- 1Université Paris-Saclay, CEA, CNRS, Inserm, Laboratoire d'Imagerie Biomédicale Multimodale Paris Saclay, Orsay, France
- 2Université Paris-Saclay, CEA, Département des Accélérateurs, de Cryogénie et de Magnétisme, Gif-sur-Yvette, France
- 3Université Paris-Saclay, CNRS, Centre de Nanosciences et de Nanotechnologies, Palaiseau, France
- 4Unité mixte de physique, CNRS, Thales, Université Paris-Saclay, Palaiseau, France
Radiofrequency (RF) coils fashioned from high-temperature superconductor (HTS) have the potential to increase the sensitivity of the magnetic resonance imaging (MRI) experiment by more than a dozen times compared to conventional copper coils. Progress, however, has been slow due to a series of technological hurdles. In this article, we present the developments that recently led to new perspectives for HTS coil in MRI, and challenges that still need to be solved. First, we recall the motivations for the implementations of HTS coils in MRI by presenting the limits of cooled copper coil technology, such as the anomalous skin effect limiting the decrease of the electric resistance of normal conductors at low temperature. Then, we address the progress made in the development of MRI compatible cryostats. New commercially available low-noise pulsed-tube cryocoolers and new materials removed the need for liquid nitrogen-based systems, allowing the design of cryogen-free and more user-friendly cryostats. Another recent advance was the understanding of how to mitigate the imaging artifacts induced by HTS diamagnetism through field cooling or temperature control of the HTS coil. Furthermore, artifacts can also originate from the RF field coupling between the transmission coil and the HTS reception coil. Here, we present the results of an experiment implementing a decoupling strategy exploiting nonlinearities in the electric response of HTS materials. Finally, we discuss the potential applications of HTS coils in bio-imaging and its prospects for further improvements. These include making the technology more user-friendly, implementing the HTS coils as coil arrays, and proposing solutions for the ongoing issue of decoupling. HTS coil still faces several challenges ahead, but the significant increase in sensitivity it offers lends it the prospect of being ultimately disruptive.
Introduction
The discovery of high-temperature superconductivity (HTS) by [1], paved the way for the application of superconductor-based technologies at temperatures above that of boiling nitrogen. Simultaneously, powerful, low vibration, reliable cryocoolers have been developed for applications at temperatures close to 90 K, the critical temperature of high-temperature superconducting YBa2Cu3O7 (YBCO) which is the technologically most useful HTS material. In microwave technology, HTS has allowed the design of detectors and filters of unmatched sensitivity/selectivity and noise figures, suitable for satellite telecommunications [2]. A down-to-earth application of HTS materials is Magnetic Resonance Imaging (MRI). Here, the use of HTS materials can be envisioned both as magnet coils and as probes for the detection of the radiofrequency (RF) MR signal. The vanishing electric resistance of the superconducting MRI coil brings electric noise levels within reach that are unachievable with more traditional copper coils, even if the latter are cooled to much lower temperatures than the HTS coil; typically, the signal-to-noise ratio (SNR) can be increased more than a dozen times [3]. In spite of the promise held, the development of HTS radiofrequency MRI coils, starting from early prototypes in the 90s [4], has been slow. Interest in this technology faded in the 2000s due to technological deadlocks that seemed insurmountable at the time. These included drawbacks such as having to employ wet cryostats (e.g. using liquid nitrogen or LN2), as well as the necessary decoupling of the RF detection coil (fashioned from a HTS material) from the (copper) RF transmission coil. Both issues induce important artifacts in MR images. Recently, however, a number of hurdles have been lifted, heralding new possibilities for HTS coils in MRI. In this paper, we review recent progress and challenges ahead.
First, we review the limitations of copper coils. Cryogenic RF coils become relevant when the dominant noise source in MRI is the RF coil used to collect the signal, i.e. in the case of nonconductive samples, or when conductive samples are probed with “small” surface antennas. With copper coils, noise reduction is readily limited by the intrinsic properties of the material. This limit can be significantly pushed by using superconductors, justifying the implementation for HTS coils in MRI with small samples or small RF coils.
We next present the progress made in the design of MRI compatible cryostats. Such cryogenic systems face the key challenges of ensuring full electromagnetic compatibility in DC and RF domains and user-friendly integration for routine practice in MR exams.
A recent study [5] has led to a much better understanding of how the static field
We then address the subject of the RF
Finally, we discuss the perspectives for further improvements of HTS coil technology in view of novel MRI applications.
