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Magnetic Resonance in Medicine 72:1793–1800 (2014)
Prototype Phantoms for Characterization of Ultralow Field Magnetic Resonance Imaging Michael A. Boss,1* John A. B. Mates,1 Sarah E. Busch,2y Paul SanGiorgio,2z Stephen E. Russek,1 Kai Buckenmaier,2 Kent D. Irwin,1 Hsiao-Mei Cho,1 Gene C. Hilton,1 and John Clarke2 Purpose: Prototype phantoms were designed, constructed, and characterized for the purpose of calibrating ultralow field magnetic resonance imaging (ULF MRI) systems. The phantoms were designed to measure spatial resolution and to quantify sensitivity to systematic variation of proton density and relaxation time, T1. Methods: The phantoms were characterized first with conventional magnetic resonance scanners at 1.5 and 3 T, and subsequently with a prototype ULF MRI scanner between 107 and 128 mT . Results: The ULF system demonstrated a 2-mm spatial resolution and, using T1 measurements, distinguished aqueous solutions of MnCl2 differing by 20 mM [Mn2þ]. Conclusion: The prototype phantoms proved well-matched to ULF MRI applications, and allowed direct comparison of the performance of ULF and clinical systems. Magn Reson Med C 2013 Wiley Periodicals, Inc. 72:1793–1800, 2014. V Key words: ultralow field; ultralow field magnetic resonance imaging; phantom; T1; proton density; spatial resolution; superconducting quantum interference device
INTRODUCTION Standardized, traceable phantoms play an important role in both clinical and research magnetic resonance imaging (MRI) (1), but have not previously been implemented with ultralow field (on the order of 100 mT) MRI (ULF MRI). Phantoms enable characterization, calibration, and 1 Electromagnetics Division, National Institute of Standards and Technology, Boulder, Colorado, USA. 2 Department of Physics, University of California, Berkeley, California, USA.
Grant sponsor: Donaldson Trust; Grant sponsor: National Institutes of Health. † Present address: NASA, Goddard Space Flight Center, Greenbelt, MD 20771. ‡ Present address: Agilent Technologies, Santa Clara, CA 95051. *Correspondence to: Michael A. Boss, Ph.D., National Institute of Standards and Technology, 325 Broadway, MS 818.03, Boulder, CO 80305. E-mail: [email protected] Correction added after online publication 29 January 2014. The units for Bp were corrected to mT and its italicization was corrected in Table 1. The text on page 3, column 1 was updated from “Figure 1c,d” to “Figure 1b,d.” Additional formatting changes were also made in accordance with our Style Guide including hyphenation corrections and figure label corrections, such as “a:” to “(a).” Received 19 July 2013; revised 6 November 2013; accepted 7 November 2013 DOI 10.1002/mrm.25060 Published online 26 November 2013 in Wiley Online Library (wileyonlinelibrary.com). C 2013 Wiley Periodicals, Inc. V
validation of magnetic resonance (MR) scanners, guide their improvement, and facilitate quality assurance. Furthermore, they provide quantitative, reproducible reference data with which to assess imaging techniques. In this article, we describe the design, fabrication, and characterization of prototype phantoms for ULF MRI. We designed and fabricated custom phantoms to characterize ULF MRI systems: a relaxation time (T1) phantom using MnCl2 as a relaxation agent, a proton-density (PD) phantom consisting of binary mixtures of H2O and D2O, and a resolution phantom. These phantoms were designed to mimic the range of T1 values measured in tissues at ULF (2,3), to cover the range of proton densities in the body and to bracket the expected spatial resolution of ULF systems. To our knowledge, previous ULF phantoms have not been designed or maintained as standard references to calibrate PD or relaxation time measurements in ULF MRI systems, as is commonly done in clinical MRI (4). Our phantoms enable characterization of ULF systems with the same metrics used in clinical systems to assess system performance. In addition, the phantoms are compatible with clinical systems, allowing direct comparison between ULF and clinical scanners. We emphasized characterization of T1 measurement because an inherent advantage of ULF MRI over clinical MRI is the enhanced T1 contrast between tissues. This enhancement arises from the fact that the spectral density of fluctuations inducing relaxation is higher at ULFs B0 1=ðgtC Þ than at clinical fields (5,6). Here, tC is the tissuedependent magnetic coherence time and c is the proton gyromagnetic ratio. For example, prostate tumors, which exhibit similar T1 values to healthy prostate tissue at clinical fields, exhibit a significantly lower T1 at ULF compared to nonpathologic tissue (2). This T1 contrast at ULF may provide a sensitive, specific, and noninvasive means of detecting pathologic tissues compared to conventional techniques at clinical field strengths of 1.5 or 3 T. The price of enhanced ULF contrast is lower signal-tonoise ratio (SNR) arising from two factors (7,8). First, the proton magnetization is reduced; this can be overcome—at least to some extent—by prepolarization of the protons (9– 11). Second, by Faraday’s Law, the voltage induced into the receiver coil scales as the rate of change of applied magnetic flux, and thus with frequency (12,13). This drawback can be eliminated by detecting the signal with a superconducting quantum interference device coupled inductively to an untuned input circuit (14–28). A further application
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FIG. 1. Configurations and photographs of phantoms. (a) The T1 plate and (b) resolution-array plate. Photographs show assembled phantoms for (c) the T1 map and (d) the PD map and resolution array. Polypropylene cylinders 19-mm tall are placed in the polyphenylene sulfide plates and hold approximately 1.5 mL of (c) Mn2þ dilutions or (d) D2O dilutions. The resolution array (b, d) consists of four arrays of holes, with diameters of 0.5, 1, 2, and 3 mm, spaced center-to-center by 1, 2, 4, and 6 mm, respectively. For imaging at clinical fields, the phantoms are filled with water to avoid susceptibility artifacts.
