High Field Imaging
Sebastian Schmitter1

1Physikalisch-Technische Bundesanstalt, Braunschweig and Berlin, Germany

Synopsis

The MRI main field strength has been constantly increased over the past decades and today, scanners with 3T, 7T and even beyond are in use. However, ultrahigh-field (>7T) systems are still mostly used in research centers although a transition into hospitals is expected. The reasons for using (ultra-)high fields are multifold and will be outlined in this presentation Along with these benefits go a larger range of challenges, which are among the reasons for the rather slow transition of UHF into clinical applications. Solutions to most of these challenges will be presented and applications will be highlighted.

Target audience

Clinicians and scientists, who are interested in the benefits, the problems and practical solutions in high and ultra-high field MRI.

Highlights

  • An increased main magnetic field strength provides higher signal-to-noise ratios (SNR) that allow for higher spatiotemporal resolution or shorter acquisition times.
  • T1 relaxation typically increases, which is highly beneficial e.g. for inflow-based angiographic sequences. T2* decreases with increasing field strength requiring shorter echo times to compensate the signal loss.
  • Higher field strength allows for increased encoding capability in parallel imaging, thus higher acceleration factors may be used and scan time may be beneficial.
  • Susceptibility effects increase with field strength. This effect is highly beneficial in SWI, QSM or fMRI, but also creates artifacts and problems concerning the imaging process.
  • Variations of the transmit B1 field (B1+) generate spatially varying flip angles and contrast. This is efficiently addressed using B1+ shimming or parallel transmission.
  • SAR and RF power are substantial problems at ultra-high field. Often RF pulses need to be stretched or other RF power reduction techniques need to be applied. SAR and/or RF power can be included in RF pulse design.
  • Applications for HF and UHF imaging range from high-resolution anatomical imaging, vascular imaging, susceptibility weighted imaging (SWI), quantitative susceptibility mapping (QSM), functional MRI (FMRI), X-nuclei imaging, spectroscopic imaging and others.

Motivation

Since the beginning of magnetic resonance imaging, the main magnetic field strength (B0) has constantly increased. Currently, most clinical MRI systems operate at B0 of 1.5 Tesla (standard field strength) or 3 Tesla (high field strength). In 1998 an 8T system was installed in Ohio and since then, the number of ultra-high field (UHF) systems (B0≥7T) has substantially increased to approximately 70 systems worldwide today.

Most of today’s UHF systems are being used for research purposes. However, the transition into hospitals for dedicated applications is expected, because recently, a CE marked/FDA approved clinical 7T system has been introduced.

Benefits

One of the major advantages of using higher fields is an increase in signal-to-noise ratio (SNR) (1,2), which is an important quantity that impacts the maximum spatial and temporal resolution and the ability to depict structures in MR images. While the magnetization increases linearly with B0, the theoretical SNR gain is more difficult to assess (3). In practice, a linear to quadratic increase in SNR with field strength has been reported, however, the SNR at UHF is spatially dependent (1,2,4). An SNR increase by a factor 6-10 from 1.5T to 7T appears to be realistic. The higher SNR is often used to increase the spatial or temporal resolution. Higher spatiotemporal resolution, however, requires an increased measurement time, thus acceleration techniques become increasingly important with rising field strength. Fortunately, the parallel imaging performance increases with B0, allowing higher acceleration factors (5). Other acceleration methods, such as KT-acceleration, compressed sensing or Multiband imaging are promising with regard to limiting scan times (6-9).

The increased SNR provided by UHF is also promising for nuclei other than protons, such as 13C, 17O, 19F, 23Na or 31P. Signals caused by those “X-nuclei” are typically orders of magnitude lower than for protons, because the concentration in the human body and the gyromagnetic ratio γ is lower compared to protons (note that the MR signal scales with γ3). The higher SNR allows for acceptable resolutions within clinically feasible scan times.

