This paper addresses the scientific and technological challenges related to coill development of wireless technoloyy frequency RF Mood enhancer natural remedies and techniques for magnetic resonance imaging MRI based on published literature together with technolgoy authors' interpretation and further considerations.
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Especially for flexible arrays, a reduction of the total amount, size, and weight of on-coil components is crucial. In this work, we focus on the realizability of completely wireless MR receive arrays by addressing and interrelating different aspects of the MR receive system.
The aim is to outline feasible and efficient approaches toward wireless communication in MRI and prospect digital wireless RF devices, highlighting the most promising strategies as well as associated benefits and challenges. Three subsystems that have to undergo significant changes for wireless MRI were identified: the MR receive signal chain, control signaling, and on-coil power supply.
Their functional blocks and respective possible physical location are depicted in Figure 1A. Different wireless transceiver positioning variants, estimated transmission distances, and angles are sketched in Figure 1B.
Figure 1. A Functional block diagram of a wireless MR receive chain consisting of three main subsystems: wireless MR signals bluecontrol signals greenand power supply orange.
B Side view of wireless transceiver positioning variants including transmission distance and angle estimations. Specific requirements that need to be met for each of the functional blocks are included, and benefits of current technology as well as current limitations or challenges encountered in their development are listed.
The following general requirements apply to all of the mentioned subsystems and corresponding components: MR compatibility no impact on MRI or component functioningpatient safety no heatinglinearity, low noise figure, low power consumption, low number of additional components, miniature component size, and minimum weight.
Figure 2. Summary of the state of the art in wireless radio frequency RF coil development. The MR signal is characterized by high signal frequency the Larmor frequencydepending on the investigated nucleus and B 0 field strength, typically in the order of 50— MHz.
Further, the DR easily reaches ~90 dB [ 28 ]. In extreme cases, especially for high-resolution 3D acquisitions at high B 0 fields, the DR can attain up to ~ dB [ 2930 ]. To enable proper signal conditioning for various imaging scenarios frequency, DR, number of receive coil elements, etc.
The choice of suitable digitization components is a critical task, as there is always a trade-off between achievable conversion rates, bit resolution, power dissipation, cost, and scalability to multi-channel systems.
In general, on-coil digitization is advantageous, as it improves signal and phase stability, yielding better image quality, and offers easier scalability to multi-channel systems [ 181931 ].
For component selection, the main challenges are related to the MR signal properties. To date, commercially available high-speed ADCs dedicated to MRI are limited to bit [ 3233 ], insufficient for some imaging scenarios with very high DR. Concerning the sampling rate, one possibility is direct sampling at the Nyquist rate, employing ADCs capable of sampling at high rates greater than twice the Larmor frequency [ 34 ].
However, the essential imaging information of the MR signal lies only within a small signal bandwidth maximum 1—2 MHzdetermined by the maximum gradient strength and the field of view FOVmodulated onto the carrier wave at the Larmor frequency.
Therefore, demodulation of the amplified analog RF signal to baseband around zero frequency or to an intermediate frequency IF by mixing with a local oscillator LO signal on-coil before conversion to digital data is possible.
This significantly lowers the ADC sampling rate requirement. This was shown with broadband on-coil receivers for optical fiber transmission of digital signals from two [ 21 ] or four [ 24 ] wrist coil channels at 1.
Direct undersampling corresponds to sampling at lower than twice the maximum frequency and digital demodulation at the same time. This technique was applied for single receive elements at 0.
Multi-channel scalable solutions in combination with optical fibers were proposed for in-field receivers with one ADC per coil element at 1. Recent research also demonstrated a digital RF front end adaptable for 16 channels and useable from 1.
Direct under sampling approaches are useful, as no analog conversion step is needed prior to digitization, and the amount of on-coil components is usually low.
Care has to be taken to remove signal ambiguities, e. Therefore, especially with discrete components, the form factor and power consumption of the receiver increase. Taken together, the required data rate for wireless transmission depends on the digitization approach and ADC bit resolution for any MR receive system with a specific B 0 field strength, imaging bandwidth, and number of coil elements.
Sequence parameters, such as the receive duty cycle the ratio between acquisition and repetition timealso influence the effective data rate.
Estimations of up to 2. Digital strategies for DR compression and coil-wise demodulation can be combined to efficiently reduce the data size to one-third of the original amount [ 41 ].
Nonetheless, with digital compression directly after digitization, the number of components and, therefore, also the power needed on-coil will increase [ 2242 ]. To give an estimate for the minimum data rate requirement, we take a modern clinical MRI setup at 3 T with 64 RF receive elements as a reference.
Our estimation is in line with other published values [ 4344 ], only differing in terms of assumed ADC bit resolution, receive duty cycle, or number of receive elements.
For digital wireless communication in MRI, apart from achievable data rate and power consumption, lossless spatial transmission is an important criterion. First MR data transfer tests based on the These approaches are clearly impractical for wireless MRI. More recent attempts were conducted with higher carriers in the 5 GHz band This Wi-Fi approach is interesting as small client routers, used in most portable devices nowadays, are available, providing sufficient spatial range for MRI.
Efficient data throughput could be improved up to Mbps, suitable for low-channel and low-bandwidth MRI. Recently, out-of-bore experiments with shielded WiGig dongles [ 51 ] have shown transmission rates of — Mbps over 3—5. This Wi-Fi standard meets our estimated minimum data rate requirement for a modern clinical MRI setup and is therefore viable for wireless coil arrays.
Also, the shorter spatial transmission range of one of the presented 60 GHz links [ 43 ] is sufficient for some transceiver positioning variants see Figure 1B. Optical wireless communication OWC [ 5253 ] with visible, infrared, or ultraviolet light carriers i.
An MR-compatible OWC front end has been tested for 2 m analog positron emission tomography detector signal transmission [ 56 ], but the technology has not yet been exploited for MR signals. Unlike Wi-Fi, high-speed OWC mostly requires a direct line of sight between transceivers, although some systems can even communicate via diffuse light reflections [ 57 ].
Suitable components for Li-Fi Light-Fidelity, i. Wireless digital MR signal transmission appears feasible with current Wi-Fi strategies under the condition that appropriate measures for data rate reduction prior to wireless transmission are implemented on-coil, e.
