• Fast non-field-cycled non-cryogenic ULF MRI is possible in a simple platform

• Coil design in the Johnson noise-dominated regime is different from the usual body-noise dominated

• Increased fractional B0 homogeneity at ULF allows b-SSFP over large FOVs in simple magnets.

• Low magnetic susceptibility allows for imaging around implants • Strategies for contrast at ULF include quantitative MRF.

• Opportunities for in vivo Overhauser DNP exist at ULF.

Specialty area: Portable MRI

Speaker name: Matthew S. Rosen, mrosen@cfa.harvard.edu

Target audience: MRI scientists

Outcome/Objectives: Attendees will gain insight into strategies for efficient signal acquisition at ultra-low field MRI, as well as new opportunities for purpose-built portable scanners operating in the millitesla regime.

Purpose: This presentation is intended to provide a understanding of the need for different approaches to to MRI at high- and ultra-low field, both in terms of hardware, acquisition strategies, and approaches to contrast.

MRI is a powerful, non-invasive technique for revealing the internal structure and function of the human body with a rich range of biological contrasts. Despite considerable improvements in imaging quality and speed, the underlying technology remains remarkably unchanged compared to the first generation scanners that emerged on the market 30 years ago. The fact that very strong Tesla-scale magnetic fields are needed to overcome the intrinsic lack of sensitivity of NMR-based methods continues to dominate scanner construction, and drives both pricing and scanner siting requirements. These massive magnets preclude truly portable operation in many environments including surgical intervention, triage and primary care suites.

A promising approach to portable MRI is operation ultra-low magnetic field where scalable electromagnets become practical. The ultra-low field (ULF) regime is defined [1] when the magnetic field used for signal detection is below 10 mT. Operation at ULF enables imaging in environments where high magnetic fields would be contraindicated (such as in the presence of nearby ferrous materials), and raises the potential for scanners to be built at significantly reduced total cost, and with open geometry designs that ease patient handling and positioning.

In an effort to improve the performance of operation at low and ultra-low magnetic field, Macovski and Conolly introduced the concept of pre-polarized MRI (also known as PMRI) in 1993 [2], which employs a strong, inhomogeneous pulsed magnet field to generate increased nuclear polarization, and a second much weaker homogeneous magnetic field for signal detection. This PMRI strategy has been the acquisition strategy for nearly all low- and ultra-low field MRI systems since its introduction, however pre-polarized ULF MRI suffers from intrinsically long acquisition times, most of which is uncompressible, that result from the time needed to generate nuclear polarization.

In our recently published work [3], we demonstrate a novel approach to fast and efficient brain ULF MRI at 6.5 mT with no pre-polarization nor cryogenics, combining under-sampling strategies with a high performance fully refocused steady-state-based acquisition on a simple, inexpensive platform (Fig. 1). With a novel inductive single channel detector, and robust hardware, we have achieved the fastest 3D MRI of the living human brain in the ULF regime compared to the state-of-the-art as reported in the literature [4]-[8].

Ultra-low field acquisition strategy:

High performance imaging at ultra-low magnetic field focuses on substantially reducing acquisition time using fast imaging techniques. This is especially important when significant signal averaging is required as a result of low Boltzmann polarization. Our approach to fast imaging at ULF is 3D balanced steady state free precession sequences (b-SSFP) [9]. Unlike traditional gradient- and spin-echo techniques, b-SSFP sequences dynamically refocus spin magnetization following measurement (Fig. 2), eliminating the extra delays typically used for T₂ decay and T₁ recovery. This considerably reduces acquisition times and provides the highest SNR per unit time of all imaging sequences [9], [10]. Our use of undersampled b-SSFP at very low magnetic field is a a powerful strategy for high-speed MRI at 6.5 mT, and allows us to achieve more than a 100-fold time savings compared to traditional gradient echo imaging in this regime [11].

Balanced sequences are very sensitive to the amount of spin dephasing that occurs between consecutive RF pulses (the pulse repetition time, TR), and typical banding artifacts are expected to appear within a range of ±1/(2*TR) Hz that result from inhomogeneity in the static magnetic field [9]. This sets a strict requirement on the absolute magnetic field homogeneity over the field-of-view (FOV), which for operation at 3 T is a very challenging sub-PPM level. In the ULF regime, however, the fractional homogeneity requirement is three orders of magnitude lower, significantly easing the engineering burden for low-field magnet design. In our system, a TR=22.5 ms is completely immune to banding artifacts for up to 160 ppm inhomogeneity at 6.5 mT. Furthermore, magnetic susceptibility differences are significantly reduced at ULF, preventing off-resonance b-SSFP artifacts. As a result, provided reasonable magnetic field homogeneity, b-SSFP at very low magnetic field alleviates the necessity of ultra-short TRs and provides good image quality over a large FOV without the need for sophisticated ultrafast gradient power amplifiers.

