Engineers who are interested in utilizing a dielectric resonator as a MRI coil. Cylindrical dielectric resonators (CDR) operating in TE_01δ mode and HEM_11δ 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 TE_01δ mode of the CDR.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). The coil efficiency (Eff_coil) 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 Eff_coil 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 Eff_coil 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 Eff_coil and SNR are more than order of magnitudes at both ends of the phantom; the minimum improvements are about 11% in Eff_coil 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.