The T1 relaxation constant increases with B0 field strength (10). A longer T1 value in combination with higher SNR is particularly beneficial for inflow-based MR-angiography techniques, such as time-of-flight (TOF) (11,12). In this case the contrast benefits twice from higher B0: the longer T1 values result in stronger suppression of the static background tissue while the full magnetization of the freshly inflowing blood provides stronger signal compared to lower fields. Blood flow quantification techniques such as 4D flow MRI similarly benefit from higher fields (13). For other sequences, however, a longer T1 constant may need to be compensated by longer TR in order to achieve acceptable contrast. While T1 increases with static field strength the T2* relaxation constant decreases. In practice, this is rather challenging, as it results in signal loss and requires shorter echo times to compensate this effect. The magnetic susceptibility χ of the tissue causes small deviations ΔB0 of the local magnetic field. Because ΔB0 scales linearly with B0, susceptibility induced changes in ΔB0 will increase at UHF, thus susceptibility weighted imaging (SWI) or quantitative susceptibility mapping (QSM) benefit from UHF (14,15). The sensitivity towards the susceptibility of deoxygenated blood also increases, leading to an increased blood-oxygenation level dependent (BOLD) contrast (2).

MR spectroscopy benefits substantially from higher magnetic fields due to an increased chemical shift (16). Thus, the frequency difference between different metabolite resonances increases linearly with B0 allowing for higher spectral resolution and improved spectral quantification.

Challenges

In order to excite the spins a radiofrequency (RF) field is applied by the RF coil, consisting of a transmit magnetic (B1+) and an electric component (E). The frequency of the RF field increases linearly with B0 and at 7T the resulting RF wavelength achieves values of ~11cm in the tissue. Such wavelength is in the order of organ sizes and results in spatially heterogeneous B1+ amplitudes/phases at UHF (17) causing spatially dependent flip angles and unwanted spatially varying contrast. In practice, most commercially available head coils generate a birdcage or circularly polarized (CP) mode during excitation that leads to high FA in the brain center and reduced FA in the brain periphery. Typical values in the periphery are about 50% of the center B1+ amplitude values at 7 Tesla, however, even complete FA voids can occur depending on the RF coil and field strength.

Another substantial challenge is related to the power of the applied RF field. In a simple model, a quadratic increase of the RF power with field strength is expected while a more detailed analysis shows deviations from the quadratic increase towards slightly lower values starting at around 3T (3). Higher RF power has substantial impact on MR hardware and MR safety. The required peak RF voltages increase, thus stronger amplifiers are required. More importantly, the specific absorption rate (SAR), i.e. the RF power deposited in the tissue increases. With increasing B0 the focus shifts from the global SAR, i.e. the SAR averaged over a large body region or the entire body, to the local 10g-averaged SAR. Both SAR quantities are limited according to IEC guidelines; in practice, the local SAR becomes more important at UHF because it typically reaches the limits first. Although the total RF power fed to the RF coil can be measured during MR imaging and thus can provide an estimate for the global SAR, the local SAR cannot be assessed in vivo with sufficient accuracy. Therefore, electromagnetic simulations are conducted for different human body models and the spatial distribution of E and B1+ field are calculated as well as resulting SAR values are derived. Further challenges at high and ultra-high field are related to the increase susceptibility that creates strong ΔB0 particularly at air-tissue boundaries. At 7T, ΔB0 values corresponding to frequencies of >400 Hz are observed in the brain above the nasal cavities and within the temporal lobes. This typically leads to signal reduction or signal voids in gradient echo images. Delta B0 also impacts balanced steady-state free precession (SSFP) imaging, which shows banding artifacts in regions with strong ΔB0. This effect is frequently observed in body imaging at 3 Tesla (18). Furthermore, B0 inhomogeneity also creates distortions, particularly in single shot acquisitions such as EPI. To limit this effect, acquisition time needs to be reduced by parallel imaging or other methods.

Another substantial practical challenge at UHF is the ability to detect the cardiac cycle using an electrocardiogram (EKG). The magneto-hydrodynamic (MHD) effect induces an elevated T-wave in the EKG signal, which in practice often leads to false detection of the R-peak (19).

