Decoupling between coils in a flexible phased-array using stacked circumferential shielding
Jhy-Neng Tasso Yeh1 and Fa-Hsuan Lin1

1Institute of Biomedical Engineering, National Taiwan University, Taipei, Taiwan

Synopsis

We propose a flexible coil array design consisting of circumferential shielding stacked on circular coils to achieve robust decoupling between coil when the array is either bended or on a flat plane. Two types of circumferential shielding were tested through numerical simulation. The results demonstrated that both types can have good decoupling on a curved surface with S21 < -17.1 dB. The maximum |B1| for the array coil with Type-II (two-parallel-ring) shielding was 63% larger than that with Type-I (single-strip) shielding. Future work will empirically construct these coils to compensate simulation inaccuracies and validate their performance In Vivo.

PURPOSE

MRI phased-array coil can achieve large FOV and high SNR simultaneously if coils in an array can be sufficiently decoupled. Coil decoupling can be achieved by either overlapping coil elements1 or adding inductors and capacitors between coils2. Note that the anatomical structure can vary significantly among clinical applications, a flexible coil array allowing for a higher filling factor for different subjects is desired3-4. However, decoupling by overlapping coil restricts the mechanical flexibility of the coil array. Coil decoupling by adding inductors and capacitors has the disadvantage of ohmic contact loss. Here we propose to decouple between coils in an phased-array by using circumferential shielding stacked on top of the coil (Fig.1). We hypothesize that such a stacked structure is particularly useful when the coil array may be mechanically bended, because clear gaps between coils allows bending. Also the stacked structure permits exchanged for extendable anatomical range. Here we studied two types of circumferential shielding (Fig.2) and compared their performance using full-wave electromagnetic simulation. Simulation inaccuracies and variations of components can be easily compensated by further fine-tuning the practical coils in the future. Our results demonstrate that both types achieved good isolation between coils (S21 lower than -17.1dB) and provided localized sensitivity B1 regardless of the array was flat or bended.

METHODS

The simulation setup was illustrated in Fig.25. Two decoupled PEC loops were located above the center of the RF coils at right in order to evaluate isolation between coils using S12. A liquid-filled cylindrical phantom was included to simulate the loading, which reduced the quality factor (Q) such speeded up to meet convergence criteria. We simulated two side-by-side circular coils (diameter = 50 mm, width = 5 mm, thickness = 0.02mm) built by copper sheet adhering on FR4 substrate with a 2 mm gap between coils. Each circular coil used three evenly distributed tuning capacitors. The capacitor value was tuned to drive resonance frequency to approach the 1H Larmor frequency (125 MHz) at 3T. The electromagnetic simulation was done using HFSS (Ansys, Canonsburg, USA). All analysis parameters used the following parameters: max ΔS = 0.02, minimal coverage pass = 2, mixed-order basis functions, sweep-type = fast. Two types of circumferential shielding were modeled: type-I (Fig.2C) was a single strip with width = 8 mm6; Type-II (Fig.2D) consisted of two parallel rings (width = 1mm, gap = 8mm). The sizes of both shielding were adjusted to achieve the best decoupling. To demonstrate the flexibility as changing the relative position of two coils on decoupling, each coil array was either placed on the flat top of the phantom or bended to fit to the curvature of the phantom. Maps of |B1| were evaluated by formulas described in Ref.7.

RESULTS

Fig.3 shows |B1| maps when coils were placed on the flat top of the phantom without and with decoupling shielding, respectively. Coupled B1 field was apparent beneath the RF coils at left. The coupled |B1| reduced to 3.1% and 6.8% with type-I and type-II shielding, respectively. The max |B1| beneath the RF coil at right and S21 with type-I and type-II shielding were {6.31x10-6(a.u.), -20.4(dB)} and {8.41x10-6(a.u.), -19.6(dB)}, respectively. Fig.4 shows |B1| maps when coils were bended to fit to the curvature of a 200mm-diameter cylinder. The coupled B1 field without shielding was about 85% of flat-top results in Fig.3. With type-I and type-II shielding, the coupled |B1| reduced to 2.5% and 4.9% and the resonant frequency drifted ≤ 0.75 MHz. The max |B1| beneath the RF coil at right and S21 were {3.74x10-6(a.u.), -18.2(dB)} and {5.66x10-6(a.u.), -17.2(dB)}, respectively. Figs.5 shows the |B1| maps on the curvatures of 180mm and 220mm-diameter cylinders. Since bending an array inevitably changed the relative positions of two coils, the mutual inductance was no longer the same and the coupling between coils was significantly different after such a geometrical deformation. However, both types of circumferential shielding had similar B1 distribution and penetration-depth, successfully maintained the image quality in these scenarios.

DISCUSSION

We used numerical simulation to demonstrate that a flexible coil array consisting of circumferential shielding stacked on circular RF coils can be used to acquire high quality MRI even when the array was mechanically bended. While we only demonstrated an array of two coils, we expect that more channels with circumferential shielding can be integrated into an array in order to increase FOV while maintaining the same decoupling between coils. Also our simulation results suggest that type-II (two-parallel-ring) structure performs better than type-I (single-strip). Future work will empirically construct these coils and validate their performance In Vivo.

Acknowledgements

This study was supported by Ministry of Science and Technology, Taiwan (MOST 104-2314-B-002-238, MOST 103-2628-B-002-002-MY3), National Health Research Institute, Taiwan (NHRI-EX104-10247EI), and Ministry of Economic Affairs, Taiwan (100-EC-17-A-19-S1-175).

References

1. Romer PB, et al. Magn Reson Med. 1990;16:192-225.

2. Lee RF, et al. Magn Reson Med. 2002;48:203-213.

3. Keil B, et al. Magn Reson Med. 2011;66:1777-1787.

4. Duan Y, et al. Proc ISMRM 2013. #4383.

5. Martius S, et al. Proc ISMRM 2015. #3099.

6. Gilbert KM, et al. Magn Reson Med. 2011;67:1487-1496.

7. Collins CM, et al. Magn Reson Med. 2001;45:684-691.

Figures

Fig.1. The flexible coil array design: (A). Two pairs of gapped coil loops (orange color) with stacked circumferential shielding (blue color) illustrate the flexibility of different bending angles. Toy-alike connectors are used to join two parts together. (B). Possible application of flexible phased-array coil on the human head.

Fig.2. 3D model of simulation setups: (A&B). Circular RF coils without stacked shielding structure. (C&D). Circular coils with type-I or type-II circumferential shielding. The cylindrical phantom (yellow color) was included to simulate the human head characteristics (diameter: 200mm, length: 200mm, core σ: 0.9 S/m, εr: 42.9, shell thickness: 5 mm).

Fig.3. |B1| maps when coils were placed on the flat top of a 200mm-diameter cylindrical phantom: (A). circular coils without shielding, (B). with type-I shielding, (C). with type-II shielding. The maximum |B1| for the array coil with type-II shielding was 33% larger than that with type-I shielding.

Fig.4. |B1| maps when coils were bended to fit to the curvature of a 200mm-diameter cylindrical phantom: (A). circular coils without shielding, (B). with type-I shielding, (C). with type-II shielding. The maximum |B1| for the array coil with type-II shielding was 51% larger than that with type-I shielding.

Fig.5. |B1| maps when coils were bended to fit to the curvatures of 180mm (Up) and 220mm (Down) diameter cylindrical phantoms: (A). circular coils with type-I shielding, (B). with type-II shielding. The maximum |B1| for the array coil with type-II shielding was 63% larger than that with type-I shielding.



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