of standardized phantoms would be to compare the performance of ULF scanners at different institutions. In this article, we present measurements of our prototype phantoms at both microtesla fields and clinical field strengths. METHODS Phantoms The phantoms consist of short acrylic cylinders, similar in shape and size to hockey pucks, containing multiple polypropylene cylinders inserted into polyphenylene sulfide plates in a single plane (Fig. 1). We chose acrylic for its ready availability and ease of machining, and polyphenylene sulfide for its structural rigidity, low water absorption, and resistance to elevated temperatures and chemicals. There are two different polyphenylene sulfide plate designs: one holds 10 T1 or PD cylinders (Fig. 1a); the other holds six cylinders and includes carefully machined holes of precise dimensions to serve as a resolution array (Fig. 1b). Although the structure of these phantoms could cause susceptibility artifacts at clinical field strengths, the effects of susceptibility mismatch are greatly reduced at ULF, and image distortion is negligible (28). To minimize susceptibility artifacts in clinical strength images, each phantom has a fill port to immerse its contents in water. T1 Array The T1 phantom enables us to vary T1 controllably by changing the concentration of MnCl2 in water. The relax-
ation rates of protons in water in the presence of Mn2þ ions at both ultralow and clinical fields are well-studied with field-cycling nuclear magnetic resonance (NMR) relaxometry (29), and show that 1/T1 scales linearly with concentration. MnCl2 dilutions are easy to prepare in exact concentrations, are less toxic than other commonly used relaxation agents such as NiCl2, are stable over long periods, and have a T1 field-dependence that allows the same solutions to be used at both ULF and clinical field strengths. When well-sealed in the polypropylene cylinders, the solutions do not leak out or evaporate, and thus provide long-term phantoms that can be used for intrascanner and interscanner performance assessment. A stock solution of MnCl2 was mixed at a concentration of 1:399660:0033 mg of Mn per gram of aqueous solution, corresponding to 25:47660:060 mol/m3, or mM [Mn2þ]. Here and elsewhere in this manuscript, the measurement uncertainty is the standard deviation of the measurement. The concentration was verified by inductively coupled plasma mass spectroscopy. The stock was diluted with deionized water to 10 different concentrations by mass: 0.080, 0.100, 0.120, 0.170, 0.190, 0.210, 0.230, 0.280, 0.300, and 0.320 mM, to within 60:002 mM. These concentrations were chosen to encompass the range of measured T1 in vivo at ULF (50–200 ms, or 1/T1 ¼ 5 20 s1, and to determine how accurately concentration could be resolved at both ultralow and clinical fields. Aliquots of these solutions were pipetted into the small polypropylene cylinders (Fig. 1a,c) for MRI. The unused solutions were retained in standard 15-mL polypropylene centrifuge tubes for further NMR and MRI analysis.
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FIG. 2. Pulse sequence for 2D imaging and T1 mapping (tev varied). The static field B0 is applied continuously. The prepolarizing field Bp is turned on for a period tp and switched off adiabatically, allowing the spins to relax for a time tev in B0, before a standard spin-echo imaging sequence is applied. The frequency encoding field gradients are applied continuously, whereas the pulsed phase encoding gradients are applied only between the 90 and 180 pulses. The superconducting quantum interference device gradiometer acquires the echo. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
PD Array and Resolution Array The PD phantom consists of dilutions of D2O in H2O, in concentrations by mass of 5, 10, 20, 40, 80, and 100 percent H2O, pipetted into the six cylinders shown in Figure 1b,d. These proton densities are representative of those observed in vivo. The resolution array (Fig. 1d) was designed to determine the resolving capability of ULF MRI, typically on the order of 1 mm (16). It consists of four arrays of holes, of diameter 0.5, 1.0, 2.0, and 3.0 mm, spaced center-tocenter by 1, 2, 4, and 6 mm, respectively. For measurement, the phantom was immersed in Mn-doped water, which provided the proton signal, while simultaneously decreasing T1 to enable shorter repetition times (TRs). Clinical Field Strength NMR and MRI We measured relaxation times with a 1.5T NMR spectrometer to reflect the most common clinical MRI field strength. To prepare the NMR samples, we pipetted 70 mL of MnCl2 solution into a 100-mm long, 2-mm outer diameter borosilicate glass capillary, previously sealed at one end, leaving a headspace of 78 mm above the liquid. We degassed the solution, backfilled the headspace with
helium, and flame-sealed it approximately 10 mm from the end of the capillary. We centered each sample in a standard 5-mm diameter NMR tube using polytetrafluoroethylene spacers and loaded it into the spectrometer. We characterized each sample with an inversion-recovery pulse sequence with 20 logarithmically spaced inversion recovery times (TI). The data were averaged over three such sequences, and fitted to determine T1. We imaged the phantoms with a clinical 3T MRI system using an eight-channel head coil. Images consisted of 2-mm slices, a 100 100-mm2 field of view, and a 256 256 matrix size. We acquired T1 data with a fastspin echo inversion recovery sequence, with an echo time (TE) and TR of 8 and 4000 ms, respectively. PD and resolution-array data were obtained with a preclinical MRI scanner operating at a field strength of 1.5 T, using a spin-echo sequence, with a TE/TR of 16/3500 ms. We manually selected regions-of-interest (ROIs) in the resulting images, averaged the signals within them, and fitted the data to determine T1. A similar ROI selection process was used for the PD images. ULF NMR and MRI We operated the ULF MRI system using a superconducting quantum interference device, coupled to a superconducting gradiometer, to detect the signal (16,30). We adjusted the imaging field, B0, from 107 to 128 mT to find the NMR frequency at which the environmental noise, primarily harmonics of 60 Hz, was minimized. The proton spins in the sample were prepolarized