Wi-Fi protocol-dependent or component-related drawbacks, e. WiGig 60 GHz seems to be a promising strategy because of high data rate capability, sufficient transmission range, and low power consumption, although full functioning of WiGig hardware on-coil during an MR scan and the effect on image quality still have to be examined.
Also, the final interfacing of the chosen wireless WiGig transceiver to a digital RF coil still has to be demonstrated and can be challenging, as it requires the smooth interaction of various on-coil components. So far, Wi-Fi technology benefited from rapid development pushed by the portable device industry; therefore, we think that the implementation of future high-performance Wi-Fi transceivers in RF coils is an aspect to be followed up by the research community.
Alternatively, OWC strategies could be investigated for wireless MR signal transmission. With OWC, the wireless transmission of uncompressed, directly digitized MR signals could be envisioned, which is advantageous with respect to miniaturized device size and low system complexity but is questionable concerning a limited on-coil power budget.
Striving for full removal of coil cabling, a bidirectional wireless link is indispensable as signals must be sent not only from the coil to the MR scanner but also from the scanner control unit to the coil, mainly for triggering, synchronization, and in some cases, control of B 0 shimming.
Trigger signals need to be distributed to the coil electronics, e. Wireless detuning triggers transmitted via a MHz antenna during an MRI scan at 1. Presumably, these trigger signals could also be applied to activate power-consuming components preamplifiers, ADCs only during signal reception.
A stable clock, phase-synchronous with the MRI, controlling on-coil electronics such as ADC or down-conversionis critical. Clock jitter, which decreases the effective number of ADC bits and creates image artifacts, must be limited.
For synchronization of MR unit and in-bore receivers, one method is to physically transmit the MRI master clock to the receiver, which has been demonstrated with 1. This requires additional on-coil clocking electronics e.
In contrast, on-coil clock generators can be used but are particularly impaired by gradient induction; therefore, free-running oscillator information has to be sent to the MR system alongside sampled data to detect and correct for frequency and phase errors as well as time offsets, i.
This often requires both additional hardware and software in the wireless receive system [ 62 — 65 ]. Several MRI applications benefit from localized on-coil B 0 shimming with DC currents on the RF coil elements compensating for B 0 inhomogeneities [ 66 ].
High shim currents themselves cannot be wirelessly transmitted but can be wirelessly controlled, which has been successfully demonstrated by 2.
: MRI coil technology| An Easy Guide to MRI Coils Types | The additional conductor loss associated with coaxial loops appears to arise from a longer conductor pathway that includes both outer and smaller inner conductors. It has been instrumental and indispensible in jumpstarting MR Whole Body Imaging. With 35 mm inner diameter, this array coil is designed for mouse body examinations on multiple receiver instruments. While the 1. A , — Get closer with adaptive MRI coils. |
| MRI RF Coils | RF Coil Technical Details | Bruker | Technolpgy PubMed Google Scholar Schaller, B. Lu JY, Responsible alcohol use Tecnnology, Pauly J, Scott G. Increased fat metabolism, administered, and supervised the study technoogy acquired funding to support it. These coils tehnology available for BioSpec and PharmaScan 7 T, BioSpec 9. Höfflin J, Fischer E, Hennig J, Korvink JG. Will not scan High Noise or Grainy Images Artifact Homogeneity Non-uniform or shaded images Mechanical Problem Please specify below. Citation: Nohava L, Ginefri J-C, Willoquet G, Laistler E and Frass-Kriegl R Perspectives in Wireless Radio Frequency Coil Development for Magnetic Resonance Imaging. |
| Types of Volume Coils | Poor MRI coil technology Fat Saturation problem. Coill Card Please call with credit card information. Then Cholesterol management tips out to our team below! MAGNETOM Verio Proven 3T clinical imaging. This imaging strategy has matured and has a great potential to turn into a very significant examination in the near future. |
| Monthly Updates - Systems & Parts | The experiment was repeated three times and the mean value was reported. All experiments were performed on a GE Discovery MR 3 T system. In-vivo images were acquired using the same fast spin echo sequence. The SNR values were calculated according to Method 4 described in the NEMA Standards Publication MS R, R , as the ratio of the mean pixel value of signal within the specified ROI divided by the standard deviation of the noise calculated in the background region of the image, well removed from the phantom and any visible artifacts. The SNR maps were calculated by dividing the entire image by the standard deviation of the noise calculated as described above. The study was performed with ethical approval from the Weill Cornell Medicine Institutional Review Board and in accordance with all applicable regulations. Informed consent was obtained from the volunteer. Roemer, P. The NMR phased array. Article CAS Google Scholar. Vaughan, J. RF Coils for MRI Wiley, Google Scholar. Sodickson, D. Simultaneous acquisition of spatial harmonics SMASH : Fast imaging with radiofrequency coil arrays. Pruessmann, K. SENSE: Sensitivity encoding for fast MRI. Griswold, M. et al. Generalized autocalibrating partially parallel acquisitions GRAPPA. Article Google Scholar. Winkler, S. Evaluation of a flexible channel screen-printed pediatric MRI coil. Radiology 1 , — Fryar, C. Anthropometric reference data for children and adults: UnitedStates, — In: Statistics NCfH, ed. AIR Technology. Development and clinical implementation of very light weight and highly flexible AIR technology arrays. The ISMRM 25th Annual Meeting and Exhibition. Honolulu, HI, USA; Stickle, Y. A novel ultra-flexible high-resolution AIR Adaptive imaging receive channel bilateral phased array for 3T brachial plexus MRI. Saniour, I. Acoustically transparent and low-profile head coil for high precision magnetic resonance guided focused ultrasound at 3 T. Attenuation of the dark band artifact in MR-guided focused ultrasound using an ultra-flexible high-sensitivity head coil. Hardy, C. Imaging 28 5 , — Mager, D. An MRI receiver coil produced by inkjet printing directly on to a flexible substrate. IEEE Trans. Imaging 29 2 , — Jia, F. Knee MRI under varying flexion angles utilizing a flexible flat cable antenna. NMR Biomed. Corea, J. Screen-printed flexible MRI receive coils. Frass-Kriegl, R. Flexible channel coil array for high-resolution magnetic resonance imaging at 3 Tesla. PLoS One 13 11 , Hosseinnezhadian, S. A flexible channel transceiver array of transmission line resonators for 7 T MRI. Article ADS CAS Google Scholar. Zhang, B. A high-impedance detector-array glove for magnetic resonance imaging of the hand. Nohava, L. In Proceedings of the 27th Annual Meeting of ISMRM, Montreal, Canada. Ruytenberg, T. Shielded-coaxial-cable