RF Coil design:

The design of inductive detection coils for use in ULF MRI presents a different set of challenges to those present in conventional high-field MRI. In particular, issues of coil resistance and probe bandwidth manifest differently. In conventional MRI, the dominant source of noise is the presence of small currents in the lossy sample (the so-called “body noise” regime) to which a characteristic sample resistance RS is attributed. Both the sample and the coil contribute to Johnson noise but in practice RS is much larger than the coil resistance RC (i.e. RS>>RC), and thus RC can be neglected in SNR calculations. However, at low field, RS becomes much smaller and RC becomes the dominant noise contribution (i.e., the so-called Johnson noise dominated regime). To minimize the coil resistance in a simple design, larger diameter wire or stranded litz wire can be used, but one needs to consider the impact this has on coil bandwidth. Given the maximum imaging gradient strength of ~ 1 mT/m attainable in our 6.5 mT scanner, a 20 cm (head-sized) FOV will span a frequency encode bandwidth of ~ 10 kHz. This sets the minimum bandwidth needed for the detection circuit so as to not significantly convolve the coil response function with the object being imaged. At our Larmor frequency of 276 kHz, this corresponds to a maximum coil Q of ~30. A single channel 30 turn inductive coil for operation at 276 kHz (Fig. 3) was designed and built using 3D printing fused deposition modeling technology and multi-strand litz wire[12]. The hemispheric spiral design results in a very homogeneous magnetic field [13], [14] over the volume of interest that is everywhere orthogonal to our transverse main magnetic field B0, making it suitable for both RF transmit and receive. The number of turns in the coil was chosen to obtain the inductance needed to obtain the desired Q.

In addition to high-efficiency and homogeneity, this coil design is well suited for imaging in a forward deployed trauma setting, as the coil former can be printed in customized sizes for various head sizes or to ensure compatibility with broad classes of head pathology in trauma settings and with access holes so as to not impact workflow or interfere with other critical monitoring devices.

Three-dimensional under-sampled images acquired at 6.5 mT in 6 minutes are shown in Fig 4 for each of the three spatial orientations (axial, coronal, and sagittal). In the brain, the two hemispheres and the cerebellum are distinct, and cortical tissue can be distinguished from white matter. Liquid compartments, here CSF, appear in bright grey and white. In b-SSFP, contrast is related to the ratio of the imaged sample [9]. At high field, liquids and tissue typically have rather different relaxation times but at 6.5 mT their ratio (T₂/T₁) is of order unity resulting in the distinct PD-weighted contrast shown.

As described above, the reduced fractional homogeneity requirement of b-SSFP in the ULF regime is a key feature for neurocritical care. Shown in Fig. 5 are axial images acquired at 3 T and 6.5 mT scanner in a subject who had had undergone surgical brain resection of the right frontoparietal cortex following open head trauma, and a methyl methacrylate prosthesis covered by titanium mesh was placed following resection. Very significant reduction of magnetic susceptibility artifact due to the titanium mesh is observed in the ULF images, and we see that the degree of brain resection is markedly over-estimated at 3 T. ULF MRI is compatible with metallic objects in the head from a safety perspective, and in addition enables imaging in the presence of these materials and visualization of tissue health in regions that would otherwise be obscured by phase aberration leading to lack of image intensity. We speculate that low field MRI will also provide more accurate imaging of penetrating TBI in the presence of bullets and shrapnel.

Our imaging work at 6.5 mT demonstrates the shortest acquisition times and highest SNR per unit time in ULF MRI to date owing to our use of modern sparse sampling strategies and a fully refocused sequence in an optimized electromagnet scanner. These images were acquired without pre-polarization techniques, at a fixed magnetic field and with a simple single channel inductive detector. With an eye towards optimization, we note that for a given spatial resolution, the minimum TR—and consequently the total scan time—is limited by the maximum attainable time-integrated gradient strength. The maximum gradient strength in the LFI is currently ~1 mT·m^-1, resulting in a minimum TR of ~23 ms. Weak gradients especially impact phase encoding in balanced sequences like b-SSFP, as every phase-encode pulse is paired with an opposite polarity rewinding pulse. An increase in gradient strength would allow shorter phase encode pulses, thus decreasing total imaging time while maintaining SNR, provided that image distortion from non-linear magnetic fields that accompany the desired encoding gradient (the so called “concomitant field” artifacts [15]) can be mitigated. At 6.5 mT, an increase in gradient strength in the range of 2-5×, combined with efficient strategies to eliminate concomitant field artifacts [16]-[18], can reasonably be envisioned.

A key challenge in obtaining clinically relevant MRI images at ULF is the ability to acquire T1 and/or T2 relaxation-weighted images, and thereby provide contrast to different types of tissue. Typically, magnetization prepared gradient-echo, and spin-echo sequences are used to obtain relaxation-weighted images, but these types of imaging experiments become prohibitively time consuming at ultra-low magnetic fields where signal averaging and recovery of the longitudinal magnetization are required. To this, we believe that b-SSFP-based “magnetic resonance fingerprinting” (MRF) [19] is an ideal strategy to obtain contrast at ULF, and have shown some very encouraging preliminary results at 6.5 mT [20].

We contend that ULF MRI scanners operating at this expected level of performance could complement traditional MRI by relieving hospital congestion and shortening triage delays. Outside of the radiology suite, mobile ULF scanners might be deployable during military conflicts or during sport events and enable the acquisition of immediate after-trauma knowledge, typically in the case of traumatic brain injuries. Finally, ULF MRI technology may allow resource-poor environments access to MRI systems, without the strict siting requirements and high costs of conventional scanners.