Limits of Copper Coils in Magnetic Resonance Imaging
Cooling the RF coil circuit at cryogenic temperatures to improve the sensitivity of nuclear magnetic resonance (NMR) detection has been an early idea [6]. From the fluctuation-dissipation theorem, the noise spectral density generated by a linear dissipative medium is proportional both to its temperature and to its power dissipation rate. A drastic reduction of the coil noise is thus expected as a result of drop of the coil resistance as the temperature is lowered. Conversely, neither the sample temperature nor its equivalent damping resistance into the coil are affected, so that the noise from the sample will eventually limit the improvement in sensitivity [7]. The earliest proofs of concept have been established with copper coils cooled to 4.2 K using liquid helium (LHe) and at 77 K with liquid nitrogen, respectively, to increase the sensitivity in 13C high-resolution spectroscopy [8], and in low-field clinical MRI [9]. A dramatic improvement in sensitivity was reported in NMR spectroscopy, whereas for MRI the gain was rather modest due to the large noise induced by the human body. A few years later, the potential of HTS ceramics for replacing copper had been demonstrated with prototypes of RF coils, at 77 K for low-field MRI [10], 10 K for MR microscopy at high field [4] and 25 K for 1H detection in a 400 MHz spectrometer [11]. Electrical losses of more than two orders of magnitude below that of copper were reported. Since these early achievements, however, the cooled-copper RF coil technology has spread widely in the markets of high-resolution NMR spectroscopy [12] and preclinical micro-MRI [13]. The replacement of copper by HTS materials, on the other hand, has remained a research topic in industrial and academic labs [3,14–18]. The objective of this section is to provide guidelines to assess the coil material performance in terms of ultimate SNR performance and report the limits achievable with copper.
Using the reciprocity principle [7], the contribution of the RF coil to the final signal-to-noise ratio in MR images can be expressed by the RF sensitivity factor
that represents the time-domain SNR available from the coil per unit magnetic moment and unit acquisition bandwidth [19] at the Larmor frequency
where
In the case of lossless samples (i.e. when
with
Figure 1 displays the ultimate sensitivity gain for RF coils at different cooling temperatures as compared to copper at room-temperature. For normal metals, the resistivity
with
becomes comparable or greater than 1, as showcased in Figure 1A (grey curves). The factor
FIGURE 1. (A) Expected resistivity
Figure 1B also displays the normalized RF sensitivity factor expected for a typical HTS material at 77 K. A first key feature of HTS is the partition of the conduction electrons between a coherent Cooper pair state that carries the lossless supercurrent, and a dissipative contribution from excited quasiparticle states. A second is that the electromagnetic field is extinguished over a surface layer, the thickness of which corresponds to the frequency-independent London penetration depth
The depth
Another consideration with cryogenic probes comes from the thermal insulation wall inserted between the sample and the low temperature RF coil. To preserve the field-of-view accessible with the coil, the thickness of the insulator needs to be accommodated with a larger coil radius. Consequently, looking back at Equation 1, this tends to mitigate the sensitivity gains predicted above by reducing the
Progress in Magnetic Resonance Imaging Compatible Cryostats
The need for cryogenic means has long been seen as a serious bottleneck to the dissemination of cooled RF coil technology in MR applications. Many advances have been made during the last 20 years in the emerging field of superconductive electronics where the requirement for low-cost, compact, secure, and easy-to-use cryogenic devices has been a strong motivation [23]. However, MR detection meets severe additional constraints of compatibility with the presence of RF as well as of large static magnetic fields: no metallic parts can be present near the RF coil and no magnetic materials nor moving metallic parts can be in proximity to the magnet bore. Specific materials such as PVC, fiberglass composites, polystyrene or polyurethane foams, quartz, glass, or ceramics are generally involved in the cryostat body. Thus, initial developments [19] mostly relied on homemade cryogenic designs. First trials involved immersion of the coil circuit in LHe (4.2 K) or LN2 (77 K) [8,9], allowing intrinsically uniform and temperature-stabilized cooling of volume coils or large surface coils. To reduce mechanical and electrical instabilities due to boil-off and bubbling, the trend was then to replace the direct liquid immersion by a flow of cold nitrogen [24] or helium [25] gas, or by a cold finger made of high-performance dielectric material such as sapphire [11,26].
Another particular constraint that arises during the study of human subjects, animal models and biological samples is to maintain the living system at room temperature all the while keeping a small separation from the cooled RF coil. This can be provided either by a plain insulating wall [10], by a double-wall vacuum layer [25], or by a single-wall evacuated cryostat containing the RF coil [27]. The latter allowed the reduction of the insulator thickness to less than 1 mm, which is particularly suitable for high-resolution surface imaging such as that of the human skin and rodents in a preclinical setting. Nevertheless, using conventional superinsulation with metallized layers to block thermal radiation through the insulating window does not fit the requirement for RF transparency, so an amount of thermal leakage still occurs. Thus, insufflation of a warm stream of an inert gas such as N2 into a reserved space along the outer cryostat surface has been applied to preserve the temperature stability of small living samples [24]. As pointed out in the previous section, the requirement for a minimized insulation thickness is not as severe with HTS coils as with copper coils since the formers offer much higher ultimate SNR gains.
More than 15 years after the early demonstrations of RF coils cooled by direct fluid immersion in the academic world, more elaborated closed-cycle cooled RF probes [28] have become available as standard accessories for analytical NMR spectroscopy [29]. In the new probes, typically, circulating helium gas transfers the cooling power from a distant two-stage Gifford–McMahon refrigerator to the RF coil, through heat exchangers at a regulated temperature ranging between 20 and 30 K. Temperatures of 20–30 K are low enough to improve the sensitivity of copper RF probes by a ratio of four when compared to room temperature, which enables 16 times faster acquisitions at constant SNR. Moreover, an integrated preamplifier cooled to below 80 K by the first stage of the Gifford-McMahon machine ensures a low noise contribution from the receiver, even with noiseless samples. Such new technology has been expanding rapidly since the late 1990s to face the high throughput demand in analytical chemistry. Its impact is tremendous in large scientific domains such as genomics, proteomics, and drug discovery, where the structural and dynamic analysis of large proteins is very time consuming, only small amounts of the substance under study are available, and less sensitive nuclei such as 13C and 15N must be investigated [12].