in a pulsed magnetic field Bp ranging from 80 to 100 mT, which was turned off adiabatically after a duration tp , before initiating the imaging sequence (9,31), so that the initial magnetization, M0, is proportional to Bp . We performed three types of ULF measurements: T1 using NMR, T1 maps using two-dimensional (2D) MRI, and 2D PD images. All three measurements were based on the spin-warp pulse sequence shown in Figure 2 (32), which we term “prepolarization evolution”: following their prepolarization, the spins decay for a duration tev toward their intrinsic polarization in the static field, B0, which is three orders of magnitude smaller than Bp . A 90 excitation pulse is followed by a phase encoding gradient (GPE ) pulse and a 180 refocusing pulse after a duration s to generate a spin echo. The period between the 90 and the echo center is the TE; as the echo forms, signal is acquired for a period of time, Tacq . During the experiment, a frequency encoding gradient (GFE ) is continuously applied. Typical imaging parameters are listed in
Table 1 ULF MRI Parameter Values Parameter
T1-NMR
T1-MRI
Proton density
Resolution
B0 (mT ) Bp (mT) GFE (mT/m) GPE (mT/m) tp (ms) tev (ms) TE (ms) Tacq (ms)
123–126 80–100 90 140 to þ140 300 15–610 72 102
124 80–100 90 140 to þ140 300 10–610 72 102
107 80–100 90 140 to þ140 400 7 76 102
128 80–100 108 200 to þ200 400 7 74 102
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FIG. 3. Representative NMR data, for 0.19 mM [Mn2þ]. (a) Inversion recovery at 1.5 T and (b) prepolarization evolution at 124 mT . Signal is in arbitrary units such that the standard deviation for each data point is 61. Solid lines are fits to the data. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
Table 1. For PD images, tev is chosen to be as short as possible (10 ms), and a single image is acquired. For T1 image sets, a 2D image is taken at each of several values of tev . We acquire 20 averages at 18 different values of tev , requiring a total measurement time of 15–20 min per tev . For T1 measurements of homogeneous samples with ULF NMR, the phase encoding gradient is turned off, and we repeat the sequence with different values of tev , fitting the resulting data to an exponential decay. To acquire the high-resolution image of the spatial resolution phantom, we use a higher frequency encoding gradient GFE (108 mT =m) and higher phase encoding gradients GPE (200 to þ200 mT =m) than in typical imaging. The directions of the fields and gradients are: B0 along the z-axis (horizontally oriented), Bp along the x-axis (vertically oriented), GFE along z and GPE along y. RESULTS T1 from NMR We used NMR to measure T1 of the retained MnCl2 dilutions in the 15-mL centrifuge tubes. At 1.5 T, as the inversion time, TI, is increased, the signal recovered from a negative value to its positive thermal equilibrium value in a characteristic time T1 (Fig. 3a). At ULF (123– 126 mT ), as the evolution time tev was increased the signal decayed to the equilibrium magnetization (Fig. 3b). The ULF measurement was performed twice for each of the 10 dilutions used in the T1 phantom. The relaxation rates, 1/T1, scaled linearly with Mn2þ concentration at both field strengths (Fig. 4). At 1.5 T, the slope of the best linear fit, the relaxivity, was 7:2060:06 s1 mM1. At ULF, the relaxivity was 59:160:3 s1 mM1, an order of magnitude larger.
as a set they exhibit linear relaxivities of 6:660:2 s1 mM1 at 3 T, and 60:460:8 s1 mM1 at 124 mT . The MR images acquired at 3 T and ULF were recorded as magnitude data, which biases the ROI intensities at low SNR. Consequently, the ULF relaxation times (and by extension, the relaxivity) were obtained by fitting the meansquared value of the ROI to an exponential decay with a finite offset (the image noise floor). This fitting procedure minimizes bias in the measured values due to low SNR: for the approximately 150 pixels used in the ULF ROIs, this results in a relaxation rate bias of less than 0.1%. PD Results The PD array was imaged at 1.5 T and 107 mT to determine the signal intensity as a function of proton concentration. Initial results indicated that the ULF signal intensities did not scale with the 1.5T signal intensities as expected. Rather, the signal intensities were dependent on the orientation of the phantom. The six PD vials were replaced with six identical vials of 5 mM NiCl2, each containing 1:00060:005 mL, and imaged at ULF. Circular ROIs drawn inside the vials in the resulting image revealed a spatial dependence of the signal
T1 from MRI The T1 phantom was imaged at 3 T with an inversion recovery pulse sequence (Fig. 5a), with a total acquisition time of approximately 1 h. A prepolarization decay pulse sequence (Figs. 2 and 5b) was used to acquire an ULF MRI, 2D T1-map of the same phantom in approximately 5 h. We calculated T1 on a pixel-by-pixel basis from the images acquired at 3 T and 124 mT to generate T1 maps (Fig. 6a,b). There is a clear distinction between the 0.10, 0.20, and 0.30 mM [Mn2þ] series in both sets of images. Furthermore, when the relaxation times of circular ROIs are compared (Fig. 6c,d), each dilution is distinguishable, and
FIG. 4. Relaxation rates and residuals as a function of Mn2þ concentration determined by NMR. (a) 1/T1 at 1.5 T (triangles) and ULF (123–126 mT, circles). Standard deviations of the data are less than 0.1 and 1%, respectively. (b) Residuals from the fits yield a relaxivity uncertainty of 60:06 s1 mM1 at 1.5 T and 60:3 s1 mM1 at ULF. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 5. Representative MRI data for a single pixel of the 0.19 mM [Mn2þ] solution. (a) Inversion recovery at 3 T and (b) prepolarization evolution at 124 mT . Signal is in arbitrary units such that the standard deviation for each point is 61. Solid lines are fits to the data. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
intensity (Fig. 7). The PD vials were restored to their original positions, and the phantom was imaged in the same orientation as in the NiCl2 experiment. The signal magnitude of each PD vial, averaged over circular ROIs, was subsequently corrected using the results of the NiCl2 data. The corrected results showed a linear dependence on H2O concentration, and matched the 1.5T data remarkably well (Fig. 8).
