position, phase, and / or frequency. In a typical design, an MR imaging device (sometimes called an MRI scanner) has a housing with a central bore within which the MR examination area is disposed.

Summary of the Invention

Problems to be Solved by the Invention

[0003] The static B0 magnetic field is generated by solenoid magnet windings wound around the central bore and housed within the MR scanner housing. These solenoid magnet windings are often superconducting windings in today's MRI scanners, and the housing has a liquid helium (LHe) reservoir for cooling the superconducting windings. The magnetic field gradient coils are also disposed within the housing around the central bore.

[0004] In the case of a human subject, a body coil is used that is arranged concentrically around the bore to provide MR excitation, and such a body coil is generally a cylindrical birdcage coil, a TEM coil, or a variation thereof. Alternatively, a local coil arranged near the anatomical structure of the body to be imaged is used for excitation. MR readout is typically performed using a local MR receive coil arranged near the anatomical structure to be imaged. The local MR receive coil and the local MR excitation coil (if used) may be the same coil or different coils. For various reasons, the MR receive coil (and, if used, the MR excitation coil) can have an MR coil including one or more coil elements, each coil element being configured as a loop coil, although other coil element designs are also known.

[0005] In today's coil designs, a single loop is used in a B0 magnetic field with a design criterion of at least 3 tesla. However, such a single loop used in a B0 magnetic field of 3T or higher is affected by receive |B1| magnetic field distortion, and the MRI signal-to-noise ratio (SNR) decreases.

[0006] The following discloses specific improvements to overcome these and other problems.

Means for Solving the Problems

[0007] In some embodiments disclosed herein, a magnetic resonance (MR) receive coil has a conductive loop that resonates at the Larmor frequency of a B0 magnetic field with a design criterion of at least 3 tesla, and a conductive trace configured to detune the conductive loop from the Larmor frequency in response to a direct current flowing through the conductive trace. The conductive trace further has a receive |B1| magnetic field uniformity improvement inductor disposed at an intermediate point along the conductive trace.

[0008] In some embodiments disclosed herein, a method for optimizing the received |B1| magnetic field uniformity of an MR receive coil includes placing the MR receive coil in a B0 magnetic field of at least 3 tesla with the MR receive coil coupled to a dielectric. The MR receive coil has a conductive loop that resonates at the Larmor frequency of the B0 magnetic field of the design criteria 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, and the conductive trace further has a receive |B1| magnetic field uniformity improvement inductor disposed on a portion of the conductive trace oriented parallel to the design criteria B0 magnetic field. The method further includes adjusting one or more components of the MR receive coil having at least the receive |B1| magnetic field uniformity improvement inductor to optimize the received |B1| magnetic field uniformity of the MR receive coil disposed in the B0 magnetic field of the design criteria and coupled to the dielectric.

[0009] One advantage is to provide an MRI receive coil having reduced received |B1| magnetic field distortion.

[0010] Another advantage is to provide an MRI receive coil having a conductive trace disposed on one side of a single-loop coil.

[0011] Another advantage is to increase the SNR of the MR scanner.

[0012] Another advantage is to provide an MRI receive coil having a conductive trace with a receive |B1| magnetic field uniformity improvement inductor, and such receive |B1| magnetic field uniformity improvement inductor is disposed on a portion of the conductive trace oriented parallel to the design criteria B0 magnetic field.

[0013] Certain embodiments may not provide any of the above advantages, may provide one, two, three or more, or all of the above advantages, and / or may provide other advantages that will be apparent to those skilled in the art upon reading and understanding the present disclosure.

[0014] The present disclosure can take the form of various components and combinations of components, as well as various steps and combinations of steps. The drawings are for the purpose of illustrating preferred embodiments only and should not be construed as limiting the present disclosure.