Based on the closed-cycle technology already developed for NMR probes, cryocooled surface probes designed for high-resolution imaging of the mouse brain [30–32] have become commercially available for routine operation on micro-MRI systems up to 15.2 T. This has given rise to growing numbers of cutting-edge structural and functional imaging applications [13]. Typically, the probes are made of 15–20 mm circular or rectangular copper coil elements bent onto a cold finger. Their half-cylindrical shell packaging leaves a large access to the animal head and offers a warm contact with a temperature-controlled ceramic wall, ensuring a minimum coil-sample distance of about 1 mm. Compared to a similar probe operating at room temperature, in vivo image acquisition with a 30 K quadrature surface probe at 9.4 T exhibited a SNR gain of about 2.5 when averaged over the mouse brain [31].
The rather high temperature range targeted with the new closed-cycle technology allows a much more robust operation than achievable with the early cryostats operating at 10 K and below. Notably, the new technology offers a fully automated control and does not require to refill the system with cryogenic fluids. However, designs until now were not versatile enough to allow for a change of RF coil configuration or orientation. The set-up is expensive and cumbersome because a high cooling power is required from the Gifford-McMahon unit, the rotary valve of which has to be located a few meters away from the MRI magnet for electromagnetic compatibility issues. Moreover, while cooling to 30 K is needed to take the most out of copper RF coils, HTS materials already exhibit sufficient performances at temperatures above 50 K. Thus, research activity is still ongoing in academic labs with simple all-in-bore 77 K static cryostats offering some flexibility and sufficient autonomy for acquisition times of several hours [3,33–36].
In order to optimize the handling of HTS coils, cryogenic systems providing intermediate temperature control above 50 K [15,37–40], and variable temperature set-ups for frequency tuning [39,41] have been proposed. Along with a higher working temperature, an important trend to alleviate the charge on closed-cycle cryocoolers is to drastically reduce the distance between the refrigerator machine and the cooled RF coil. Pulsed tube refrigerators without moving parts close to the cold finger have been used at close proximity to the MRI magnet, i.e. on the patient bed [42], or even inside the magnet within the RF coil cryostat [37]. The latter solution was designed to allow a variable orientation of the RF coil with a very compact design; however, it actually needs to be operated in a downward orientation (i.e. upright for the pulse tube) because of gravity issues for ensuring an efficient flow cooling.
The use of a refrigerant was eliminated in the cryogen-free system proposed by [40], later described by [42], and displayed in Figure 2. The cold source, provided by a commercial single-stage pulse tube cryocooler connected to a remote motor, is located directly on the non-magnetic cryostat vessel. The cold head is thermally coupled to the HTS coil by a series of high-conductivity pure aluminum flexible straps used in combination with copper and sapphire supports, as depicted in Figure 2B. This solution allows the transfer of sufficient cooling power to the RF coil; offers great adaptability to the RF coil bedding geometry; minimizes static and RF magnetic field perturbations induced by the system; and minimizes the transfer of mechanical vibrations to the HTS coil. Moreover, a non-magnetic heat exchanger tip anchored to the coil bed and a fully automated control system enable the control of the coil temperature to within 0.1 K, [41,42]. Thus, the coil frequency can be tuned through temperature as a parameter. These aforementioned characteristics, as well as the reasonable weight (70 kg) and compact nature of the cryostat, allow an installation as close as possible to the measurement zone of the MRI scanner. In fact, this system is cryogenist-free—i.e. no cryogenic knowledge is required to use it -, autonomous, and adaptable to any MRI structure that stays in the examination room during the experiments.
FIGURE 2. Images of the cryogen-free cryostat. (A) The cryostat and associated equipment. (B) Detailed 3D drawing of the general layout and cryogenic scheme of the system. (C) The HTS coil signal is probed by inductive coupling via the coupling coil. Matching and tuning of the HTS coil to the acquisition chain of the MRI scanner is performed with the coupling coil and the tuning coil, respectively. Fine tuning is also achieved via temperature control of the HTS coil. (D) Position of the HTS coil inside the cryostat, approximately 1.7 mm under the imaging zone. (E) Cryostat on the patient bed which can be inserted in the bore of the 1.5 T MRI scanner. This system was fully described in [42].
Mitigating the Effects of High-Temperature Superconducting Coil Diamagnetism
Another recent advance was made by better understanding the interaction between the HTS coil and the static magnetic field
The static field inhomogeneity artifacts due to static field expulsion from HTS coil were recently investigated [5] for different cooling conditions and temperatures of the HTS coil. This was done by measuring the MR signal at 1.5 T of a water sample in the vicinity of a HTS coil: a 12 mm multi-turn Transmission Line Resonator [44] (MTLR) with a pair of six-turn YBCO spirals (Tc = 86 K) on both sides of a Al2O3 substrate. In the aforementioned study, the HTS coil was non-resonant to avoid RF
In Figure 3, we present the main results of this study. Images acquired after zero-field cooling displayed important signal losses and phase perturbations (Figure 3A). The observed artifacts were predominant closer to the HTS coil and shared the same characteristic dimensions. In comparison, images acquired after field-cooling the coil in the MR scanner were unaltered. The phase perturbation was also shown to have a strong temperature dependence (Figure 3B); the amplitude of the perturbation was continuously reduced as the temperature increased. Using a phenomenological model of the coil superconducting diamagnetism, it was shown that these artifacts were caused by a small yet unavoidable misalignment, less than 1°, of the HTS coil in the static field
FIGURE 3. MR images of a water sample in the vicinity of a HTS coil (A) after different cooling conditions of the coil (at 58 K), and (B) for different temperatures of the coil after zero-field cooling. The slice shown is parallel to the HTS coil plane, ∼4.2 mm above.