Resolution-Array Results Images from the resolution array obtained with 3T MRI (Fig. 9a) and 128 mT MRI (Fig. 9b) were acquired in 30 s and 36 min, respectively. The 3T images clearly resolved three resolution arrays (1, 2, and 3-mm holes), whereas the 128 mT image clearly resolved the 2 and 3-mm holes. We convolved a perfect digital representation of the
FIG. 6. Grayscale T1 maps at (a) 3 T and (b) 124 mT . Mn2þ concentrations increase from top to bottom and from left to right. White circles show representative ROIs. Measured relaxation times within these ROIs yield relaxivity at (c) 3 T and (d) 124 mT . Solid lines show linear fits, and subplots below show residuals. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 7. Identical vials of NiCl2 imaged with the ULF scanner. (a) Characterization of the spatial dependence of the signal intensity. (b) Variation of intensity as a function of vial position. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
resolution array with a series of Gaussian functions of increasing width, and compared the resulting images with our MRI-acquired image. We then fit the data to obtain the width in y and z simultaneously using a least squares fitting procedure. By fitting the images to this Gaussian degradation, we found 0.8-mm resolution at 3 T, and 1.8 and 2.6-mm resolution at 128 mT in the frequency encoding and phase encoding directions, respectively. DISCUSSION Our ULF NMR measurements demonstrate that the MnCl2 dilutions are appropriate for ULF, span a range of clinically relevant T1 values (50–200 ms) and verify the T1-resolving power of the ULF system. The ULF relaxivity is an order of magnitude greater than that at 1.5 or 3 T. The 1.5T and ULF relaxivity results agree with published
values for the relaxation times of protons in the presence of manganese ions (29). Furthermore, fits to both the 1.5T and ULF NMR data intersect zero concentration at 1=T1 ¼ 0:3660:01 and 0:3660:04 s1, respectively, consistent with the relaxation time of water at both clinical field strengths and ULF. The residuals from the fit showed no systematic error and 61% error bars, readily distinguishing dilutions differing by 0.02 mM [Mn2þ]. The ULF and clinical field relaxation times are also comparable to those of tissues in vivo, making the phantom suite appropriate for T1 standards for both field values. The NMR and MRI measurements of the ULF relaxivity agree to within 2.2%. At 1.5 T, the signal intensity scales with increasing PD. The ULF PD results closely match the results obtained at 1.5 T. However, it was necessary to map out the spatial sensitivity with identical vials of NiCl2 to
FIG. 8. Signal intensity at 107 mT vs. signal intensity at 1.5 T. The blue circles represent the raw, uncorrected data acquired at ULF; red circles are corrected using the results from the NiCl2 phantom. The corrected ULF data are in agreement with the 1.5 T data to within less than 61%. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 9. Grayscale MR images and slice profiles of the resolution inset phantom. (a, c) 3 T and (b, d) 128 mT . Acquisition times were 30 s and 36 min, respectively. Slice profiles through the 3, 2, and 1-mm diameter holes indicated in (a) and (b) by dotted lines are consistent with measured image resolutions of 0.8 mm at 3 T and 1.8 mm at 128 mT . The images have slightly different configurations: the imaging of the resolution array at 3 T was performed simultaneously with the PD measurement, while the 128 mT image was acquired with only a single cylinder from the PD array, acting as an alignment key. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
correct the ULF PD intensity data. This use of the phantom revealed that the signal intensity in the ULF scanner was somewhat spatially dependent. This angular inhomogeneity was subsequently traced to a misalignment of the gradiometer. Although the inhomogeneity does not directly affect T1 or resolution results, it is important to correct for it in deriving PD data from ULF images. The sensitivity map, combined with the corrected PD results, demonstrates that the ULF scanner is capable of measuring PD as accurately as an MRI scanner operating at clinical field strengths. Our measurements of the resolution array at 128 mT agreed with the expected system spatial resolution, set by limitations on gradient strength and SNR to approximately 2 mm. The 3 T scanner was able to resolve the 1-mm features; for the given measurement parameters, the ULF resolution obtained is comparable to those obtained by clinical scanners, albeit with a longer scan time. The feature sizes on the phantom are appropriate for evaluating future improvements of ULF MRI performance.
CONCLUSIONS Traceable phantoms allow us to establish performance metrics for MRI scanners and make meaningful comparisons among them. Phantoms facilitate the improvement of scanner capabilities and the advancement of MRI techniques. Traceable standardized phantoms have been developed for these purposes in clinical MRI; such phantoms have not previously been implemented for ULF MRI. There are many advantages of ULF scanners, such as greater T1 contrast, lack of susceptibility artifacts, low cost, and light weight, which compensate for low SNR. Given the difference in longitudinal relaxation times between cancerous and healthy tissue at ULF, there is a clear potential clinical application for ULF MRI. ULF MRI may provide a simpler, more versatile, and noninvasive method of imaging for diagnosis and treatment monitoring of certain cancers, and potentially other pathologies. It is desirable to characterize ULF system