Brief Description of the Drawings

[0015]

Figure 1

Figure 2

Figure 3

Figure 4

Figure 5

Figure 6

Embodiments for Carrying Out the Invention

[0016] For example, at a high B0 magnetic field of 3T or more, the distortion of the │B1│ magnetic field caused by high frequency (RF) becomes a problem. For example, at 3T, the magnetic resonance (MR) resonance frequency is 128 MHz corresponding to an RF wavelength of 2.34 meters in air. This wavelength is shortened inside the patient to a length comparable to the cross-sectional diameter of the patient due to the dielectric effect. The shortened wavelength causes structural and destructive interference of the │B1│ magnetic field inside the patient that distorts the │B1│ magnetic field.

[0017] The following discloses the use of conductive copper traces sometimes included in an MR receive coil to perform detuning of an MR receive coil loop during a transmit phase in order to suppress received │B1│ magnetic field distortion and enhance the uniformity of the received │B1│ magnetic field. In some embodiments, additional inductors are included in the copper wire, and in some embodiments, they are arranged at symmetric positions along the z-direction (the direction of the B0 magnetic field). The MR receive coil can further have end inductors L1 and L2 for a DC path through for performing detuning of the MR receive coil loop during the transmit phase from time to time.

[0018] In addition to providing additional inductors, such as center inductors at the symmetry center, for suppressing received │B1│ magnetic field distortion, the following discloses a method for calibrating a coil loop to enhance received │B1│ magnetic field uniformity. In one approach, the inductors of the MR receive coil (excluding the inductors added for suppressing received │B1│ magnetic field distortion) have a fixed inductance value (e.g., 1.8 microhenries in one non-limiting exemplary embodiment), and the calibration method includes adjusting the inductance of the additional inductor to optimize the received │B1│ magnetic field uniformity. A phantom representing the dielectric mass of a patient can be used in the calibration method, or alternatively, electromagnetic simulations of the coil and the dielectric mass can be used. The disclosed calibration approach in some embodiments involves adjusting a center inductor symmetrically arranged on the coil loop along the z-direction. The optimization is an empirical process.

[0019] Receiving │ B1 │ Calibration of the coil loop to optimize magnetic field uniformity is typically performed once at the factory where the MR receiving coil is manufactured. The calibrated coil loop can then be installed as one of the coil loops of an MR coil having an array of such coil loops. When the MR coil is used for patient imaging, the calibrated MR coil loop should be positioned with respect to the B0 magnetic field and the patient in the same spatial orientation as was used in the calibration. In practice, the calibrated coil loop is typically part of a coil array (e.g., a head coil array or a body coil array) that is sized and shaped to couple to a particular part of the human anatomical structure, and thus, proper positioning of the coil loop in clinical use is naturally achieved.

[0020] Referring to FIG. 1, an exemplary magnetic resonance (MR) imaging system or apparatus 10 has a magnetic resonance (MR) imaging scanner, such as an MR scanner that generates a magnetic field of at least 3 Tesla (3T). As shown in FIG. 1, the MR apparatus 10 has a housing or gantry 2 having various components shown in FIG. 1, for example, by way of non-limiting illustration, a superconducting magnet 4 that generates a static magnetic field (B0) schematically shown in FIG. 1, a magnetic field gradient coil 6 for superimposing a gradient magnetic field on the B0 magnetic field, a whole-body radio frequency (RF) coil 8 for applying an RF pulse to excite and / or spatially encode magnetic resonance in an imaged patient disposed in the MR bore 12 or other MR examination area, and / or others. The magnet 4 and the gradient magnetic field coil 6 are arranged concentrically around the bore 12. A patient robot bed 14 or other patient support enables a medical patient, a patient undergoing a medical screening, or other patient being imaged to be loaded into the MR bore 12 for imaging. The magnetic resonance excited in the object being imaged is read out by the MR receive coil 18. FIG. 1 shows that an exemplary MR coil 16 is an array of coil loops 18, i.e., the MR receive coil 18. It will be understood that the coil array 16 can generally have 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 within the examination area (i.e., the MR bore 12), as shown in FIG. 1.