Two important conclusions can be drawn from these results. 1) The
Decoupling the High-Temperature Superconducting Coil During Transmission
Another possible source of artifacts comes from the magnetic coupling between the HTS surface coil and the volume coil used during transmission. Copper coil technology typically relies on diodes to deactivate the reception coil during transmission. Unfortunately, these standard techniques cannot be implemented with HTS materials, as the physical connections required for incorporating discrete components in the coil circuit degrade its quality factor and add too much noise. Decoupling is hence a major technological challenge for HTS coils. A solution can be found through the highly nonlinear electrical properties presented by HTS materials [46]. The quality factor
Setup 1—The HTS surface coil is used in a dedicated cryostat [50] shown schematically in Figure 4, cooled in Earth’s magnetic field and characterized with a VNA and a method inspired from [47]. The electrical parameters of the HTS coil are obtained through the measurement of its RF response using an inductive coupling approach involving a measurement probe overcoupled with the HTS coil. The measurement probe is a 10 mm diameter copper loop placed in front of the HTS coil at a 5 mm distance. The incident power produced by the VNA and transmitted to the HTS coil via the measurement probe is then converted in equivalent
FIGURE 4. Schematic drawing of the interior of the dedicated cryostat used in Setup 1.The VNA is connected to a measurement probe inductively coupled to the HTS coil, which is secured on a cold finger. The temperature is monitored using a thermosensor (Pt100). The system was fully described by [50].
First, this setup simulates an MRI sequence. A high amplitude pulse with
Secondly, the setup performs a swept-frequency analysis where the modulus and phase of the compensated reflection coefficient
Setup 2—Then, the HTS coil is cooled inside the MR compatible cryostat [42] integrated into a 1.5 T clinical MRI (Achieva, Philips) shown in Figure 2. Fine frequency tuning of the HTS coil at
FIGURE 5. (Setup 2)–(A) Pictures of the HTS coil, water sample and copper coil. MR acquisitions obtained with (A) the HTS coil at 60 K and (B) the copper coil at room temperature for different
With this second setup, we first connected the coupling coil to the MRI acquisition RF channel to acquire the HTS coil signal during reception. A reference measurement was obtained using a room temperature copper coil of similar dimensions (Figure 5A). Then, the imaging sample was removed and the coupling coil signal
Figure 6 presents the results obtained with Setup 1. Figure 6A displays the dependance between the
FIGURE 6. (Setup 1)–(A) Measured quality factor
The effects of the RF power on the HTS material are shown in Figures 6B–G, which displays the modulus and phase of
In Figure 5, we present the results acquired with Setup 2. The spatial inhomogeneities observed in the MR images of Figures 5B–C are due to a non-uniform flip angle of the spins over the sample. This is produced during transmission by the re-emission of the spatially non-uniform
These two setups studied the nonlinear electrical properties of the HTS coil in Earth’s magnetic field (Setup 1) and in a field of 1.5 T (Setup 2). We observed differences between the behavior of
Discussion
In this work, we reviewed some of the challenges and recent advances made in the implementation of HTS coil in MRI. The key feature of HTS coils is that they do not share the lower limit on resistivity experienced by copper coils, although the
In order for these applications to be developed, the HTS coil must be integrated into a cryostat that achieves both MRI compatibility and good thermal insulation between the cryogenic probe and the room temperature imaging subject. Although this has proven to be an important technological challenge, recent years have seen the design of new and more robust cryostats, with the notable development of the cryogen-free and autonomous system that we described [42]. While we focused our attention primarily on closed-cycle technology, onboard open-cycle cooling approaches dedicated to small RF coils show promising perspectives to address the challenge of cooling high-density flexible HTS coil arrays, i.e. arrays where the relative positions of the coil elements can be modified. Such individual onboard cooling units or, alternatively, distributed refrigerant streams, could be implemented respectively with a compact N2 Joule-Thomson refrigerator [55] or a microfluidic LN2 exchanger [56]. The actual challenges of such a device are improving the compactness and reducing the cooling time while keeping a standalone, user-friendly, and cryogen-free solution. Compactness will be achieved by reducing the dimensions, and therefore the weight of the device. Consequently, this will reduce the heat stored in the system at room temperature as well as the thermal resistance between the cooling system and the coil, allowing for faster cooling or less cooling power. This cannot be accomplished without a fine optimization of the thermal capacity of the thermal link with the cooling power of the cryocooler at all stages of cooling. These improvements will bring the device even closer to the MRI measurement zone and impose having even more non-magnetic material entering its composition. Using new polymeric materials—robust, lighter, and compatible with vacuum and MRI scanners–, will fulfill both requirements.