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performance quantitatively by use of a standardized phantom. The prototype phantoms proved to be well-matched to superconducting quantum interference device-based ULF MRI systems and can already be used to characterize their T1 sensitivity and spatial resolution. Their use facilitates comparison between different scanners at both ULF and clinical field strengths through common metrics such as SNR, imaging time, resolution, and spatial homogeneity of the signal detection sensitivity, and can guide scanner development. The next generation of phantoms will incorporate design changes based on lessons learned from the prototypes, improving stability, facilitating maintenance, providing traceability, and allowing for standardized comparisons of system performance over time and across scanners. ACKNOWLEDGMENTS The authors gratefully acknowledge the assistance of Dr. Hui Dong in identifying the source of the spatially inhomogeneous response of the ULF MRI system. Imaging at 3 T was performed at the University of Colorado Health Sciences Center (Aurora, CO). The 1.5T NMR and MRI data were acquired on a vertical bore system and a horizontal bore scanner, respectively, at the National Institute of Standards and Technology (Boulder, CO). Ultralow field NMR and MRI experiments were conducted at the University of California, Berkeley. Work at UC Berkeley was funded by the Donaldson Trust and the National Institutes of Health. K. B. gratefully acknowledges receipt of a fellowship from the Deutsche Forschungsgemeinschaft. REFERENCES 1. Gunter JL, Bernstein MA, Borowski BJ, Ward CP, Britson PJ, Felmlee JP, Schuff N, Weiner M, Jack CR. Measurement of MRI scanner performance with the ADNI phantom. Med Phys 2009;36:2193–2205. 2. Busch S, Hatridge M, M€ oßle M, Myers W, Wong T, M€ uck M, Chew K, Kuchinsky K, Simko J, Clarke J. Measurements of T1-relaxation in ex vivo prostate tissue at 132 mT. Magn Reson Med 2012;67:1138–1145. 3. Zotev VS, Matlashov AN, Savukov IM, Owens T, Volegov PL, Gomez JJ, Espy MA. SQUID-based microtesla MRI for in vivo relaxometry of the human brain. IEEE Trans Appl Supercond 2009;19:823–826. 4. Russek SE, Boss M, Jackson EF, Jennings DL, Evelhock JL, Gunter JL, Sorensen AG. Characterization of NIST/ISMRM MRI system phantom. In Proceedings of the 20th Annual Meeting of ISMRM, Melbourne, Australia, 2012. Abstract 2456. 5. Koenig SH, Brown RD. Relaxometry of solvent and tissue protons: diamagnetic contributions. Magnetic resonance imaging. Physical principles and instumentation. Vol. 2. Philadelphia: Saunders; 1987. 6. Lee SK, M€ oßle M, Myers W, Kelso N, Trabesinger AH, Pines A, Clarke J. SQUID-detected MRI at 132 mT with T1-weighted contrast established at 10 mT–300 mT. Magn Reson Med 2005;53:9–14. 7. Myers W, Slichter D, Hatridge M, Busch S, M€ oßle M, McDermott R, Trabesinger A, Clarke J. Calculated signal-to-noise ratio of MRI detected with SQUIDs and Faraday detectors in fields from 10 mT to 1.5 T. J Magn Reson 2007;186:182–192. 8. M€ oßle M, Myers WR, Lee SK, Kelso N, Hatridge M, Pines A, Clarke J. SQUID-detected in vivo MRI at microtesla magnetic fields. IEEE Trans Appl Supercond 2005;15:757–760. 9. Packard M, Varian R. Free nuclear induction in the earth’s magnetic field. Phys Rev 1954;93:941. 10. Qiu L, Zhang Y, Krause HJ, Braginski AI, Burghoff M, Trahms L. Nuclear magnetic resonance in the earth’s magnetic field using a nitrogen-cooled superconducting quantum interference device. Appl Phys Lett 2007;91:072505.
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Magnetic Resonance in Medicine 72:1793–1800 (2014)
Prototype Phantoms for Characterization of Ultralow Field Magnetic Resonance Imaging Michael A. Boss,1* John A. B. Mates,1 Sarah E. Busch,2y Paul SanGiorgio,2z Stephen E. Russek,1 Kai Buckenmaier,2 Kent D. Irwin,1 Hsiao-Mei Cho,1 Gene C. Hilton,1 and John Clarke2 Purpose: Prototype phantoms were designed, constructed, and characterized for the purpose of calibrating ultralow field magnetic resonance imaging (ULF MRI) systems. The phantoms were designed to measure spatial resolution and to quantify sensitivity to systematic variation of proton density and relaxation time, T1. Methods: The phantoms were characterized first with conventional magnetic resonance scanners at 1.5 and 3 T, and subsequently with a prototype ULF MRI scanner between 107 and 128 mT . Results: The ULF system demonstrated a 2-mm spatial resolution and, using T1 measurements, distinguished aqueous solutions of MnCl2 differing by 20 mM [Mn2þ]. Conclusion: The prototype phantoms proved well-matched to ULF MRI applications, and allowed direct comparison of the performance of ULF and clinical systems. Magn Reson Med C 2013 Wiley Periodicals, Inc. 72:1793–1800, 2014. V Key words: ultralow field; ultralow field magnetic resonance imaging; phantom; T1; proton density; spatial resolution; superconducting quantum interference device
INTRODUCTION Standardized, traceable phantoms play an important role in both clinical and research magnetic resonance imaging (MRI) (1), but have not previously been implemented with ultralow field (on the order of 100 mT) MRI (ULF MRI). Phantoms enable characterization, calibration, and 1 Electromagnetics Division, National Institute of Standards and Technology, Boulder, Colorado, USA. 2 Department of Physics, University of California, Berkeley, California, USA.
Grant sponsor: Donaldson Trust; Grant sponsor: National Institutes of Health. † Present address: NASA, Goddard Space Flight Center, Greenbelt, MD 20771. ‡ Present address: Agilent Technologies, Santa Clara, CA 95051. *Correspondence to: Michael A. Boss, Ph.D., National Institute of Standards and Technology, 325 Broadway, MS 818.03, Boulder, CO 80305. E-mail: [email protected] Correction added after online publication 29 January 2014. The units for Bp were corrected to mT and its italicization was corrected in Table 1. The text on page 3, column 1 was updated from “Figure 1c,d” to “Figure 1b,d.” Additional formatting changes were also made in accordance with our Style Guide including hyphenation corrections and figure label corrections, such as “a:” to “(a).” Received 19 July 2013; revised 6 November 2013; accepted 7 November 2013 DOI 10.1002/mrm.25060 Published online 26 November 2013 in Wiley Online Library (wileyonlinelibrary.com). C 2013 Wiley Periodicals, Inc. V