[0021] Typically, the MR coil 16 is arranged in a fixed orientation with respect to the B0 magnetic field generated by the magnet 4. For example, an exemplary coil 16 is a body coil that is placed on the patient's torso in a fixed orientation. As a result, each MR receiving coil 18 that makes up the coil array 16 is also typically placed on the patient in a fixed spatial orientation corresponding to the B0 magnetic field. It should also be noted that the MR coil array 16 (and thus the MR receiving coil loop 18 that is its component) is placed on or in close proximity to the patient. This advantageously provides a strong electromagnetic coupling between the patient and the MR receiving coil loop 18. However, in some embodiments, at high magnetic fields such as |B0| = 3 tesla or more, magnetic field distortion may occur due to the dielectric effect of the patient. As described above, this is the result of the wavelength being shortened inside the patient at high magnetic fields such that it becomes comparable to the cross-sectional diameter of the patient.

[0022] Referring now to FIG. 2, one exemplary MR receive coil 18 of the MR coil (array) 16 of FIG. 1 is shown. The MR receive coil 18 has, for example, a conductive loop 20 that resonates at the Larmor frequency of a B0 magnetic field with a design criterion of at least 3T. The exemplary conductive loop 20 is a single loop of copper, a copper alloy, or another conductive material, and is formed, for example, as a copper layer deposited on a circuit board, a plastic sheet, a plastic former, or other electrical insulating substrate, or as an independent metal loop. The conductive loop 20 has at least one capacitor (four of which are labeled C1 - C4 and shown in FIG. 2), and in an exemplary embodiment, further has an inductor (labeled L4 in FIG. 2) disposed in parallel across the ends of capacitor C4, forming an LC detuning circuit for detuning the MR receive coil 18 during the transmit phase of the whole body coil 8. The exemplary conductive loop 20 is rectangular, although other geometric shapes, such as circular, are contemplated. It should be further understood that the bulk of the conductive loop 20, such as a copper layer, can have gaps at the positions of the capacitors C1 - C4, and the connections of the capacitors across these gaps collectively form the conductive loop 20. The conductive loop 20 is conductive (and resonant) at the Larmor frequency, but can be non - conductive in the DC conductive loop 20 due to the presence of the capacitors C1 - C4 that act as blocking capacitors to block the flow of DC current around the conductive loop 20.

[0023] Also, the MR coil 18 has a conductive trace 22 separate from the conductive loop 20, and such conductive trace 22 is arranged alongside the conductive loop 20 at a distance close enough to enable an inductive coupling between the conductive loop 20 and the conductive trace 22 at the Larmor frequency (and in its vicinity). 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. As shown in FIG. 2, the conductive trace 22 runs parallel to the conductive loop 20 along a part of the conductive loop 20. In an exemplary embodiment, the conductive trace 22 runs parallel to the conductive loop 20 along half of the conductive loop 20, although deviations from this half - portion are contemplated.

[0024] The conductive trace 22 has a plurality of inductors (denoted as L1 to L3 in FIG. 2). A plurality of (exemplarily two) inductors (e.g., L1 and L2) can be coupling inductors that connect both ends of the conductive trace 22 to the conductive loop 20. At least one of the inductors (exemplary inductor L3) has a reception |B1| magnetic field uniformity improvement inductor. The reception |B1| magnetic field uniformity improvement inductor L3 is adjusted to optimize the reception |B1| magnetic field uniformity of the MR reception coil 18 when the MR reception coil 18 is disposed within the B0 magnetic field of the design criteria, and the leg of the conductive trace 20 has a reception |B1| magnetic field uniformity improvement inductor that is oriented in parallel with the B0 magnetic field of the design criteria. In the exemplary embodiment of FIG. 2, the MR reception coil loop 18 is disposed in the full coil array 16 (see FIG. 1), and in this exemplary embodiment, the B0 magnetic field is oriented in parallel with the leg of the conductive trace 20 having the inductor L3. With respect to the exemplary X-Z coordinate system shown in FIG. 2, this corresponds to the B0 magnetic field being oriented in the Z direction, and similarly, the leg of the MR reception coil loop 18 having the reception |B1| magnetic field uniformity improvement inductor L3 is also oriented in the Z direction in parallel with the B0 magnetic field.