The question of the optimal working temperature for HTS coils remains open. On the one hand, we have seen that working at temperatures significantly lower than
The issue of tuning and decoupling of the HTS coil also comes into play for the selection of its working temperature. Until now, our studies were conducted using HTS coils working at around 80 K in transceiver mode. As stated earlier, the HTS coil design is chosen so as to resonate close to the Larmor frequency, and a fine tuning of its resonance frequency is achieved by positioning a closed-circuit copper loop [3]. Thanks to the new cryostat, this fine tuning can now be performed by a fine control of the temperature, as proposed by [42], since the coil resonance frequency depends on the temperature [39]. Also, matching the HTS coil to the source impedance was achieved by a contactless matching technique using an inductive pick-up loop tuned at the Larmor frequency. The coupling coefficient is manually adjusted, modifying the distance between the pick-up loop and the HTS coil, to reach the optimal 50
In parallel, until the issue of the magnetic coupling between the HTS coil and the transmission volume coil is solved, the former can still be used in transceiver mode (i.e. as the transmitter and receiver) Typically, the transmit level of the coupled volume coil is manually adjusted to account for the presence of the HTS coil, which effectively becomes the source of the RF excitation flipping the sample magnetization during transmission. From the perspective of the volume coil, this leads to flip angle values of one to two orders of magnitude smaller than the ones currently used in MR protocols in order to achieve a real π/2-flip angle in the region of interest (ROI) These precautions are necessary to limit the
To that end, the approach illustrated in our experimental results exploits the non-linear electrical response of HTS materials to minimize the quality factor during transmission. Although the coil under study did not reach resistance levels sufficient to achieve proper decoupling, it did provide a proof of concept for this technique. To go further, the HTS coil have to reach its dissipative state for lower RF
Conclusion
HTS coil for MRI has come a long way in the last 30 years. Cryogenics have advanced sufficiently to allow the design of MRI compatible cryostats that allow one to reliably operate HTS coils. In addition, we now better understand how HTS coils interact with the superposed magnetic fields
Data Availability Statement
The raw data supporting the conclusion of this article will be made available by the authors, without undue reservation.
Author Contributions
AL: coordinated the writing of the paper; was involved in the redaction of many sections of the article; and acquired parts of the data shown. GA and BB: designed the MRI compatible cryostat shown in the paper; and both were involved in the writing of the section “Progress in MRI compatible cryostats” which displays their work. JB and CvdB: helped design the experiment—Setup 2—shown in the paper; were involved in the writing of the section “Decoupling the HTS coil during transmission;” provided expertise during the discussion of the experimental results LD: wrote the state-of-the-art sections for “Limits of copper coils in MRI” and “Progress in MRI compatible cryostats” MPQ: coordinated the research project; involved in the design of the experiments shown in this work; main writer of the section “Decoupling the HTS coil during transmission.” All authors actively contributed to the writing of this paper.
Funding
Experiments were performed on the 1.5 T MRI platform of CEA/SHFJ, affiliated to the France Life Imaging network and partially funded by the network (Grant ANR-11-INBS-0006), under Grant No. ANR-14-CE17-0003 (SupraSense project) and LabeX LaSIPS (ANR-10-LABX-0040-LaSIPS).
Conflict of Interest
The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.
Acknowledgments
The authors thank Dr. Jean-Christophe Ginefri for his contribution to the development of HTS coils within our research group. The authors also thank Rose-Marie Dubuisson for her help with the 1.5 T MRI scanner platform, and Georges Willoquet for his work on the electronics implemented in both cryostats used in this work, as well as Isabelle Saniour and Michel Geahel for their past contributions to this project.
References
1. Bednorz JG, Müller KA (1986) Possible highT C Superconductivity in the Ba-La-Cu-O System. Z für B - Condens Matter 64, 189–93. doi:10.1007/BF01303701
2. Lancaster MJ (1997, Passive Microwave Device Applications of High-Temperature Superconductors). Passive Microwave Device Applications of High-Temperature Superconductors. United Kingdom: Cambridge University Press doi:10.1017/CBO9780511526688
3. Poirier-Quinot M, Ginefri J-C, Girard O, Robert P, Darrasse L (2008) Performance of a Miniature High-Temperature Superconducting (HTS) Surface Coil for In Vivo Microimaging of the Mouse in a Standard 1.5T Clinical Whole-Body Scanner. Magn Reson Med 60, 917–27. doi:10.1002/mrm.21605
4. Black RD, Roemer PB, Mogro‐Campero A, Turner LG, Rohling KW (1993) High Temperature Superconducting Resonator for Use in Nuclear Magnetic Resonance Microscopy. Appl Phys Lett 62, 771–3. doi:10.1063/1.108574
5. Labbé A, Dubuisson R-M, Ginefri J-C, Van Der Beek CJ, Darrasse L, Poirier-Quinot M (2020) Static Field Homogeneity Artifacts Due to Magnetic Flux Expulsion by HTS Coils for High-Resolution Magnetic Resonance Imaging. Appl Phys Lett 117, 254101. doi:10.1063/5.0033894
6. Hoult DI, Richards RE (1976) The Signal-To-Noise Ratio of the Nuclear Magnetic Resonance experiment. J Magn Reson (1969) 24, 71–85. doi:10.1016/0022-2364(76)90233-X
7. Hoult DI, Lauterbur PC (1979) The Sensitivity of the Zeugmatographic experiment Involving Human Samples. J Magn Reson (1969) 34, 425–33. doi:10.1016/0022-2364(79)90019-2
8. Styles P, Soffe NF, Scott CA, Crag DA, Row F, White DJ, et al. (1984) A High-Resolution NMR Probe in Which the Coil and Preamplifier Are Cooled with Liquid Helium. J Magn Reson (1969) 60, 397–404. doi:10.1016/0022-2364(84)90050-7
9. Hall AS, Barnard B, McArthur P, Gilderdale DJ, Young IR, Bydder GM (1988) Investigation of a Whole-Body Receiver Coil Operating at Liquid Nitrogen Temperatures. Magn Reson Med 7, 230–5. doi:10.1002/mrm.1910070211
10. Hall AS, Alford NM, Button TW, Gilderdale DJ, Gehring KA, Young IR (1991) Use of High Temperature Superconductor in a Receiver Coil for Magnetic Resonance Imaging. Magn Reson Med 20, 340–3. doi:10.1002/mrm.1910200218
11. Anderson WA, Brey WW, Brooke AL, Cole B, Delin KA, Fuks LF, et al. (1995) High-sensitivity NMR Spectroscopy Probe Using Superconductive Coils. Bull Magn Reson 17, 98–102.