validation of magnetic resonance (MR) scanners, guide their improvement, and facilitate quality assurance. Furthermore, they provide quantitative, reproducible reference data with which to assess imaging techniques. In this article, we describe the design, fabrication, and characterization of prototype phantoms for ULF MRI. We designed and fabricated custom phantoms to characterize ULF MRI systems: a relaxation time (T1) phantom using MnCl2 as a relaxation agent, a proton-density (PD) phantom consisting of binary mixtures of H2O and D2O, and a resolution phantom. These phantoms were designed to mimic the range of T1 values measured in tissues at ULF (2,3), to cover the range of proton densities in the body and to bracket the expected spatial resolution of ULF systems. To our knowledge, previous ULF phantoms have not been designed or maintained as standard references to calibrate PD or relaxation time measurements in ULF MRI systems, as is commonly done in clinical MRI (4). Our phantoms enable characterization of ULF systems with the same metrics used in clinical systems to assess system performance. In addition, the phantoms are compatible with clinical systems, allowing direct comparison between ULF and clinical scanners. We emphasized characterization of T1 measurement because an inherent advantage of ULF MRI over clinical MRI is the enhanced T1 contrast between tissues. This enhancement arises from the fact that the spectral density of fluctuations inducing relaxation is higher at ULFs B0 1=ðgtC Þ than at clinical fields (5,6). Here, tC is the tissuedependent magnetic coherence time and c is the proton gyromagnetic ratio. For example, prostate tumors, which exhibit similar T1 values to healthy prostate tissue at clinical fields, exhibit a significantly lower T1 at ULF compared to nonpathologic tissue (2). This T1 contrast at ULF may provide a sensitive, specific, and noninvasive means of detecting pathologic tissues compared to conventional techniques at clinical field strengths of 1.5 or 3 T. The price of enhanced ULF contrast is lower signal-tonoise ratio (SNR) arising from two factors (7,8). First, the proton magnetization is reduced; this can be overcome—at least to some extent—by prepolarization of the protons (9– 11). Second, by Faraday’s Law, the voltage induced into the receiver coil scales as the rate of change of applied magnetic flux, and thus with frequency (12,13). This drawback can be eliminated by detecting the signal with a superconducting quantum interference device coupled inductively to an untuned input circuit (14–28). A further application
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FIG. 1. Configurations and photographs of phantoms. (a) The T1 plate and (b) resolution-array plate. Photographs show assembled phantoms for (c) the T1 map and (d) the PD map and resolution array. Polypropylene cylinders 19-mm tall are placed in the polyphenylene sulfide plates and hold approximately 1.5 mL of (c) Mn2þ dilutions or (d) D2O dilutions. The resolution array (b, d) consists of four arrays of holes, with diameters of 0.5, 1, 2, and 3 mm, spaced center-to-center by 1, 2, 4, and 6 mm, respectively. For imaging at clinical fields, the phantoms are filled with water to avoid susceptibility artifacts.
of standardized phantoms would be to compare the performance of ULF scanners at different institutions. In this article, we present measurements of our prototype phantoms at both microtesla fields and clinical field strengths. METHODS Phantoms The phantoms consist of short acrylic cylinders, similar in shape and size to hockey pucks, containing multiple polypropylene cylinders inserted into polyphenylene sulfide plates in a single plane (Fig. 1). We chose acrylic for its ready availability and ease of machining, and polyphenylene sulfide for its structural rigidity, low water absorption, and resistance to elevated temperatures and chemicals. There are two different polyphenylene sulfide plate designs: one holds 10 T1 or PD cylinders (Fig. 1a); the other holds six cylinders and includes carefully machined holes of precise dimensions to serve as a resolution array (Fig. 1b). Although the structure of these phantoms could cause susceptibility artifacts at clinical field strengths, the effects of susceptibility mismatch are greatly reduced at ULF, and image distortion is negligible (28). To minimize susceptibility artifacts in clinical strength images, each phantom has a fill port to immerse its contents in water. T1 Array The T1 phantom enables us to vary T1 controllably by changing the concentration of MnCl2 in water. The relax-
ation rates of protons in water in the presence of Mn2þ ions at both ultralow and clinical fields are well-studied with field-cycling nuclear magnetic resonance (NMR) relaxometry (29), and show that 1/T1 scales linearly with concentration. MnCl2 dilutions are easy to prepare in exact concentrations, are less toxic than other commonly used relaxation agents such as NiCl2, are stable over long periods, and have a T1 field-dependence that allows the same solutions to be used at both ULF and clinical field strengths. When well-sealed in the polypropylene cylinders, the solutions do not leak out or evaporate, and thus provide long-term phantoms that can be used for intrascanner and interscanner performance assessment. A stock solution of MnCl2 was mixed at a concentration of 1:399660:0033 mg of Mn per gram of aqueous solution, corresponding to 25:47660:060 mol/m3, or mM [Mn2þ]. Here and elsewhere in this manuscript, the measurement uncertainty is the standard deviation of the measurement. The concentration was verified by inductively coupled plasma mass spectroscopy. The stock was diluted with deionized water to 10 different concentrations by mass: 0.080, 0.100, 0.120, 0.170, 0.190, 0.210, 0.230, 0.280, 0.300, and 0.320 mM, to within 60:002 mM. These concentrations were chosen to encompass the range of measured T1 in vivo at ULF (50–200 ms, or 1/T1 ¼ 5 20 s1, and to determine how accurately concentration could be resolved at both ultralow and clinical fields. Aliquots of these solutions were pipetted into the small polypropylene cylinders (Fig. 1a,c) for MRI. The unused solutions were retained in standard 15-mL polypropylene centrifuge tubes for further NMR and MRI analysis.
