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Making RF coils MR-invisible by additive manufacturing using magnetically filled polymer
Markus Weiger1, Johan Overweg1, Amelie Viol1, Lauro Singenberger1, Thomas Schmid1, Emily Louise Baadsvik1, and Klaas Paul Pruessmann1
1ETH Zurich and University of Zurich, Zurich, Switzerland

Synopsis

Keywords: Non-Array RF Coils, Antennas & Waveguides, Challenges, Materials

Motivation: In MRI of tissues with short T2, materials of RF coils may be detected and can cause background artefacts. Current solutions to this problem either compromise sequence performance or impose restrictions on coil design.

Goal(s): Making RF coils MR-invisible.

Approach: The material used for constructing the coil housing is filled with ferrimagnetic material, leading to effective signal spoiling. Coil formers are created by additive manufacturing using custom filaments made from magnetite-filled polymer.

Results: Unwanted signals from the RF coil are eliminated by using coil formers made from magnetically filled polymer. Hence, background-free short-T2 imaging is enabled.

Impact: RF coils made MR-invisible by using magnetically-filled materials simplify coil design and manufacturing, and improve the performance of MRI of tissues with short T2, such as bone, tendon, lung, or myelin.

Introduction

In MRI, in addition to signals from the targeted object, unwanted signals from parts of the MR scanner – in particular RF coils – may also be detected, and cause artefacts. Usually, for such parts, materials that exhibit relatively fast signal decay are used, preventing detection with common echo times (TE). However, there is continuously growing interest in sequences with ultra-short or zero TE1, which are employed to image tissues with short T2. Therefore, background artefacts from RF coils have become a significant issue1-4.
Approaches to avoid this problem at the sequence level include using a large field of view (FOV)1,5,6, subtraction of background signal7, T2-selective magnetisation preparation8, or outer volume suppression9. However, these methods require sequence modifications, compromise sequence performance, or offer only limited efficacy. Alternatively, materials with minimal 1H content such as glass and PTFE are used4,10-14, but this imposes restrictions for coil design, manufacturing, and handling.
In the present work, a new concept for eliminating background signal from RF coils is proposed, utilising local B0 distortion by magnetic materials. Such signal spoiling has previously been achieved with metallised textiles or by wrapping with magnetic wire12. Here, this idea is exploited more rigorously in that material used for constructing the coil housing is filled with ferrimagnetic microparticles15.
We demonstrate this concept by filling a polymer with magnetite and using it for additive manufacturing of coil formers. Formers with different base materials and filler load were assessed with respect to background signal, coil performance, and field homogeneity. The resulting utility is demonstrated by imaging in a phantom and in vivo.

Methods

To compare formers made from different materials, a 1H-free birdcage coil was designed to be mechanically self-supporting. This allowed for exchanging a hollow cylinder which would normally act as the former (Figure 1a, b). The quadrature, low-pass birdcage was equipped with a matching network allowing for a large loading range.
For the custom-made materials, polylactic acid (PLA, (C3H4O2)n) pellets were heated and extruded to produce filaments for additive manufacturing (Precision 350, 3devo, Utrecht, The Netherlands). For magnetic filling, either 2.5vol% or 5.0vol% of magnetite (Fe3O4) powder (particle size 1-20 µm, Kremer Pigmente, Germany) were added to the PLA pellets before heating.
Experiments were performed using a 3T Philips Achieva system equipped with a high-performance gradient16, a custom RF chain5, and fast transmit-receive switches17. For short-T2 imaging, a PETRA18 sequence with 3D isotropic geometry and a dead time of 10 µs was used.

Results

With all formers, comparable coil performance concerning quality factor, power efficiency and image SNR was found. Magnetic forces on the magnetite-filled formers were weak relative to their own weight.
Figure 1c shows free induction decay signals for four different formers. PMMA, which serves as a reference, exhibits higher signal than PLA. With magnetite filling, the PLA signal amplitude at 10 µs is reduced by 91% and 98%, respectively. Accordingly, in the large-FOV images in Figure 2, the pure PMMA and PLA formers are clearly visible. With magnetite filling, the signal drops to noise level or below.
Figure 3 shows effects of the magnetic material on the static field in the imaging volume after second-order shimming. Field distortion is observed, which is larger than usually aimed at for standard MRI. However, in short-T2 imaging considerably higher gradient strengths are applied, making local resonance offsets less critical.
Given the results of Figures 2 and 3, the material with lower filling load is considered a good compromise between signal reduction and field distortion and was therefore used for successive experiments. Figure 4 demonstrates that in short-T2 imaging with object-matched FOV, aliased background artefacts are successfully eliminated with the proposed approach. Similarly, in-vivo experiments yielded clean images unaffected by background signal (Figure 5).