12. Kovacs H, Moskau D, Spraul M (2005) Cryogenically Cooled Probes-A Leap in NMR Technology. Prog Nucl Magn Reson Spectrosc 46, 131–55. doi:10.1016/j.pnmrs.2005.03.001
13. Niendorf T, Pohlmann A, Reimann HM, Waiczies H, Peper E, Huelnhagen T, et al. (2015) Advancing Cardiovascular, Neurovascular and Renal Magnetic Resonance Imaging in Small Rodents Using Cryogenic Radiofrequency Coil Technology. Front Pharmacol 6, 255. doi:10.3389/fphar.2015.00255
14. Wosik J, Xue L, Xie L-M, Kamel MR, Nesteruk K, Bankson JA (2007) Superconducting Array for High-Field Magnetic Resonance Imaging. Appl Phys Lett 91, 183503. doi:10.1063/1.2801384
15. Nouls JC, Izenson MG, Greeley HP, Johnson GA (2008) Design of a Superconducting Volume Coil for Magnetic Resonance Microscopy of the Mouse Brain. J Magn Reson 191, 231–8. doi:10.1016/j.jmr.2007.12.018
16. Lin I-T, Yang H-C, Chen J-H (2012) Enlargement of the Field of View and Maintenance of a High Signal-To-Noise Ratio Using a Two-Element High-T C Superconducting Array in a 3T MRI. PLoS One 7, e42509. doi:10.1371/journal.pone.0042509
17. Ramaswamy V, Hooker JW, Withers RS, Nast RE, Brey WW, Edison AS (2013) Development of a 13C-Optimized 1.5-mm High Temperature Superconducting NMR Probe. J Magn Reson 235, 58–65. doi:10.1016/j.jmr.2013.07.012
18. Laistler E, Poirier-Quinot M, Lambert SA, Dubuisson R-M, Girard OM, Moser E, et al. (2015) In Vivo MR Imaging of the Human Skin at Subnanoliter Resolution Using a Superconducting Surface Coil at 1.5 Tesla. J Magn Reson Imaging 41, 496–504. doi:10.1002/jmri.24549
19. Darrasse L, Ginefri JC (2003) Perspectives with Cryogenic RF Probes in Biomedical MRI. Biochimie 85, 915–37. doi:10.1016/j.biochi.2003.09.016
20. Matula RA (1979) Electrical Resistivity of Copper, Gold, Palladium, and Silver. J Phys Chem Reference Data, 8(4), 1147–298. doi:10.1063/1.555614
21. Sondheimer EH (1952) The Mean Free Path of Electrons in Metals. Adv Phys 1, 1–42. doi:10.1080/00018735200101151
22. Wang Y, Su HT, Huang F, Lancaster MJ (2007) Measurement of YBCO Thin Film Surface Resistance Using Coplanar Line Resonator Techniques from 20 MHz to 20 GHz. IEEE Trans Appl Supercond 17, 3632–9.