[0025] The reception |B1| magnetic field uniformity improvement inductor L3 is disposed at an intermediate point along the conductive trace 22. In one example, the reception |B1| magnetic field uniformity improvement inductor L3 is disposed at an intermediate point of the conductive loop 20. In another example, the reception |B1| magnetic field uniformity improvement inductor L3 is disposed together with one of the capacitors of the conductive loop 20 (i.e., capacitor C4 as shown in FIG. 2).

[0026] Figures 3 - 5 show examples of experimental data of the MR coil 18. Figure 3 shows a simulation setup of the MR coil 18 having a single conductive loop 20 disposed on a simulated dielectric or mass 24 that simulates the effect on a patient, with a conductive trace 22 disposed on the left half of the conductive loop 20. The simulated coil 18 is oriented such that a leg including a capacitor C4 is oriented parallel to the B0 magnetic field (along the Z direction using an exemplary X - Y - Z coordinate system) adjacent to the receive |B1| magnetic field uniformity improvement inductor L3. The loop 20 simulated in the embodiments of FIGS. 4 and 5 has dimensions of 92 mm along the z - axis (main magnet B0 direction) and 112 mm along the x - axis (patient's left - right direction). The width of the conductive loop 20 is 6 mm and the width of the conductive trace 22 is 2 mm. In this embodiment, the conductive trace 22 is disposed inside the loop 20 along the leg having the capacitor C4, and the conductive trace 22 is disposed 2 mm away from the conductive loop 20. In another embodiment, the conductive trace 22 can be disposed below or above the single conductive loop 20. In practice, the width of the conductive trace 22 can be made narrower than the conductive loop 20 in order to reduce the installation area of the additional trace 22. Further, the conductive trace 22 can share a portion of the single conductive loop 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 conductive trace 22 can be part of the DC path. In another embodiment, the conductive trace 22 can be a stand - alone conductive trace. Since the L4 value has the lowest sensitivity for correcting the receive |B1| magnetic field distortion, a DC voltage source (not shown) is appropriately placed at the position of L4.

[0027] To calibrate the MR receive coil 18 using this simulation (or using a physical setup such as a physical instance of the coil 18 and a physical body 24 made of a biological or synthetic material with dielectric properties close to those of the patient), the single conductive loop 20 is tuned to a resonance of 128 MHz (adjusted to the Larmor frequency for 3T MRI, or more generally to the design reference B0 magnetic field strength), and (in the case of this particular example) the four capacitors C1 - C4 are set to 18 pF. Next, the four inductors L1 - L4 are placed along the DC path, with L1 (1.8 μH) and L2 (1.8 μH) at the two ends of the conductive trace 22, L4 (1.8 μH) placed in parallel with C4 within the single loop 20. Inductor L3 is at the midpoint of the conductive trace 22 and is adjusted to optimally suppress the received |B1| magnetic field distortion. In a specific example, it has been found that an inductance value of 1.69 μH for inductor L3 is suitable for this purpose.

[0028] During 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 for controlling the bias diode due to detuning of the single conductive loop 20. During the receive phase of the receive coil 18, the DC path functions as an additional conductive trace for providing received |B1| magnetic field uniformity correction. In practice, a resistor R is optionally used in the DC path to adjust the bias DC current. The resistance of the resistor R can be adjusted to optimize the received |B1| magnetic field uniformity correction and to satisfy the DC bias current. In this example, the resistance of the resistor R is selected to be a value of 40 Ω and is placed within the conductive trace 22 next to the inductor L3, although the resistor R can also be placed at other locations within the conductive trace 22.

