Numerical evaluation of the optimal coupling scheme of a cylindrical dielectric resonator operating at 600 MHz (14T)
Wei Luo1, Rui Liu2, Thomas Neuberger3,4, and Michael T Lanagan1,2

1Material Research Institute, University Park, PA, United States, 2Department of Engineering Science and Mechanics, University Park, PA, United States, 3Huck Institute of Life Science, University Park, PA, United States, 4Department of Biomedical Engineering, University Park, PA, United States

### Synopsis

To maximize the energy transfer to the cylindrical dielectric resonator utilized in magnetic resonant imaging probe head, a three-loop coupling method was investigated using electromagnetic field simulations. The simulation results demonstrate the supreme performance of this coupling method and verify the previous preliminary experimental results.

### Targeted audience

Engineers who are interested in utilizing a dielectric resonator as a MRI coil.

### Introduction

Cylindrical dielectric resonators (CDR) operating in TE01δ mode and HEM11δ mode have been used in recent years as MRI probes in high-field magnetic resonance imaging (MRI) (B0 ≥ 7T)[1-5]. Electromagnetic field simulations were used to optimize the power transfer to the CDRs and the results were compared to experimental data [6]. A three-loop inductive coupling method is shown by both experimental[6] and simulation results to have the best coupling to the TE01δ mode of the CDR.

### Methods

An MRI probe made of CDR with a three-loop inductive coupling method[6] was modeled (Figure 1$b$) and simulated using the FDTD method. The CDR ($ε_r$ = 173, outer diameter = 46.1 mm, height = 33.7 mm) has a hollow bore (diameter: 5.32 mm) in its center for sample placement. A three-turn copper solenoid coil (height: 27 mm, diameter: 48 mm, wire diameter: 0.69 mm) was wrapped around the CDR for coupling. The second simulated CDR MRI probe was coupled by a single loop coil (diameter: 14.8 mm wire diameter: 1 mm) as shown in Figure 1$a$. In this model, the CDR kept the same geometry but its $ε_r$ is lowered to 160 which would provide some tunability at 14T in this setup. The single loop coupling coil was located right beneath the CDR. In both models, a cylindrical ‘muscle’ phantom (diameter: 4.32 mm, length: 35 mm, $ε_r$ = 56, $σ$ = 0.85 S/m) was placed at the center of the CDR. Both MRI probes with muscle phantom were positioned at the center of a copper shield (diameter: 55 mm, length: 150 mm) and tuned to 600 MHz using two copper pieces located on the top and bottom faces of the CDR (not shown in Figure 1). The coil efficiency, $|B_1^+|\sqrt{P_{dissp}}$, and SNR maps in the phantom were calculated and compared. $|B_1^+|$ is the magnitude of the larger circularly‐polarized magnetic component of the EM field generated by the CDR MRI probe and $\sqrt{P_{dissp}}$ is the total power dissipated in the phantom, coupling coil, CDR, and copper shield. Their center line profile were also extracted and analyzed. All EM field simulations were performed in a commercial package (XFdtd, Remcom, USA).

### Results and Discussion

The coil efficiency (Effcoil) and normalized SNR maps (FA = 90˚) from both CDR MRI probes in the cross section of the phantom along the phantom axis is shown in Figure 2 and Figure 3. Both, the Effcoil and the SNR from the MRI probe with the three-loop inductive coupling method (Figure 2$b$ and Figure 3$b$) are higher at every point across the phantom than the one using a single coupling loop (Figure 2$a$ and Figure 3$a$). The intensity distributions of the Effcoil and SNR from both MRI probes are similar; the maximum locates at the phantom center and the value falls off towards to the ends of the phantom. This also shows in their center line profiles (Figure 4 and Figure 5). Furthermore, Figure 4 and Figure 5 show that, with the three-loop inductive coupling method, the maximum improvement of Effcoil and SNR are more than order of magnitudes at both ends of the phantom; the minimum improvements are about 11% in Effcoil and 10% in SNR at the phantom center, respectively. The optimal imaging length, where its SNR < 0.95×max(SNR) within the phantom, increases more than 46% from 10.5 mm for the MRI probe using a single coupling loop to 15.4 mm for the one using the three-loop inductive coupling method. In conclusion, we expect a better imaging quality across the full-length of the phantom using the CDR MRI probe with the three-loop inductive coupling method. The maximum improvement would be located at both ends of the MRI probe and a maximum of 10% increase is expect at the probe center.

### Acknowledgements

NIH grant R24 MH106049 and NSF grant DBI-1353816

### References

1] Haines et al., JMR, 2002;200:349-53. [2] Neuberger et al., CONCEPT MAGN RESON B, 2002;33B:109-14. [3] Aussenhofer et al., MRM, 2004;68:1325-31. [4] Aussenhofer et al., NMR Biomed, 2011;26:1555-61. [5] Aussenhofer et al., JMR, 2014;243:122-9. [6] Liu et al., ISMRM 2015;23:3102.

### Figures

Figure 1: Simulation models of CDR MRI probe with single coupling coil ($a$) and three-loop coupling method ($b$). The muscle phantom (dark blue) is position in the center of both probes. The shield is partially shown (yellow) for illustration purpose.

Figure 2: The Effcoil of the CDR MRI probe with single coupling coil method ($a$) and the one with three-loop coupling method ($b$).

Figure 3: The SNR maps using the CDR MRI probe with single coupling coil method ($a$) and the one with three-loop coupling method ($b$).

Figure 4: The center line profile of the Effcoil shown in Figure 2.

Figure 5: The center line profile of the SNR maps shown in Figure 3.

Proc. Intl. Soc. Mag. Reson. Med. 24 (2016)
0388