coils as receive and transceive array elements for 7T human MRI. Adriany, G. Nordmeyer-Massner, J. Mechanically adjustable coil array for wrist MRI. Lopez Rios, N. Size-adaptable channel receive array for brain MRI in human neonates at 3 T. Czerny, R. Malko, J. A flexible mercury-filled surface coil for MR imaging. CAS PubMed Google Scholar. Chitambar, C. Medical applications and toxicities of gallium compounds. Public Health 7 5 , — Rumble, J. CRC Handbook of Chemistry and Physics. Barta, R. How thin can you go? Performance of thin copper and aluminum RF coil conductors. Edelstein, W. The intrinsic signal-to-noise ratio in NMR imaging. Zhu, S. Ultrastretchable fibers with metallic conductivity using a liquid metal alloy core. Mehmann, A. On the bending and stretching of liquid metal receive coils for magnetic resonance imaging. Varga, M. Adsorbed eutectic GaIn structures on a neoprene foam for stretchable MRI coils. Automatic resonance frequency retuning of stretchable liquid metal receive coil for magnetic resonance imaging. Imaging 38 6 , — Port, A. Detector clothes for MRi: A wearable array receiver based on liquid metal in elastic tubes. Byron, K. An MRI compatible RF MEMs controlled wireless power transfer system. Theory Tech. Article ADS Google Scholar. Conductive elastomer for wearable RF coils. ISMRM Vincent, J. Conductive thread-based stretchable and flexible radiofrequency coils for magnetic resonance imaging. Hayes, C. Noise performance of surface coils for magnetic resonance imaging at 1. Xie, J. Cholleti, E. Highly stretchable capacitive sensor with printed carbon black electrodes on barium titanate elastomer composite. Sensors 19 1 , 42 Paul, C. Inductance: Loop and Partial Wiley, Fassler, A. Soft-matter capacitors and inductors for hyperelastic strain sensing and stretchable electronics. Smart Mater. Igreja, R. Analytical evaluation of the interdigital electrodes capacitance for a multi-layered structure. Actuators A 2—3 , — COMSOL Multiphysics. Ecoflex by Smooth-On. Sigma-Aldrich, Gallium-Indium eutectic. Duan, Q. Characterization of a dielectric phantom for high-field magnetic resonance imaging applications. Dielectric phantom recipe generator. Accessed 03 Mar 21 Download references. This work was supported by the National Institutes of Health under NIH R00EB, and GE Healthcare. thanks Terry Ching of Singapore University of Technology and Design for helpful discussions on material selection and preparation. and S. would like to acknowledge Muc Chu, Jojo Borja, and Jonathan Dyke of the Citigroup Biomedical Imaging Center for the helpful technical discussions. Authors thank Cynthia Fox for her valuable assistance in the proofreading of the manuscript. Department of Radiology, Weill Cornell Medicine, New York, NY, , USA. Department of Radiology, Hospital for Special Surgery, New York, NY, , USA. Elizaveta Motovilova, Ek Tsoon Tan, Hollis G. Victor Taracila, Jana M. You can also search for this author in PubMed Google Scholar. conceptualized work, performed numerical simulations, built coils, designed experiments, collected data, analyzed data, prepared figures, and wrote the manuscript. conceptualized work, performed imaging experiments and supported data analysis, V. conceptualized work and supported data analysis, J. built coils, T. supported data collection, J. built coils, H. conceptualized work, F. conceptualized work and supported data analysis, D. conceptualized work, supported data analysis, and provided supervision, S. conceptualized work, designed experiments, analyzed data, provided supervision and wrote the manuscript. All authors revised the manuscript. Correspondence to Elizaveta Motovilova or Simone A. Springer Nature remains neutral with regard to jurisdictional claims in published maps and institutional affiliations. Open Access This article is licensed under a Creative Commons Attribution 4. The images or other third party material in this article are included in the article's Creative Commons licence, unless indicated otherwise in a credit line to the material. If material is not included in the article's Creative Commons licence and your intended use is not permitted by statutory regulation or exceeds the permitted use, you will need to obtain permission directly from the copyright holder. Reprints and permissions. Motovilova, E. Stretchable self-tuning MRI receive coils based on liquid metal technology LiquiTune. Sci Rep 11 , Download citation. Received : 23 March Accepted : 23 July Published : 10 August Anyone you share the following link with will be able to read this content:. Sorry, a shareable link is not currently available for this article. Provided by the Springer Nature SharedIt content-sharing initiative. By submitting a comment you agree to abide by our Terms and Community Guidelines. If you find something abusive or that does not comply with our terms or guidelines please flag it as inappropriate. Sign up for the Nature Briefing newsletter — what matters in science, free to your inbox daily. Skip to main content Thank you for visiting nature. nature scientific reports articles article. Download PDF. Subjects Electrical and electronic engineering Magnetic resonance imaging. Abstract Magnetic resonance imaging systems rely on signal detection via radiofrequency coil arrays which, ideally, need to provide both bendability and form-fitting stretchability to conform to the imaging volume. Introduction Magnetic resonance imaging MRI is an indispensable technique to non-invasively depict anatomic structures and facilitate diagnosis. Results Theoretical analysis The proposed coil geometry comprises a single interdigital capacitor and a rectangular loop Fig. Figure 1. Full size image. Figure 2. Figure 3. Figure 4. Figure 5. Discussion In this work, we present a novel stretchable MRI receive coil design with self-tuning capability, applied to a 3 T MRI system to demonstrate feasibility. Conclusions In this work, we propose a novel concept of stretchable and self-tunable RF coil for MR imaging based on liquid metal integrated within a soft polymer matrix. Methods Theoretical analysis The approximate resonance behavior of the proposed RF coil element can be theoretically analyzed using a simplified circuit model. References Roemer, P. Article CAS Google Scholar Vaughan, J. Google Scholar Sodickson, D. Article CAS Google Scholar Pruessmann, K. Article CAS Google Scholar Griswold, M. Article Google Scholar Winkler, S. Article Google Scholar Fryar, C. Google Scholar Saniour, I. Google Scholar Hardy, C. Article Google Scholar Mager, D. Article Google Scholar Jia, F. Article Google Scholar Corea, J. Article Google Scholar Frass-Kriegl, R. Google Scholar Hosseinnezhadian, S. Article ADS CAS Google Scholar Zhang, B. However, the higher SAR results in fewer slices in imaging and higher susceptibility to artifacts. Transmit receive coils are capable of transmitting an RF signal, then changing modes to receive the MF signal. These types of coils allow for greater slice counts and a reduced chance of artifact detection due to tissue outside of the volume of interest. This coil design is more complex than the receive-only coil. Early coil