Discussion

In this work, a new concept for making RF coils MR-invisible was proposed and successfully demonstrated in short-T2 imaging with short dead time and high bandwidth. The particular implementation by filling polymers for additive manufacturing with magnetic material offers high flexibility for coil design and production.
The main side effect of adding magnetic materials is reduced main field homogeneity in the imaging volume. Hence, the choice of the filling load should balance signal suppression and field distortion. However, the effect on homogeneity could be greatly reduced by increasing the former length and profiling its thickness. In addition, the amount of magnetic material could be reduced by choosing base materials with lower proton density or by printing in lower-density mode, including small-scale voids.

Acknowledgements


References

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7. Weiger M, Wu M, Wurnig MC, Kenkel D, Jungraithmayr W, Boss A, Pruessmann KP. Rapid and robust pulmonary proton ZTE imaging in the mouse. NMR Biomed 2014;27:1129-1134.

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11. Rösler MB, Weiger M, Schmid T, Brunner DO, Froidevaux R, Pruessmann KP. Ultrasonic soldering on glass for the construction of MRI coils with minimized background signal in short-T2 images. In Proceedings of the 33rd Annual Scientific Meeting of ESMRMB, Vienna, Austria, 2016. p.87.

12. Weiger M, Brunner DO, Schmid T, Froidevaux R, Rösler MB, Gross S, Pruessmann KP. A virtually 1H-free birdcage coil for zero echo time MRI without background signal. Magn Reson Med 2017;78:399-407.

13. Horch RA, Wilkens K, Gochberg DF, Does MD. RF coil considerations for short-T2 MRI. Magn Reson Med 2010;64:1652-1657.

14. Ozen AC, Ilbey S, Jia F, Idiyatullin D, Garwood M, Nixdorf DR, Bock M. An improved intraoral transverse loop coil design for high-resolution dental MRI. Magn Reson Med 2023;90:1728-1737.

15. Weidenfeller B, Höfer M, Schilling F. Thermal and electrical properties of magnetite filled polymers. Composites Part A: Applied Science and Manufacturing 2002;33:1041-1053.

16. Weiger M, Overweg J, Rösler MB, Froidevaux R, Hennel F, Wilm BJ, Penn A, Sturzenegger U, Schuth W, Mathlener M, Borgo M, Börnert P, Leussler C, Luechinger R, Dietrich BE, Reber J, Brunner DO, Schmid T, Vionnet L, Pruessmann KP. A high-performance gradient insert for rapid and short-T2 imaging at full duty cycle. Magn Reson Med 2018;79:3256-3266.

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Figures

Figure 1: Birdcage RF coil with exchangeable cylindrical former designed to investigate different former materials. a) Quadrature birdcage with low-pass design of 100 mm diameter and 120 mm length including a former, a 1H-free holder made from PTFE, a matching network (T/M), and cable traps. b) The different cylindrical formers of 3 mm thickness and 150 mm length investigated in this work. c) Rapidly decaying FID signals acquired with a former inside the coil but without any other signal source. PLA has lower signal than PMMA, which is strongly reduced for magnetically filled PLA.

Figure 2: Images with large FOV acquired from a doped water bottle (diameter 65 mm, length 140 mm, T2 ≈ 35 ms) using the birdcage coil with different formers. For each former, a transverse and a coronal slice are shown with linear and logarithmic grey scale to emphasise the background. Both PMMA and PLA formers are clearly visible. With lower magnetite filling load, the former signal is nearly removed, and with higher filling load, it is entirely eliminated. Imaging parameters: FOV 210 mm, resolution 1 mm, BW 1 MHz, encoding time 70 µs, TR 1, hard pulse 1 µs, flip angle 3.6°, and scan time 1:24 min.

Figure 3: Field maps acquired in the water bottle with the three PLA formers after second-order shimming. Pure PLA shows slight field deviations caused by the cylindrical former and the bottle. With magnetite filling, the non-uniformity is increased while keeping the same pattern. Furthermore, a bulk field offset occurs (bottom right), which can readily be addressed. Both effects are increased with higher filling load.

Figure 4: Short-T2 imaging in a rubber phantom (T2* ≈ 500 µs) using PLA formers without or with magnetite filling. A FOV of 80 mm was chosen to match to the phantom size. With pure PLA, the signal of the former is aliased into the FOV, leading to background artefacts. With magnetite filling, the PLA signal is sufficiently reduced to provide an entirely clean background. Imaging parameters: resolution 0.45 mm, angular undersampling 1.37, BW 250 kHz, encoding time 346 µs, TR 1 ms, hard pulse 2 µs, flip angle 2.9°, NSA 8, and scan time 9:41 min.

Figure 5: In vivo imaging of the human wrist using the birdcage coil with the PLA former with 2.5% magnetite filling. High quality, background-free images are obtained. The high resolution and SNR obtained for short-T2 tissues enable depiction of tendon and cortical bone with positive contrast. Imaging parameters: FOV 130 mm, resolution 0.55 mm, angular undersampling 1.88, BW 500 kHz, encoding time 226 µs, TR 1 ms, hard pulse 2 µs, flip angle 3.3°, NSA 7, and scan time 10:57 min.

Proc. Intl. Soc. Mag. Reson. Med. 32 (2024)
1564
DOI: https://doi.org/10.58530/2024/1564