Solutions

Variations of the FA can be reduced by using adiabatic RF pulses instead of standard RF pulses. However, a limitation of these pulses is the fact that they typically require higher RF power. Another promising solution for this problem consists in parallel transmission (pTX) of different RF pulses on multiple transmit (TX) coil elements (20,21). In B1-shimming (3), a sub-case of pTX, all coil elements share the same RF pulse shape, but individual (time-constant) transmit phases and/or amplitudes are set for each RF coil element in such a way that the spatial B1+ variations are minimized. This approach has proven to be successful in a large range of applications and targets, particularly for targets within the body (22). In ‘true’ parallel transmission (pTX) N independent RF pulses are applied to the N TX coil elements. This approach is particularly beneficial for localized excitations. An intermediate solution consists of ‘spokes pulses’ for the slice selective case and ‘kt-pulses’ for the non-selective case (23, 24). Both pulses consist of composed, B1-shimmed sub-pulses (e.g. SINC or RECT pulses) together with gradient blips in between the sub-pulses. The combination of blips and B1-shimming allows for improved homogeneity within the target volume compared to B1-shimming alone. RF power and SAR can be addressed in several ways. Stretching a given RF pulse by a factor f reduces the RF power by 1/f. Similarly, the VERSE principle (25)can be applied, which allows stretching only those parts of the RF pulse with high RF amplitude. State-of-the-art pTX RF pulse designs include the global SAR, the local SAR as well as the RF peak power as a constraint (26). In practice, a tradeoff between SAR/power and excitation fidelity needs to be found.

ΔB0 is typically reduced by B0 shimming; typically second-order shimming is performed but some systems are equipped with shims of even higher orders. On the excitation side, ΔB0 can be incorporated into the RF pulse design if necessary. Increased imaging speed that is needed to address longer acquisition times due to higher spatial resolution is addressed using parallel imaging, by kt-acceleration (e.g. kt-blast or kt-GRAPPA), by compressed sensing or by applying Multiband (6-8). Body applications at ultra-high field are still challenging, in this case a single universal shim setting (such as the CP mode for the head) that can be applied for all body targets and body shapes, does not seem to exist. However, the so-called ‘TIAMO’ technique combines two complementary shim solutions, which is a practical and robust approach (27). Body applications that require cardiac triggering are more difficult to perform at 3T and particularly at UHF because of the MHD effect. A solution to this problem was proposed by Frauenrath et al (28), who presented an acoustic cardiac triggering system that operates similar to a stethoscope.

Applications

While clinical 3T systems were already applied for head imaging soon after their introduction, the transition from 1.5T towards 3T for body applications such as the heart has been rather slow. Although the time scale is larger, a similar trend can be observed for applications at 7T. Currently most studies are targeting the brain while body applications remain rare. Most applications aim for higher resolution or stronger contrast using structural, functional or spectroscopic imaging. Applied techniques which appear to benefit most from UHF are angiographic techniques, such as TOF, susceptibility weighted imaging and QSM, gradient and spin echo functional MRI, diffusion weighted imaging, spectroscopic MRI and X-nuclei imaging (2,14-16). These techniques have been applied to a large range of diseases and disorders, including tumor imaging, microvascular diseases, multiple sclerosis, epilepsy, dementia, Alzheimer’s disease, Parkinson’s disease and other neurodegenerative diseases (14,29-32). Functional MRI at 7T in combination with high acceleration factors using parallel imaging and multiband allows for whole-brain fMRI imaging with 1mm isotropic resolution (33). Similar protocols are applied in the Human Connectome Project as well as in other functional studies performed worldwide.

Conclusion

Although UHF MRI is connected to a variety of challenges, practical solutions have been presented over the past decade for a major part of the problems. This paves the way towards anatomical, functional and spectroscopic head imaging for clinical applications. Body imaging at UHF is substantially more challenging which makes the transition into clinical routine much more difficult.

Acknowledgements

No acknowledgement found.