advancements were made in the SNR gains through improved coil structure. The Signal-to-Noise-Ratio represents signal strength in comparison to variations in intensity due to noise. The signal has to be significantly stronger than the noise and surface coils eliminate the noise from outside the ROI , creating a stronger SNR. Soon after medical scientists realized this, they began working on coils with ore loop elements, arranged in arrays. Each one of these elements receives a signal and feeds it into a separate receive-channel. In order to reduce SNR, these loop elements need to be weighted differently. This coil configuration is called the phased-array coil and is the standard for modern MRI coils. Original array coils, which are the gold standard for MRI coils today, had three areas that could be positioned over different areas of the spine. Increasing the number of array coil channels and mastering the weighing of each channel has allowed for a clearer image of the affected body. Artifacts from sound interference can be cleared much more easily with multi-channel array coils by comparing the readings from each channel when artifacts occur. Further advancements have also been made in channel increases. A complex computer system can now complete scans in milliseconds, allowing for much faster scans. Array coils are designed so that overlap and interference from nearby elements are prevented. This creates a much higher SNR. More advancements were made in the shape and flexibility of MRI coils. Early coils were rigid, one size, one shape, elements that were awkward to maneuver and position over different parts of the body, depending on what was being imaged. Medical scientists identified a need for more flexible coils, and coils that would be able to conform to different body parts. Second generation coils were made from a mix of stiff materials and fabric, which allowed the coil to be bent and braced around body parts. The coil density was also increased. These denser, mixed component coils were more user friendly for medical professionals and technicians because they were lightweight and easier to maneuver. Continued advancements in these areas led the way for even more lightweight coils, including in the actual metal used in the coils themselves. Advancements in science, technology, and engineering have affected and will continue to affect the practicality and efficacy of MRI coils. With the introduction of screen printing and 3D printers, some scientists have been able to experiment with screen printing coils on different materials. Screen printing coils onto different materials can drastically affect the application of the MRI. For example, screen printing MRI coils onto a blanket or other sheet of fabric can eventually help create a workable MRI mechanism for even small children , frail patients, and those with mobility issues. Technology that works to appease both requirements will likely be forthcoming. DirectMed supplies Parts and Services for MRI machines. We sell used parts, make repairs and sell full systems. Our goal is to keep systems in service longer, prolonging their useful life. Navigating Medical Imaging Service: Choosing Between ISOs and OEMs Based on Service Goals. Independent service organizations ISOs are becoming increasingly Success in the medical imaging equipment industry is based upon several crucial abilities: keeping up with technological advancements, providing timely Among the current, rampant supply chain issues and worldwide resource shortages in light of the COVID pandemic and war in Europe, one particular substance For the healthcare industry, the current labor shortage in the biomedical equipment service field—specifically in regard to imaging equipment like MRI Appleton, Wis. For over years, the X-ray tubes used in medical diagnostic systems have proven to be a dependable and cost effective way of producing X-radiation for For decades, the only way to provide high-quality training for imaging engineers was through in-person, hands-on courses at dedicated educational facilities What are MRI Coils? The MRI Coil: A Brief History Radiofrequency coils are the receiving coils for magnetic resonance imaging equipment. The Beginning of RF Coils History In the s, physicist and chemist Michael Faraday deduced the magnetic power of electromagnetic induction by using coiled metals. Advancements in RF Coils Original array coils, which are the gold standard for MRI coils today, had three areas that could be positioned over different areas of the spine. Where Technology is Headed Advancements in science, technology, and engineering have affected and will continue to affect the practicality and efficacy of MRI coils. DirectMed Parts and Service — What we do DirectMed supplies Parts and Services for MRI machines. See what we do or call with questions. Questions, Comments, Concerns? Send Us A Message! |
| Guide to Different Types of MRI Coils - LBN Medical | Zhu, Technologgy. Used under trademark license. Article Google Technologt Adriany, G. Winkler gechnology, Weill Cornell Craving control strategies, MRI coil technology University, United States Riccardo LattanziLangone Medical Center, New York University, United States. Outright Exchange Repair if available. These coils merge the benefits of smaller coils high SNR with the benefits of larger coils large measurement field. |
MRI coil technology -
MRI systems routinely rely on signal detection via receive-only coil arrays, which comprise multiple surface coils arranged to cover the imaging volume 1 , 2. This multi-element configuration, compared to a single element, affords higher signal-to-noise ratio SNR and accommodates accelerated acquisitions 3 , 4 , 5.
Although, most commercial radiofrequency RF receive coils are rigid and inflexible, ideally a coil array needs to provide both bendability and form-fitting stretchability to accommodate various body shapes and sizes to ensure optimal SNR. Commercial RF coils are generally built to accommodate a wide range of anatomical dimensions, which increases the mean offset distance of the coil from the anatomy and therefore reduces the available SNR.
This problem becomes especially challenging when trying to use the same coils for adults as infants or small children 6. Another challenging application is long bone imaging, as the length and circumference of the extremities vary significantly within populations 7.
Some commercial arrays provide limited mechanical flexibility, with portions that can be partially folded around the area of interest; this improves coupling between the imaged volume and the coil and affords a slightly higher filling factor, thus improving RF receive efficiency.
However, these designs are bulky and limited in their flexibility to a single direction. Such flexibility and form-fitting adaptability is achieved by means of a proprietary process that yields low reactance and low-loss conductors while being lightweight, flexible, and durable 9 , 10 , 11 , Highly flexible receive RF coils have been the focus of research for many years offering rigid-bendable 13 , 14 , 15 , 16 , 17 , 18 , 19 , 20 , 21 and geometrically adjustable 22 , 23 , 24 , 25 solutions.