[0029] To demonstrate the advantageous suppression of received |B1| magnetic field distortion provided by the disclosed MR coil 18 having the received |B1| magnetic field uniformity improvement inductor L3 adjusted as such, using the same simulation settings as in FIG. 3, a conventional single-loop coil is modeled without additional conductive traces. The conventional loop is also tuned to the same resonance of 128 MHz using the same four capacitors of 18 pF. Both MR receive coil loop designs are simulated by being placed 10 mm above a cylindrical uniform phantom 24 with a diameter of 150 mm that simulates the dielectric properties of the patient by having respective conductivity of 0.6 S / m and relative permittivity of 78. The phantom 24 has a length of 120 cm along the Y direction shown in FIG. 3. A high-frequency (RF) voltage source V (not shown) is placed in parallel with capacitor C3, and the simulated signal-to-noise ratio (SNR) is calculated across the central transverse slice of the phantom 24 for comparison.

[0030] FIG. 4 shows the calculated SNR of a conventional single loop (without the received |B1| magnetic field uniformity correction disclosed herein, upper figure) and the calculated SNR of the disclosed MR coil 18 having loop 20 and conductive trace 22 (having the received |B1| magnetic field uniformity improvement inductor L3 adjusted to 1.69 μH as described above, lower figure). The SNR lines along the x-axis at a depth of y = 12 mm from the surface of the phantom are plotted in FIG. 4. As can be seen from the figure, the MR coil 18 with received |B1| magnetic field uniformity correction (lower figure) has improved left-right SNR uniformity compared to the conventional loop (upper figure).

[0031] FIG. 5 is a table showing the values of the ratio of the SNR of the MR coil 18 to that of a conventional loop coil across the central transverse slice (Z = 0 plane) of the phantom 24, which shows SNR improvement (>1) over most of the region. Each box represents a spatial position in the X - Y plane, and the value labeling the box is the ratio of the SNR of the MR coil 18 with received |B1| magnetic field uniformity correction to the SNR of the MR coil 18 without received |B1| magnetic field uniformity correction. Of note are the bold SNR values that highlight where the SNR uniformity has been improved by the received |B1| magnetic field uniformity correction.

[0032] Referring to FIG. 6 and continuing to refer to FIGS. 1 and 2, an exemplary method 100 for optimizing the received |B1| magnetic field uniformity of the MR coil 18 is schematically shown as a flowchart. To start method 100, in process 102, the MR receive coil 18 to be adjusted is placed within an MR scanner or other source of a B0 magnetic field of a design criterion, and the receive coil 18 is within the coil array 16 (see FIG. 1), placed on the dielectric phantom 24 (see FIG. 3), and oriented with respect to its B0 magnetic field. In process 102, the MR coil 18 is placed on the phantom such that the MR coil 18 is placed within a B0 magnetic field of a design criterion of at least 3 teslas, and the MR receive coil 18 is coupled to the dielectric 24 (i.e., the phantom). In process 104, one or more components of the MR receive coil 18 are adjusted. For example, at least the receive |B1| magnetic field uniformity improvement inductor L3 is adjusted to optimize the received |B1| magnetic field uniformity of the MR receive coil 18 placed within the B0 magnetic field of a design criterion and coupled to the dielectric. As another example, the register R of the trace 22 is adjusted to optimize the received |B1| magnetic field uniformity of the MR receive coil 18 and to satisfy the DC bias current for detuning of the MR receive coil 18. The impedances of the additional inductors L1, L2 may be fixed. The adjustment process 104 can be performed on the phantom or by electromagnetic simulation.

[0033] The present disclosure has been described with reference to preferred embodiments. Those skilled in the art may come up with modifications and changes upon reading and understanding the foregoing detailed description. The exemplary embodiments are intended to be construed to include all such modifications and changes insofar as they fall within the scope of the appended claims or their equivalents.

Claims

1. A magnetic resonance (MR) receiving coil, comprising: a conductive loop resonating at the Larmor frequency of a B0 magnetic field with a design criterion of at least 3 Tesla; a conductive trace configured to detune the conductive loop from the Larmor frequency in response to a direct current flowing through the conductive trace, the conductive trace further having a receive |B1| magnetic field uniformity improvement inductor disposed at an intermediate point along the conductive trace; The MR receiving coil having the above.