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FIG. 2. Pulse sequence for 2D imaging and T1 mapping (tev varied). The static field B0 is applied continuously. The prepolarizing field Bp is turned on for a period tp and switched off adiabatically, allowing the spins to relax for a time tev in B0, before a standard spin-echo imaging sequence is applied. The frequency encoding field gradients are applied continuously, whereas the pulsed phase encoding gradients are applied only between the 90 and 180 pulses. The superconducting quantum interference device gradiometer acquires the echo. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
PD Array and Resolution Array The PD phantom consists of dilutions of D2O in H2O, in concentrations by mass of 5, 10, 20, 40, 80, and 100 percent H2O, pipetted into the six cylinders shown in Figure 1b,d. These proton densities are representative of those observed in vivo. The resolution array (Fig. 1d) was designed to determine the resolving capability of ULF MRI, typically on the order of 1 mm (16). It consists of four arrays of holes, of diameter 0.5, 1.0, 2.0, and 3.0 mm, spaced center-tocenter by 1, 2, 4, and 6 mm, respectively. For measurement, the phantom was immersed in Mn-doped water, which provided the proton signal, while simultaneously decreasing T1 to enable shorter repetition times (TRs). Clinical Field Strength NMR and MRI We measured relaxation times with a 1.5T NMR spectrometer to reflect the most common clinical MRI field strength. To prepare the NMR samples, we pipetted 70 mL of MnCl2 solution into a 100-mm long, 2-mm outer diameter borosilicate glass capillary, previously sealed at one end, leaving a headspace of 78 mm above the liquid. We degassed the solution, backfilled the headspace with
helium, and flame-sealed it approximately 10 mm from the end of the capillary. We centered each sample in a standard 5-mm diameter NMR tube using polytetrafluoroethylene spacers and loaded it into the spectrometer. We characterized each sample with an inversion-recovery pulse sequence with 20 logarithmically spaced inversion recovery times (TI). The data were averaged over three such sequences, and fitted to determine T1. We imaged the phantoms with a clinical 3T MRI system using an eight-channel head coil. Images consisted of 2-mm slices, a 100 100-mm2 field of view, and a 256 256 matrix size. We acquired T1 data with a fastspin echo inversion recovery sequence, with an echo time (TE) and TR of 8 and 4000 ms, respectively. PD and resolution-array data were obtained with a preclinical MRI scanner operating at a field strength of 1.5 T, using a spin-echo sequence, with a TE/TR of 16/3500 ms. We manually selected regions-of-interest (ROIs) in the resulting images, averaged the signals within them, and fitted the data to determine T1. A similar ROI selection process was used for the PD images. ULF NMR and MRI We operated the ULF MRI system using a superconducting quantum interference device, coupled to a superconducting gradiometer, to detect the signal (16,30). We adjusted the imaging field, B0, from 107 to 128 mT to find the NMR frequency at which the environmental noise, primarily harmonics of 60 Hz, was minimized. The proton spins in the sample were prepolarized in a pulsed magnetic field Bp ranging from 80 to 100 mT, which was turned off adiabatically after a duration tp , before initiating the imaging sequence (9,31), so that the initial magnetization, M0, is proportional to Bp . We performed three types of ULF measurements: T1 using NMR, T1 maps using two-dimensional (2D) MRI, and 2D PD images. All three measurements were based on the spin-warp pulse sequence shown in Figure 2 (32), which we term “prepolarization evolution”: following their prepolarization, the spins decay for a duration tev toward their intrinsic polarization in the static field, B0, which is three orders of magnitude smaller than Bp . A 90 excitation pulse is followed by a phase encoding gradient (GPE ) pulse and a 180 refocusing pulse after a duration s to generate a spin echo. The period between the 90 and the echo center is the TE; as the echo forms, signal is acquired for a period of time, Tacq . During the experiment, a frequency encoding gradient (GFE ) is continuously applied. Typical imaging parameters are listed in
Table 1 ULF MRI Parameter Values Parameter
T1-NMR
T1-MRI
Proton density
Resolution
B0 (mT ) Bp (mT) GFE (mT/m) GPE (mT/m) tp (ms) tev (ms) TE (ms) Tacq (ms)
123–126 80–100 90 140 to þ140 300 15–610 72 102
124 80–100 90 140 to þ140 300 10–610 72 102
107 80–100 90 140 to þ140 400 7 76 102
128 80–100 108 200 to þ200 400 7 74 102
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FIG. 3. Representative NMR data, for 0.19 mM [Mn2þ]. (a) Inversion recovery at 1.5 T and (b) prepolarization evolution at 124 mT . Signal is in arbitrary units such that the standard deviation for each data point is 61. Solid lines are fits to the data. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
Table 1. For PD images, tev is chosen to be as short as possible (10 ms), and a single image is acquired. For T1 image sets, a 2D image is taken at each of several values of tev . We acquire 20 averages at 18 different values of tev , requiring a total measurement time of 15–20 min per tev . For T1 measurements of homogeneous samples with ULF NMR, the phase encoding gradient is turned off, and we repeat the sequence with different values of tev , fitting the resulting data to an exponential decay. To acquire the high-resolution image of the spatial resolution phantom, we use a higher frequency encoding gradient GFE (108 mT =m) and higher phase encoding gradients GPE (200 to þ200 mT =m) than in typical imaging. The directions of the fields and gradients are: B0 along the z-axis (horizontally oriented), Bp along the x-axis (vertically oriented), GFE along z and GPE along y. RESULTS T1 from NMR We used NMR to measure T1 of the retained MnCl2 dilutions in the 15-mL centrifuge tubes. At 1.5 T, as the inversion time, TI, is increased, the signal recovered from a negative value to its positive thermal equilibrium value in a characteristic time T1 (Fig. 3a). At ULF (123– 126 mT ), as the evolution time tev was increased the signal decayed to the equilibrium magnetization (Fig. 3b). The ULF measurement was performed twice for each of the 10 dilutions used in the T1 phantom. The relaxation rates, 1/T1, scaled linearly with Mn2þ concentration at both field strengths (Fig. 4). At 1.5 T, the slope of the best linear fit, the relaxivity, was 7:2060:06 s1 mM1. At ULF, the relaxivity was 59:160:3 s1 mM1, an order of magnitude larger.