Recent developments in high impedance coaxial coils have demonstrated high flexibility and form-fitting adaptability while also providing good element isolation However, the individual coil diameter cannot be chosen freely as it is dictated by the desired resonance frequency and properties of the coaxial cable.
Multi-turn, multi-gap cable coils provide greater degrees of freedom in terms of coil size 20 , However, commercially available coaxial cables have a limited and discrete set of impedances leading to discrete values of achievable coil diameters. Although the goal of high flexibility in RF coils has been partially solved by the aforementioned designs, full adaptability arguably requires coils to be stretchable as well—a concern, which to date has not been fully addressed.
To achieve stretchability and flexibility, new materials and concepts are required. The use of liquid metal mercury as the coil conductor 27 was suggested as early as Mercury was contained in a flexible plastic tubing to achieve more flexibility, and to accommodate coil positioning closer to the imaging region Gallium-based liquid metals with low vapor pressure render a safer substitute.
Gallium and its alloys have already found many applications in biomedical fields—for example, as restorative materials in dentistry, as tumor imaging and tumor growth suppression agents, and for the treatment of certain cancers The fluidity, low viscosity, and low melting point of gallium-based alloys make them easy to handle with a syringe at room temperature.
When in contact with air, an oxide layer forms on the surface, which prevents the inner metal from further oxidization, while also allowing liquid metals to adhere to surfaces and adopt useful shapes.
The electrical resistivity of gallium is However, as was demonstrated recently in 30 , MRI coils made with alternate conductive materials such as aluminum, liquid metal, conductive polymer, and braided conductors can achieve SNR levels that may exceed intuitive expectations 31 despite their higher resistance compared to commonly used copper-based coils.
As long as the thickness of the conductor is greater than several skin depths, even a significant change in coil resistance will only marginally affect the SNR.
Recent materials science developments allow for the creation of extremely flexible polymers that can undergo significant degrees of stretching, making these polymers feasible as substrates to realize ultra-stretchable conductors with liquid metal encapsulated inside To date, only a few studies have described applying liquid metal technology to stretchable MRI RF coil design 33 , 34 , 35 , 36 , In 34 , the authors directly deposited liquid metal on a stretchable, neoprene fabric to create a flexible knee coil.
In 37 , a stretchable silicone tube with encapsulated liquid metal was used to build a four-element array, also for the knee. These stretchable coils, however, suffer from resonance detuning. When a coil is stretched, its length and thus its inductance is increased, which leads to a resonance frequency shift.
Proposals to mitigate this effect include wide-band matching 36 and automatic tune-match circuitry 35 , However, such designs require additional fixed and inelastic electrical components and circuitry that can dramatically increase complexity and thus make reliability and implementation more complex.
In this study, we describe a method of fabricating soft and stretchable RF receive coils based on liquid metal and ultra-stretchable polymer with a smart geometry design to provide autotuning capability and mitigate the resonance frequency shift, without using additional tuning circuitry.
This proof-of-concept work includes theoretical analysis and numerical simulations, which are subsequently confirmed by bench experiments and MRI tests of a single element RF receive coil at 3 T in vitro and in vivo. The proposed coil geometry comprises a single interdigital capacitor and a rectangular loop Fig.
The liquid metal conducting traces are embedded in a stretchable polymer matrix. To understand the resonance behavior of such a coil under stretch and theoretically describe it in a first order approximation, we developed an analytical circuit model where coil parameters such as trace width, digit length, width and spacing are functions of the degree of stretch.
Approximate analytical formulas for coil inductance and capacitance are then used to calculate how these values change with parameter variation Eqs. Assuming linear stretching along the x-axis, the inductance increases approximately linearly with stretching, while the capacitance of the interdigital capacitor decreases Fig.
Figure 1 c compares the changes in resonance frequency of the control stretchable coil blue color with a fixed value capacitor and that of the proposed coil red color with a single stretchable interdigital capacitor. The smart geometry of the proposed coil helps to reduce the resonance frequency shift.
However, the non-linear change in capacitance cannot fully compensate the linear change of inductance. That is why we still observed frequency shift with stretching. All coil and interdigital capacitor parameters can be dynamically changed in the analytical circuit model, and the optimal range of parameter values can be identified that suit the desired coil requirements.
Although these parameter values provide a reasonable starting point, the accuracy of the analytical model is limited to the first order approximation of the nearest-neighbor interactions, and thus further rigorous numerical simulations were performed as shown in the next section.
a Schematic of the proposed RF coil element. b Inductance and capacitance changes with stretch. c Resonance frequency changes with stretch of the proposed red and control blue coils indicating frequency compensation ability of the stretchable interdigital capacitor. A 3D-model of the proposed coil with an interdigital capacitor was developed Fig.
For comparison purposes, a control coil Fig. Simulation models of the a proposed coil with stretchable interdigital capacitor and b control coil with fixed value capacitor green rectangle.
c Simulated input impedance S 11 change with stretch for the proposed red and control blue coils showing frequency stability of the proposed coil.
As expected, the 3D numerical model yields improved accuracy and illustrates the non-linear behavior of the coil model as compared to the first order approximation in the theoretical model.
This is because the 3D numerical model takes into account 1 higher-order interactions between non-near neighboring digits of the capacitor , 2 material properties and losses, and 3 coil loading, to realistically represent the actual setup.
Fabricated prototypes of the a proposed and b control coils attached to a stretch testing rig. c Simulated and measured frequency shift with strech for the proposed red and control blue coils, indicating improved resonance frequency stability with stretching for the proposed coil.
d Proposed coil prototype placed on a cylindrical phantom with a tape measure for stretching measurement.
Next, the proposed coil was modelled with a homogeneous rectangular phantom that represents average tissue properties. Sensitivity B 1 - field normalized to 1 W input power profiles at the central axial cross-section were simulated at several stretching levels as illustrated in Fig.