2. The receive |B1| magnetic field uniformity improvement inductor is adjusted to optimize the receive |B1| magnetic field uniformity of the MR receiving coil when the MR receiving coil is disposed in the B0 magnetic field of the design criterion and the legs of the conductive trace including the receive |B1| magnetic field uniformity improvement inductor are oriented parallel to the B0 magnetic field of the design criterion. The MR receiving coil according to claim 1.

3. The MR receiving coil according to claim 1 or 2, wherein the receive |B1| magnetic field uniformity improvement inductor is disposed at an intermediate point of the conductive loop.

4. The MR receiving coil according to any one of claims 1 to 3, wherein the conductive trace extends parallel to the conductive loop along a portion of the conductive loop.

5. The MR receiving coil according to any one of claims 1 to 3, wherein the conductive trace extends parallel to the conductive loop along half of the conductive loop.

6. The MR receiving coil according to any one of claims 1 to 5, wherein the conductive trace has at least two additional inductors in addition to the receive |B1| magnetic field uniformity improvement inductor.

7. The MR receiving coil according to any one of claims 1 to 6, wherein the at least two additional inductors have two coupling inductors connecting both ends of the conductive trace to the conductive loop.

8. The MR receiving coil according to any one of claims 1 to 7, wherein the conductive loop has at least one capacitor.

9. The MR receiving coil according to claim 8, wherein the at least one capacitor has a capacitor of the conductive loop disposed together with the receive |B1| magnetic field uniformity improvement inductor.

10. A method for optimizing the receive |B1| magnetic field uniformity of a magnetic resonance (MR) receiving coil, comprising: disposing the MR receiving coil in a B0 magnetic field with a design criterion of at least 3 Tesla; wherein the MR receiving coil is A conductive loop resonating at the Larmor frequency of the B0 magnetic field of the design criteria, and a conductive trace configured to detune the conductive loop from the Larmor frequency in response to a direct current flowing through the conductive trace, the conductive trace further having a receive |B1| magnetic field uniformity improvement inductor disposed in a portion of the conductive trace oriented parallel to the B0 magnetic field of the design criteria, the conductive trace; having steps; adjusting at least one or more components of the MR receive coil having the receive |B1| magnetic field uniformity improvement inductor so as to optimize the receive |B1| magnetic field uniformity of the MR receive coil disposed in the B0 magnetic field of the design criteria and coupled to the dielectric; having a method.

11. The method according to claim 10, further comprising adjusting the impedance of the receive |B1| magnetic field uniformity improvement inductor while fixing the impedance of the additional inductor of the conductive trace.

12. The method according to claim 10 or 11, wherein the adjustment is performed on a phantom or by electromagnetic simulation.

13. When the MR receive coil is disposed in the B0 magnetic field of the design criteria and the leg of the conductive trace having the receive |B1| magnetic field uniformity improvement inductor is oriented parallel to the B0 magnetic field of the design criteria, further adjusting the receive |B1| magnetic field uniformity improvement inductor so as to optimize the receive |B1| magnetic field uniformity of the MR receive coil. The method according to any one of claims 10 to 12.

14. The method according to any one of claims 10 to 13, further comprising disposing the receive |B1| magnetic field uniformity improvement inductor at an intermediate point of the conductive loop.

15. The method according to any one of claims 10 to 13, further comprising disposing the conductive trace to run parallel to the conductive loop along a portion of the conductive loop.

16. The method according to any one of claims 10 to 13, further comprising disposing the conductive trace to run parallel to the conductive loop along half of the conductive loop.

17. The method according to any one of claims 10 to 16, wherein the conductive trace has at least two additional inductors in addition to the receive |B1| magnetic field uniformity improvement inductor.

18. The method according to any one of claims 10 to 17, wherein at least two additional inductors have two coupling inductors connecting both ends of the conductive trace to the conductive loop. **Claim 19** The method according to any one of claims 10 to 18, wherein the conductive loop has at least one capacitor. **Claim 20** The method according to claim 19, wherein the at least one capacitor has a capacitor of the conductive loop arranged together with the reception |B1| magnetic field uniformity improvement inductor.

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