References

  1. Vaughan JT, Garwood M, Collins CM, Liu W, DelaBarre L, Adriany G, et al. 7T vs. 4T: RF power, homogeneity, and signal-to-noise comparison in head images. Magn Reson Med. 2001;46(1):24-30.2.
  2. Ugurbil K, Adriany G, Andersen P, Chen W, Garwood M, Gruetter R, et al. Ultrahigh field magnetic resonance imaging and spectroscopy. Magn Reson Imaging. 2003;21(10):1263-81.3.
  3. Hoult DI. Sensitivity and power deposition in a high-field imaging experiment. J Magn Reson Imaging. 2000;12(1):46-67.4.
  4. Otazo R, Mueller B, Ugurbil K, Wald L, Posse S. Signal-to-noise ratio and spectral linewidth improvements between 1.5 and 7 Tesla in proton echo-planar spectroscopic imaging. Magn Reson Med. 2006;56(6):1200-10.5.
  5. Wiesinger F, Van de Moortele PF, Adriany G, De Zanche N, Ugurbil K, Pruessmann KP. Parallel imaging performance as a function of field strength--an experimental investigation using electrodynamic scaling. Magn Reson Med. 2004;52(5):953-64.6.
  6. Huang F, Akao J, Vijayakumar S, Duensing GR, Limkeman M. k-t GRAPPA: a k-space implementation for dynamic MRI with high reduction factor. Magn Reson Med. 2005;54(5):1172-84.7.
  7. Larkman DJ, Hajnal JV, Herlihy AH, Coutts GA, Young IR, Ehnholm G. Use of multicoil arrays for separation of signal from multiple slices simultaneously excited. J Magn Reson Imaging. 2001;13(2):313-7.8.
  8. Lustig M, Donoho D, Pauly JM. Sparse MRI: The application of compressed sensing for rapid MR imaging. Magn Reson Med. 2007;58(6):1182-95.9.
  9. Moeller S, Yacoub E, Olman CA, Auerbach E, Strupp J, Harel N, et al. Multiband multislice GE-EPI at 7 tesla, with 16-fold acceleration using partial parallel imaging with application to high spatial and temporal whole-brain fMRI. Magn Reson Med. 2010;63(5):1144-53.10.
  10. Rooney WD, Johnson G, Li X, Cohen ER, Kim SG, Ugurbil K, et al. Magnetic field and tissue dependencies of human brain longitudinal 1H2O relaxation in vivo. Magn Reson Med. 2007;57(2):308-18.11.
  11. Kang CK, Park CW, Han JY, Kim SH, Park CA, Kim KN, et al. Imaging and analysis of lenticulostriate arteries using 7.0-Tesla magnetic resonance angiography. Magn Reson Med. 2009;61(1):136-44.12.
  12. Schmitter S, Bock M, Johst S, Auerbach EJ, Ugurbil K, Van de Moortele PF. Contrast enhancement in TOF cerebral angiography at 7 T using saturation and MT pulses under SAR constraints: impact of VERSE and sparse pulses. Magn Reson Med. 2012;68(1):188-97.13.
  13. van Ooij P, Zwanenburg JJ, Visser F, Majoie CB, vanBavel E, Hendrikse J, et al. Quantification and visualization of flow in the Circle of Willis: time-resolved three-dimensional phase contrast MRI at 7 T compared with 3 T. Magn Reson Med. 2013;69(3):868-76.14
  14. Theysohn JM, Kraff O, Maderwald S, Barth M, Ladd SC, Forsting M, et al. 7 tesla MRI of microbleeds and white matter lesions as seen in vascular dementia. J Magn Reson Imaging. 2011;33(4):782-91.15.
  15. Haacke EM, Liu S, Buch S, Zheng W, Wu D, Ye Y. Quantitative susceptibility mapping: current status and future directions. Magn Reson Imaging. 2015;33(1):1-25.16.
  16. Tkac I, Oz G, Adriany G, Ugurbil K, Gruetter R. In vivo 1H NMR spectroscopy of the human brain at high magnetic fields: metabolite quantification at 4T vs. 7T. Magn Reson Med. 2009;62(4):868-79.17.