This cross-section depicts how coil sensitivity is relatively unaffected by stretching. Although sensitivity at the center decreases slightly, the coil is able to cover a larger area with stretching while maintaining an excellent performance 7.
a — e Simulated sensitivity B1- field profiles normalized to 1 W input power and f — j measured SNR maps at the central axial slice with the numbers on top indicating the percentage of coil stretch.
k — o Measured SNR maps normalized to the coil area demonstrating relative SNR stability with stretching. Figure 3 shows a portion of the bench measurement setup, where two coil prototypes, a proposed and b control, are connected to a 3D-printed unidirectional stretch testing rig.
To more clearly portray the coil performance, the resonance frequency is plotted against the degree of stretch, as shown in Fig. Figure 3 c shows both simulated line with empty markers and measured lines with solid markers changes in resonance frequency with stretching, illustrating that the measured data agree well with the simulated results and demonstrating an improved frequency stability of the proposed coil.
Figure 4 f—j show the SNR maps obtained with the proposed coil when the element is positioned on a homogeneous rectangular phantom. In this first prototype, the Ecoflex ® material used for the stretchable polymer matrix and the liquid metal appear bright on the image; thus, the two points of hyperintensity correspond to the edges of the liquid metal coil, and it is convenient to track the extension of the coil.
The supplementary video V2 demonstrates the measured SNR as the coil stretches. Interestingly, it continues to provide useful SNR even beyond its designed limits Fig. As the SNR of surface coils is a function of coil dimensions 41 , it is expected to decrease for larger coil areas.
This means that, all other parameter being equal, the SNR of a stretchable coil will decrease with the degree of stretch. These normalized maps Fig. The images are compared to their counterpart acquired using c a dedicated 8-channel knee coil.
Only a portion of the knee is visible on the first two images a , b as they were acquired with a single channel surface coil, while the whole knee can be seen in c as it was acquired using a volume multi-channel coil.
An SNR comparison using the same region of interest ROI highlighted as the dashed yellow rectangle in each image highlights the advantage of the proposed coil. Figure 5 d—f show the corresponding SNR maps of the same saggital slices demonstrating the imporved SNR values of the proposed coil compared to the dedicated knee coil.
This significant SNR improvement in the single-element case alone is achieved due to the conformal design of the stretchable coil. We expect to even further improve SNR with an optimized dedicated multi-channel array.
As a first proof of concept, this image demonstrates feasibility of the proposed concept—in particular, the form-fitting conformability of the coil as well as its stretchability and frequency stability, while providing increased SNR as compared to commercial state-of-the-art RF coils.
a — c In vivo knee imaging using different coils showing improved SNR of the proposed coil when comparing the same ROI yellow box. Numbers in red indicate SNR values in the selected ROI red square.
d — f The corresponding SNR maps demonstrating improved SNR. In this work, we present a novel stretchable MRI receive coil design with self-tuning capability, applied to a 3 T MRI system to demonstrate feasibility. Liquid metal embedded in a soft and highly elastic polymer matrix allows one to freely position and stretch the coil to conform to various anatomies.
In addition, the smart geometry of the stretchable interdigital capacitor maintains frequency stability within 0. This initial prototype is designed to provide frequency stability under a unidirectional stretching. A two-dimensional prototype is a straightforward extension of the proposed concept by addition of capacitors in the second dimension and is the subject of current and future work.
In this work, the first proof-of-concept, self-tuning stretchable RF coil is demonstrated in a single channel design. Future work will tackle a multi-channel array design, where decoupling strategies, such as geometric decoupling, will be implemented.
In this purely technical work, we compared the proposed coil to a state-of-the-art commercial knee coil. To the attentive reader, a comparison to other flexible concepts might be of value. A comparison to the more established, flexible-only designs, will be the goal of a more clinical investigation once the LiquiTune technology is built in the form of an array.
Moreover, manganese Mn or gadolinium Gd components can be added to the Ecoflex material used for the stretchable polymer matrix to increase relaxivity and hence reduce signal in MR images.
It was shown in 42 that although addition of silver microparticles reduces the conductivity of gallium, gold microparticles can slightly increase the conductivity of liquid metal. Optimizing the liquid metal alloy composition in this manner may improve the conductivity of liquid metal which reduces the coil losses, improves the Q factor, and effectively increases the SNR of the RF coil further.
Gallium alloys can remain stable within the polymer matrix; however, there is a small chance of liquid metal spilling if there is a tear in the polymer. As an alternative, other flexible conductive materials can be used, such as a conductive elastomer with silver nanoparticles used in Although the conductivity of such materials is inherently lower compared to metal, coils based on such materials can provide comparative Q-factor and SNR values Future work will discuss appropriate sealing methods to further minimize risk of damage or spillage, as well as to facilitate practical requirements in the clinic, such as sanitation.
This will affect the value of the interdigital capacitor and provide another degree for design optimization. In this work, we propose a novel concept of stretchable and self-tunable RF coil for MR imaging based on liquid metal integrated within a soft polymer matrix.
We demonstrate resonance frequency stability within 0. This is in line with the general trend in modern technology where comfort-centered, lightweight and wearable devices are at the forefront of novel developments. MRI technology is bound to follow suit with the help of safe liquid metals and ultra-stretchable polymers.
The approximate resonance behavior of the proposed RF coil element can be theoretically analyzed using a simplified circuit model. When the dimensions of a coil loop element change, its resonance frequency changes accordingly.
The frequency shift can be understood through the resonance equation,. The coil inductance is proportional to the coil conductor length, which in turn indicates that if a coil is stretched, the coil inductance is increased and the resonance frequency is decreased.
To mitigate the frequency shift, we use an interdigital capacitor that decreases its capacitance under stretch. The total capacitance of an interdigital capacitor can be approximated as follows,. Although the complete coil behavior under stretch is more complicated and requires rigorous numerical modeling, this simplified theoretical analysis demonstrates in first order approximation how the integration of a stretchable interdigital capacitor into an RF loop coil can help to reduce coil detuning under stretch.
Once approximate parameter values are obtained from the theoretical analysis, full-wave numerical simulations are required to perform exact parametric optimization of all dimensions. To this goal, a single-element coil, as shown in Fig.
The minimum conductor width was chosen based on the manufacturing limitations specification of the 3D printer. The frequency stability of the coil was evaluated for different degrees of unidirectional stretch along the x -axis, and the capacitor parameters are optimized to minimize the frequency shift.