23. Anders S, Blamire MG, Buchholz F-I, Crété D-G, Cristiano R, Febvre P, et al. (2010) European Roadmap on Superconductive Electronics - Status and Perspectives. Physica C: Superconductivity 470, 2079–126. doi:10.1016/j.physc.2010.07.005
24. McFarland EW, Mortara A (1992) Three-dimensional NMR Microscopy: Improving SNR with Temperature and Microcoils. Magn Reson Imaging 10, 279–88. doi:10.1016/0730-725X(92)90487-K
25. Black R, Early T, Roemer P, Mueller O, Mogro-Campero A, Turner L, et al. (1993) A High-Temperature Superconducting Receiver for Nuclear Magnetic Resonance Microscopy. Science. 259, 793–5. doi:10.1126/science.8430331
26. Miller JR, Zhang K, Ma QY, Mun IK, Jung KJ, Katz J, et al. (1996) Superconducting Receiver Coils for Sodium Magnetic Resonance Imaging. IEEE Trans Biomed Eng 43, 1197–9. doi:10.1109/10.544344
27. Ginefri J-C, Darrasse L, Crozat P (2001) High-temperature Superconducting Surface Coil for In Vivo Microimaging of the Human Skin. Magn Reson Med 45, 376–82. doi:10.1002/1522-2594(200103)45:3<376::aid-mrm1049>3.0.co;2-r
29. Hill HDW (1997) Improved Sensitivity of NMR Spectroscopy Probes by Use of High-Temperature Superconductive Detection Coils. IEEE Trans Appl Supercond 7, 3750–5. doi:10.1109/77.622233
30. Ratering D, Baltes C, Nordmeyer-Massner J, Marek D, Rudin M (2008) Performance of a 200-MHz Cryogenic RF Probe Designed for MRI and MRS of the Murine Brain. Magn Reson Med 59, 1440–7. doi:10.1002/mrm.21629
31. Baltes C, Radzwill N, Bosshard S, Marek D, Rudin M (2009) Micro MRI of the Mouse Brain Using a Novel 400 MHz Cryogenic Quadrature RF Probe. NMR Biomed. 22, 834–42. doi:10.1002/nbm.1396
33. Ginefri J-C, Poirier-Quinot M, Girard O, Darrasse L (2007) Technical Aspects: Development, Manufacture and Installation of a Cryo-Cooled HTS Coil System for High-Resolution Iin-Vvivo Imaging of the Mouse at 1.5T. Methods 43, 54–67. doi:10.1016/j.ymeth.2007.03.011
34. Hu B, Varma G, Randell C, Keevil SF, Schaeffter T, Glover P (2012) A Novel Receive-Only Liquid Nitrogen ($\Hbox{LN}_{2}$ )-Cooled RF Coil for High-Resolution In Vivo Imaging on a 3-Tesla Whole-Body Scanner. IEEE Trans Instrum Meas 61, 129–39. doi:10.1109/TIM.2011.2157575
35. Lin I-T, Yang H-C, Chen J-H (2013) A Temperature-Stable Cryo-System for High-Temperature Superconducting MR Iin-Vvivo Imaging. PLoS One 8, e61958. doi:10.1371/journal.pone.0061958
36. Sánchez-Heredia JD, Baron R, Hansen ESS, Laustsen C, Zhurbenko V, Ardenkjaer-Larsen JH (2020) Autonomous Cryogenic RF Receive Coil for 13 C Imaging of Rodents at 3 T. Magn Reson Med 84, 497–508. doi:10.1002/mrm.28113
37. Vester M, Steinmeyer F, Roas B, Thummes G, Klundt K (1997) High Temperature Superconducting Surface Coils with Liquid Nitrogen or Pulse Tube Refrigeration (Abstr). in Vancouver, BC: ISMRM 5th Scientific Meeting and Exhibition, doi:10.4095/219009
38. Darrasse L, Chiron L, Ginefri J-C, Poirier-Quinot M, Trollier T, Ravex A, et al. (2004) Cryosonde IRM : antenne IRM supraconductrice pour la microscopie des régions superficielles du corps humain et des petits modèles animaux. ITBM-RBM 25, 254–9. doi:10.1016/j.rbmret.2004.09.003
39. Lambert S, Ginefri J-C, Poirier-Quinot M, Darrasse L (2013) High-temperature Superconducting Radiofrequency Probe for Magnetic Resonance Imaging Applications Operated below Ambient Pressure in a Simple Liquid-Nitrogen Cryostat. Rev Scientific Instr 84, 054701. doi:10.1063/1.4802947
40. Authelet G, Poirier-Quinot M, Ginefri J-C, Bonelli A, Baudouy B (2017) Conceptual Design of a Cryogen-free μMRI Device. in IOP Conf Ser Mater Sci Eng, 12122.
41. Authelet G, Bonelli A, Poirier-Quinot M, Ginefri JC, Saniour I, Baudouy B (2019) All Polymer Cryogen Free Cryostat for μ-MRI Application at Clinical Field. in IOP Conf Ser Mater Sci Eng, 12156.
42. Saniour I, Authelet G, Baudouy B, Dubuisson R-M, Jourdain L, Willoquet G, et al. (2020) A Temperature-Controlled Cryogen Free Cryostat Integrated with Transceiver-Mode Superconducting Coil for High-Resolution Magnetic Resonance Imaging. Rev Scientific Instr 91, 055106, doi:10.1063/1.5143107
43. Tinkham M (1996) Introduction to Superconductivity. 2nd ed. eds. J Shira, and E Castellano New-York: McGraw-Hill.
44. Serfaty S, Haziza N, Darrasse L, Kan S (1997) Multi-turn Split-Conductor Transmission-Line Resonators. Magn Reson Med 38, 687–9. doi:10.1002/mrm.1910380424
45. Díez-Sierra J, López-Domínguez P, Rijckaert H, Rikel M, Hänisch J, Khan MZ, et al. (2020) High Critical Current Density and Enhanced Pinning in Superconducting Films of YBa2Cu3O7−δ Nanocomposites with Embedded BaZrO3, BaHfO3, BaTiO3, and SrZrO3 Nanocrystals. ACS Appl Nano Mater 3, 5542–53. doi:10.1021/acsanm.0c00814