  17. Van de Moortele PF, Akgun C, Adriany G, Moeller S, Ritter J, Collins CM, et al. B(1) destructive interferences and spatial phase patterns at 7 T with a head transceiver array coil. Magn Reson Med. 2005;54(6):1503-18.18.
  18. Schar M, Kozerke S, Fischer SE, Boesiger P. Cardiac SSFP imaging at 3 Tesla. Magn Reson Med. 2004;51(4):799-806.19.
  19. Suttie JJ, Delabarre L, Pitcher A, van de Moortele PF, Dass S, Snyder CJ, et al. 7 Tesla (T) human cardiovascular magnetic resonance imaging using FLASH and SSFP to assess cardiac function: validation against 1.5 T and 3 T. NMR Biomed. 2012;25(1):27-34.20.
  20. Katscher U, Bornert P, Leussler C, van den Brink JS. Transmit SENSE. Magn Reson Med. 2003;49(1):144-50.21.
  21. Zhu Y. Parallel excitation with an array of transmit coils. Magn Reson Med. 2004;51(4):775-84.22.
  22. Metzger GJ, Snyder C, Akgun C, Vaughan T, Ugurbil K, Van de Moortele PF. Local B1+ shimming for prostate imaging with transceiver arrays at 7T based on subject-dependent transmit phase measurements. Magn Reson Med. 2008;59(2):396-409.23.
  23. Cloos MA, Boulant N, Luong M, Ferrand G, Giacomini E, Le Bihan D, et al. kT -points: short three-dimensional tailored RF pulses for flip-angle homogenization over an extended volume. Magn Reson Med. 2012;67(1):72-80.24.
  24. Setsompop K, Alagappan V, Gagoski B, Witzel T, Polimeni J, Potthast A, et al. Slice-selective RF pulses for in vivo B1+ inhomogeneity mitigation at 7 tesla using parallel RF excitation with a 16-element coil. Magn Reson Med. 2008;60(6):1422-32.25.
  25. Conolly S, Nishimura DG, Macovski A, Glover G. Variable-rate selective excitation. J Magn Reson. 1988;78:440--77.26.
  26. Guerin B, Gebhardt M, Cauley S, Adalsteinsson E, Wald LL. Local specific absorption rate (SAR), global SAR, transmitter power, and excitation accuracy trade-offs in low flip-angle parallel transmit pulse design. Magn Reson Med. 2014;71(4):1446-57.27.
  27. Orzada S, Maderwald S, Poser BA, Bitz AK, Quick HH, Ladd ME. RF excitation using time interleaved acquisition of modes (TIAMO) to address B1 inhomogeneity in high-field MRI. Magn Reson Med. 2010;64(2):327-33.28.
  28. Frauenrath T, Hezel F, Renz W, d'Orth Tde G, Dieringer M, von Knobelsdorff-Brenkenhoff F, et al. Acoustic cardiac triggering: a practical solution for synchronization and gating of cardiovascular magnetic resonance at 7 Tesla. J Cardiovasc Magn Reson. 2010;12:67.29.
  29. Versluis MJ, van der Grond J, van Buchem MA, van Zijl P, Webb AG. High-field imaging of neurodegenerative diseases. Neuroimaging Clin N Am. 2012;22(2):159-71, ix.30.
  30. Tallantyre EC, Morgan PS, Dixon JE, Al-Radaideh A, Brookes MJ, Evangelou N, et al. A comparison of 3T and 7T in the detection of small parenchymal veins within MS lesions. Invest Radiol. 2009;44(9):491-4.31.
  31. Henry TR, Chupin M, Lehericy S, Strupp JP, Sikora MA, Sha ZY, et al. Hippocampal sclerosis in temporal lobe epilepsy: findings at 7 T(1). Radiology. 2011;261(1):199-209.32.
  32. Lupo JM, Banerjee S, Hammond KE, Kelley DA, Xu D, Chang SM, et al. GRAPPA-based susceptibility-weighted imaging of normal volunteers and patients with brain tumor at 7 T. Magn Reson Imaging. 2009;27(4):480-8.33.
  33. Ugurbil K, Xu J, Auerbach EJ, Moeller S, Vu AT, Duarte-Carvajalino JM, et al. Pushing spatial and temporal resolution for functional and diffusion MRI in the Human Connectome Project. Neuroimage. 2013;80:80-104.
Proc. Intl. Soc. Mag. Reson. Med. 26 (2018)