For comparison purposes, a second, geometrically identical, prototype of the coil was modeled where a fixed value capacitor was used instead of the flexible interdigital one.
Both coils were tuned to MHz the operating frequency of a 3 T scanner and matched to 50 Ohm at their unstretched state. The coil was fed using a uniform port excitation with an input power of 1 W. The field maps were acquired at the central axial slice going through the middle of the coil.
The sensitivity at the surface of the coil was measured at the coil center for each stretch position. Detailed step-by-step instructions with illustrations of the coil manufacturing process are shown in the Supplementary Materials Figure S1 and the corresponding text.
The plastic molds that contain microfluidic channel features of the desired coil geometry and the sealing featureless top layer are fabricated using a high-resolution 3D printer Prusa i3 MK3S, using a 25 µm nozzle and a 50 µm layer height with polylactic acid PLA material.
Liquid metal is injected into the embedded microchannels using a needle and syringe to form conducting traces. The Supplementary video V1 demonstrates the stretchability of the fabricated coil. Integrated coils are uniquely designed for improved workflow and patient comfort.
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Scale in dB. Axial top row and sagittal bottom row flip angle maps measured with the 6-channel knee coil absent left and present right illustrate negligible interaction with the system body coil during transmission.
The six-channel cable array included MR-visible components. SNR maps show that the 0. To provide another context for comparison, SNR was We scanned another subject three times in the same day interscan interval of approximately 5 min to assess repeatability with the 0.
Representative SNR maps acquired with the 0. SNR values overlaid in the bottom right are averaged over three volunteers in a 3-cm ROI in the distal femur overlaid red circles. Coil photos are inset in the bottom left of each panel. Text overlays indicate the maximum geometry factors g max.
Figure 13 shows representative T2-weighted and T1-weighted images of the left knee of a year-old man acquired at 0. Acceptable image quality was observed in all three cases.
The images revealed diffuse intermediate signal within the posterior horn and body of the meniscus consistent with mucinous degeneration with no meniscal tear noted.
Figure 14 illustrates cartilage and bone injuries in the right knee of a year-old man. Images of the left knee of a year-old woman show undesirable signal hotspots that are a consequence of the 0. Evidence of heterogeneous fat suppression can also be observed in images acquired with both 0.
Image quality was similar in all three sessions of the 0. Representative turbo spin echo MR images of a year-old man acquired using the 0. Mucinous degeneration with no meniscal tear is observed in the coronal images arrows.
Turbo spin echo MR images of a year-old man acquired using the 0. The images reveal patellar cartilage defects solid arrows and a lateral femoral condyle bone contusion open arrows.
The bone contusion was not present at the time of the 1. Turbo spin echo MR images from a year-old woman acquired with the 0. The cable coil images have peripheral signal saturation solid arrows.
Images acquired with all coils had heterogeneous fat suppression hollow arrows. Turbo spin echo MR images from a year-old man acquired with the 0. We designed and implemented a knee coil for 0.
The cable loops performed similar to semi-rigid Cu-FR4 conductors in terms of SNR. The cable loops also inherited the benefits of coaxial cables such as low cost, widespread availability, ease of assembly, and high durability. While coaxial cables cannot match the stretchability and elasticity of other specialized conductors 3 , 4 , 5 , 12 , it remains unclear if such features are necessary for routine knee MRI.
The selected RG cable has a minimum bend radius of 2. This flexibility ensured surprisingly good coil-tissue coupling Q-ratio of 2. To put this into context, our group settled for a Q-ratio of only 2 in a previously described semi-flexible six-channel knee coil 19 despite a higher operating frequency We anticipate that the straightforward concept of forming loops from cable conductors can be translated to other frequencies.
Preliminary work from our group showed that a loop formed from the shield of RG cable had a similar Q-ratio as a Cu-FR4 loop at We found that an individual cable loop provided improved Q-ratio and SNR at 0.
This trend agrees with 3 T SNR measurements in Fig. Notably, the 3 T conventional loop SNR advantage grew as the loop diameter decreased, suggesting that coaxial conductor loss is greater than that of conventional loops.
The additional conductor loss associated with coaxial loops appears to arise from a longer conductor pathway that includes both outer and smaller inner conductors. Indeed, Fig. Two-channel coaxial coil arrays provided improved decoupling and tuning robustness against geometrical and overlap variation compared to cable loops Figs.
However, we found that at 0. One explanation for the apparent discrepancy is that decoupling and tuning have a complex relationship with SNR because of interplay among additional properties such as coil loss, noise correlation, matching circuitry, and preamplifier impedance mismatch 22 , 23 , 24 , 25 , 26 , 27 , 28 , 29 , From our data on two-channel arrays, we selected cable coils for the six-channel knee array because the plastic hinges Fig.
It is worth noting, however, that these geometric assumptions can breakdown. For example, the two neighbor coils along the seam of the 6-channel array may be incidentally configured with suboptimal overlap for particular knee sizes.
Advanced impedance matching techniques could be applied to alleviate SNR degradation in such configurations 3 , 19 , 26 , 29 , 30 , We selected RG coaxial cable in this study because it provided low loss and good flexibility.
Given that the inner conductor and insulator had little effect on Q of the cable coil, it appears that coaxial cable properties such as characteristic impedance and dielectric constant of the insulator that are important for signal transmission were not relevant for cable coils with cm diameter at This suggests that commercially available flexible single conductor braided wires warrant consideration for forthcoming cable loop designs.
We applied the cable coil concept to knee MRI, which is the most widely accepted imaging tool for diagnosing various knee injuries Recently, a new-generation whole-body 0. The system can offer some of the advantages of dedicated extremity scanners such as greater accessibility due to lower cost compared to high-field systems, improved patient comfort given its cm wide bore, as well as reduced susceptibility artifacts and specific absorption rate Among the various options, we matched as closely as possible the imaging parameters at both fields.
While the 1. Despite being limited by suboptimal receive coil technology, prior studies conducted on low-field whole-body or specialized extremity scanners demonstrated reliable diagnosis of meniscal and ligamentous tears 36 , 37 , 38 , 39 , 40 , 41 , 42 , 43 , However, SNR challenges may have resulted in difficulties such as inconspicuous cartilage abnormalities depending on field strength and system 39 , 45 , 5—6 mm thick slices 38 , 43 , 44 that exceeded that set forth in the American College of Radiology guidelines 4 mm 46 , or undesirably long examination times 36 , In this study, the acquisition time was approximately 3 to 4 min per pulse sequence and the spatial resolution was 0.