46. Shen ZY (1994) High-temperature Superconducting Microwave Circuits. Boston, MA: Artech House Publishers.
47. Girard O, Ginefri J-C, Poirier-Quinot M, Darrasse L (2007) Method for Nonlinear Characterization of Radio Frequency Coils Made of High Temperature Superconducting Material in View of Magnetic Resonance Imaging Applications. Rev Scientific Instr 78, 124703. doi:10.1063/1.2825241
48. Geahel M (2018) Découplage de détecteurs radiofréquences supraconducteurs à très haute sensibilité pour la micro-imagerie par résonance magnétique. https://tel.archives-ouvertes.fr/tel-01880203/document (Accessed March 20, 2021).
49. Saniour I, Authelet G, Baudouy B, Dubuisson RM, van der Beek CJ, Darrasse L, et al. (2019) Novel Passive Decoupling Approach for High Resolution HTS RF Coils Based on the Nonlinear Electrical Properties of Superconductors. Rotterdam: ESMRMB 2019. in 36th Annual Scientific Meeting. Springer.
50. Saniour I, Geahel M, Briatico J, Beek CJvan der, Willoquet G, Jourdain L, et al. (2020) Versatile Cryogen-free Cryostat for the Electromagnetic Characterization of Superconducting Radiofrequency Coils. EPJ Tech Instrum 7, 1--11. doi:10.1140/epjti/s40485-020-00055-2
51. Clem JR, Sanchez A (1994) Hysteretic AC Losses and Susceptibility of Thin Superconducting Disks. Phys Rev B 50, 9355, 62. doi:10.1103/physrevb.50.9355
52. Frass-Kriegl R, Laistler E, Hosseinnezhadian S, Schmid AI, Moser E, Poirier-Quinot M, et al. (2016) Multi-turn Multi-gap Transmission Line Resonators - Concept, Design and First Implementation at 4.7 T and 7 T. J Magn Reson, 273, 65–72. doi:10.1016/j.jmr.2016.10.008
53. Ginefri J-C, Durand E, Darrasse L (1999) Quick Measurement of Nuclear Magnetic Resonance Coil Sensitivity with a Single-Loop Probe. Rev Scientific Instr 70, 4730–1. doi:10.1063/1.1150142
54. Poirier-Quinot M, Frasca G, Wilhelm C, Luciani N, Ginefri J-C, Darrasse L, et al. (2010) High-resolution 1.5-Tesla Magnetic Resonance Imaging for Tissue-Engineered Constructs: a Noninvasive Tool to Assess Three-Dimensional Scaffold Architecture and Cell Seeding. Tissue Eng C: Methods 16, 185–200. doi:10.1089/ten.tec.2009.0015
55. Wright AC, Song HK, Elliott DM, Wehrli FW (2005) Use of a Joule-Thomson Micro-refrigerator to Cool a Radio-Frequency Coil for Magnetic Resonance Microscopy. Rev Scientific Instr 76, 014301. doi:10.1063/1.1824340
56. Koo C, Godley RF, Park J, McDougall MP, Wright SM, Han A (2011) A Magnetic Resonance (MR) Microscopy System Using a Microfluidically Cryo-Cooled Planar Coil. Lab Chip 11, 2197–203. doi:10.1039/c1lc20056a
57. Li Z, Willoquet G, Guillot Gv., Hosseinnezhadian S, Jourdain L, Poirier-quinot M, et al. (2016) Study of Two Contact-Less Tuning Principles for Small Monolithic Radiofrequency MRI Coils and Development of an Automated System Based on Piezoelectric Motor. Sensors Actuators A: Phys. 241, 176–89. doi:10.1016/j.sna.2016.02.008
Keywords: high temperature superconductor, magnetic resonance imaging, MRI compatible cryostat, nonlinear properties of superconductors, HTS coil, cryogenics, imaging artefact
Citation: Labbé A, Authelet G, Baudouy B, van der Beek CJ, Briatico J, Darrasse L and Poirier-Quinot M (2021) Recent Advances and Challenges in the Development of Radiofrequency HTS Coil for MRI. Front. Phys. 9:705438. doi: 10.3389/fphy.2021.705438
Received: 05 May 2021; Accepted: 24 June 2021;
Published: 20 July 2021.
Edited by:
Simone Angela S. Winkler, Cornell University, United StatesReviewed by:
Michael D. Noseworthy, McMaster University, CanadaSigrun Roat, Medical University of Vienna, Austria
Elizaveta Motovilova, Weill Cornell Medicine, United States
Copyright © 2021 Labbé, Authelet, Baudouy, van der Beek, Briatico, Darrasse and Poirier-Quinot. This is an open-access article distributed under the terms of the Creative Commons Attribution License (CC BY). The use, distribution or reproduction in other forums is permitted, provided the original author(s) and the copyright owner(s) are credited and that the original publication in this journal is cited, in accordance with accepted academic practice. No use, distribution or reproduction is permitted which does not comply with these terms.
*Correspondence: Aimé Labbé, aime.labbe@universite-paris-saclay.fr; Marie Poirier-Quinot, marie.poirier-quinot@universite-paris-saclay.fr