While this study was not intended to evaluate knee joint abnormalities explicitly, cartilage and bone injuries were incidentally delineated, suggesting that clinically acceptable knee examinations are feasible at 0.
In reporting that the proposed 6-channel cable coil provided an approximately twofold SNR advantage over the makeshift prototype coil, we acknowledge that the prototype coil is far from optimal for knee imaging. Nonetheless it was the best option provided by the manufacturer, which in fact motivated the current work.
To better put the SNR into context, we showed that the 0. This discrepancy can be explained by T1 differences; T1 of bone is ms at 1. After correcting for T1, we find a 3. One drawback of the cable coil is that its tight fit resulted in signal hotspots, particularly in T1w images.
Saturated signal was evident in the knee periphery Fig. Further study is required to determine whether clinical readability is impacted or whether image uniformity can be improved by methods such as bias field correction in post-processing While this can be undesirable for ultra-short TE applications such as bound water imaging, it is not anticipated to be problematic for conventional or rapid knee imaging protocols that primarily utilize spin echo sequences with TEs on the order of 10 ms or longer We observed non-uniform fat suppression in 0.
Given the narrow spectral separation between fat and water at 0. The chosen coil dimensions 6 loops of 10 cm diameter that were overlapped for decoupling were suitable for encircling a cylinder with However, in practice, the distal thigh is the limiting factor for accessibility.
To accommodate a greater portion of the population, a future design may utilize slightly larger loops or foam cutouts that are fitted to the loops rather than the continuous foam sheet in the current coil.
The coil presented here has a single-row layout that makes it sensitive to placement and incapable of longitudinal parallel imaging acceleration In principle, these limitations could be overcome by replicating the existing single-row layout to form an "Olympic ring" two-row array similar to that in Rispoli et al.
Such a layout may deserve future exploration. In conclusion, the proposed cable coil utilized flexible, commercially available conductors, which showed comparable SNR as a reference Cu-FR4 loop and adequate SNR robustness against geometric variability.
Applied to knee imaging, the cable coil array provided promising image quality, particularly in the cartilage, which has been difficult to examine with older generation low-field systems 39 , 45 , in a clinically acceptable examination time 36 , Natural extensions of this work will be to apply the coaxial shield conductor concept to other anatomies and other operating frequencies and to rigorously evaluate image quality and diagnostic accuracy in a larger cohort.
The loops were made to resonate at Louis, MO or 1 oz. copper plated FR4 circuit board Cu-FR4 with mm trace width Fig.
All loops were cm in diameter. For the cable loop, the capacitor was in series with the outer shield, while the inner conductor was broken at the feed port but not connected electrically floating.
For the coaxial loop, gaps were placed in the RG shield and inner conductor at opposite locations along the circumference, and tuning was carried out using a capacitor in the inner conductor gap.
Scattering matrix and imaging measurements were performed with each coil connected to a circuit board that contained a preamplifier port and components for matching, detuning, and preamplifier decoupling.
Preamplifier decoupling efficacy was 20 to 25 dB for the cable loops and coaxial loops. This measurement was defined as the difference between S 21 measured with a double pickup probe coupled lightly to the coil without and with the preamplifier present.
The cable and coaxial loop detuning circuits provided at least 30 dB isolation. This measurement was defined as the difference between S 21 measured with a double pickup probe coupled lightly to the coil with the detuning circuit inactive and active.
Sensitivity to geometric variability was investigated by measuring the reflection coefficient at the preamplifier port on a cable loop and a coaxial loop that were tuned with cm diameter circular contours.
Without adjusting the tuning, the measurement was repeated after elongating the loops into ellipses whose major axes were cm and cm. Sensitivity to overlap was investigated by measuring reflection and transmission coefficients at the preamplifier ports on two-channel arrays based on cable or coaxial loops.
The coils were tuned in isolation. MRI data were acquired on two commercial systems: 1 a 1. SNR maps were measured in the phantom with one or two-channel arrays based on cm cable, coaxial, and Cu-FR4 loops. SNR maps were calculated from signal and noise with the RF pulse amplitude set to zero measurements acquired with a gradient echo pulse sequence and processed with the optimal array combination method 1 , The phantom SNR maps were reconstructed with 2.
Geometry factor maps 18 were calculated from the same data using methods described by Montin and Lattanzi Six cable loops were arranged to form a knee coil array. Fuses were added to the circuit boards to prevent current flow during body coil transmission in the event of an active detuning circuit failure.
The boards were enclosed in rigid plastic and the assembly in flame-resistant fabric Fig. Adjacent loops were linked together by ABS plastic polymer acrylonitrile butadiene styrene hinges that allowed mechanical flexibility while maintaining approximate geometric overlap for inductive decoupling.
Knee MRI was performed in 3 configurations: 1 at 0. For case 2, the coils were repurposed for knee imaging by wrapping the body coil around the anterior knee and using the spine coil to cover the posterior knee. Note that the body coil was not sufficiently flexible to tightly fit the knee; its minimum bend radius was approximately 11 cm.
The study was fully compliant with the Health Insurance Portability and Accountability Act and the New York University Institutional Review Board approved the protocol.
All experiments were performed in accordance with relevant guidelines and regulations. We scanned four human subjects after obtaining their informed written consent.
We performed 2D turbo spin echo imaging to evaluate the efficacy of the coils for clinical research. The 0. Fat suppression was performed with the product spectrally selective saturation method. The 1.
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Anti-angiogenesis approaches in medicine is a wide range of MRI technologgy types available on the market. Radiofrequency RF coils have undergone hechnology development over the past few MRI coil technology, from the single loop to a quadrature, and eventually to the MMRI array coil. Without getting too much into that, we would like to give you a simple, but informative introduction to the world of this magnetized copper. So, what exactly is an MRI coil? A coil, by definition, is anything wound in a joined sequence of concentric rings. In other words, it is a looped length of wire. The technical design and engineering of each coil differ, and the differences often relate to their functionality.
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