as a set they exhibit linear relaxivities of 6:660:2 s1 mM1 at 3 T, and 60:460:8 s1 mM1 at 124 mT . The MR images acquired at 3 T and ULF were recorded as magnitude data, which biases the ROI intensities at low SNR. Consequently, the ULF relaxation times (and by extension, the relaxivity) were obtained by fitting the meansquared value of the ROI to an exponential decay with a finite offset (the image noise floor). This fitting procedure minimizes bias in the measured values due to low SNR: for the approximately 150 pixels used in the ULF ROIs, this results in a relaxation rate bias of less than 0.1%. PD Results The PD array was imaged at 1.5 T and 107 mT to determine the signal intensity as a function of proton concentration. Initial results indicated that the ULF signal intensities did not scale with the 1.5T signal intensities as expected. Rather, the signal intensities were dependent on the orientation of the phantom. The six PD vials were replaced with six identical vials of 5 mM NiCl2, each containing 1:00060:005 mL, and imaged at ULF. Circular ROIs drawn inside the vials in the resulting image revealed a spatial dependence of the signal
T1 from MRI The T1 phantom was imaged at 3 T with an inversion recovery pulse sequence (Fig. 5a), with a total acquisition time of approximately 1 h. A prepolarization decay pulse sequence (Figs. 2 and 5b) was used to acquire an ULF MRI, 2D T1-map of the same phantom in approximately 5 h. We calculated T1 on a pixel-by-pixel basis from the images acquired at 3 T and 124 mT to generate T1 maps (Fig. 6a,b). There is a clear distinction between the 0.10, 0.20, and 0.30 mM [Mn2þ] series in both sets of images. Furthermore, when the relaxation times of circular ROIs are compared (Fig. 6c,d), each dilution is distinguishable, and
FIG. 4. Relaxation rates and residuals as a function of Mn2þ concentration determined by NMR. (a) 1/T1 at 1.5 T (triangles) and ULF (123–126 mT, circles). Standard deviations of the data are less than 0.1 and 1%, respectively. (b) Residuals from the fits yield a relaxivity uncertainty of 60:06 s1 mM1 at 1.5 T and 60:3 s1 mM1 at ULF. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 5. Representative MRI data for a single pixel of the 0.19 mM [Mn2þ] solution. (a) Inversion recovery at 3 T and (b) prepolarization evolution at 124 mT . Signal is in arbitrary units such that the standard deviation for each point is 61. Solid lines are fits to the data. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
intensity (Fig. 7). The PD vials were restored to their original positions, and the phantom was imaged in the same orientation as in the NiCl2 experiment. The signal magnitude of each PD vial, averaged over circular ROIs, was subsequently corrected using the results of the NiCl2 data. The corrected results showed a linear dependence on H2O concentration, and matched the 1.5T data remarkably well (Fig. 8).
Resolution-Array Results Images from the resolution array obtained with 3T MRI (Fig. 9a) and 128 mT MRI (Fig. 9b) were acquired in 30 s and 36 min, respectively. The 3T images clearly resolved three resolution arrays (1, 2, and 3-mm holes), whereas the 128 mT image clearly resolved the 2 and 3-mm holes. We convolved a perfect digital representation of the
FIG. 6. Grayscale T1 maps at (a) 3 T and (b) 124 mT . Mn2þ concentrations increase from top to bottom and from left to right. White circles show representative ROIs. Measured relaxation times within these ROIs yield relaxivity at (c) 3 T and (d) 124 mT . Solid lines show linear fits, and subplots below show residuals. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 7. Identical vials of NiCl2 imaged with the ULF scanner. (a) Characterization of the spatial dependence of the signal intensity. (b) Variation of intensity as a function of vial position. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
resolution array with a series of Gaussian functions of increasing width, and compared the resulting images with our MRI-acquired image. We then fit the data to obtain the width in y and z simultaneously using a least squares fitting procedure. By fitting the images to this Gaussian degradation, we found 0.8-mm resolution at 3 T, and 1.8 and 2.6-mm resolution at 128 mT in the frequency encoding and phase encoding directions, respectively. DISCUSSION Our ULF NMR measurements demonstrate that the MnCl2 dilutions are appropriate for ULF, span a range of clinically relevant T1 values (50–200 ms) and verify the T1-resolving power of the ULF system. The ULF relaxivity is an order of magnitude greater than that at 1.5 or 3 T. The 1.5T and ULF relaxivity results agree with published
values for the relaxation times of protons in the presence of manganese ions (29). Furthermore, fits to both the 1.5T and ULF NMR data intersect zero concentration at 1=T1 ¼ 0:3660:01 and 0:3660:04 s1, respectively, consistent with the relaxation time of water at both clinical field strengths and ULF. The residuals from the fit showed no systematic error and 61% error bars, readily distinguishing dilutions differing by 0.02 mM [Mn2þ]. The ULF and clinical field relaxation times are also comparable to those of tissues in vivo, making the phantom suite appropriate for T1 standards for both field values. The NMR and MRI measurements of the ULF relaxivity agree to within 2.2%. At 1.5 T, the signal intensity scales with increasing PD. The ULF PD results closely match the results obtained at 1.5 T. However, it was necessary to map out the spatial sensitivity with identical vials of NiCl2 to
FIG. 8. Signal intensity at 107 mT vs. signal intensity at 1.5 T. The blue circles represent the raw, uncorrected data acquired at ULF; red circles are corrected using the results from the NiCl2 phantom. The corrected ULF data are in agreement with the 1.5 T data to within less than 61%. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
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FIG. 9. Grayscale MR images and slice profiles of the resolution inset phantom. (a, c) 3 T and (b, d) 128 mT . Acquisition times were 30 s and 36 min, respectively. Slice profiles through the 3, 2, and 1-mm diameter holes indicated in (a) and (b) by dotted lines are consistent with measured image resolutions of 0.8 mm at 3 T and 1.8 mm at 128 mT . The images have slightly different configurations: the imaging of the resolution array at 3 T was performed simultaneously with the PD measurement, while the 128 mT image was acquired with only a single cylinder from the PD array, acting as an alignment key. [Color figure can be viewed in the online issue, which is available at wileyonlinelibrary.com.]
correct the ULF PD intensity data. This use of the phantom revealed that the signal intensity in the ULF scanner was somewhat spatially dependent. This angular inhomogeneity was subsequently traced to a misalignment of the gradiometer. Although the inhomogeneity does not directly affect T1 or resolution results, it is important to correct for it in deriving PD data from ULF images. The sensitivity map, combined with the corrected PD results, demonstrates that the ULF scanner is capable of measuring PD as accurately as an MRI scanner operating at clinical field strengths. Our measurements of the resolution array at 128 mT agreed with the expected system spatial resolution, set by limitations on gradient strength and SNR to approximately 2 mm. The 3 T scanner was able to resolve the 1-mm features; for the given measurement parameters, the ULF resolution obtained is comparable to those obtained by clinical scanners, albeit with a longer scan time. The feature sizes on the phantom are appropriate for evaluating future improvements of ULF MRI performance.
CONCLUSIONS Traceable phantoms allow us to establish performance metrics for MRI scanners and make meaningful comparisons among them. Phantoms facilitate the improvement of scanner capabilities and the advancement of MRI techniques. Traceable standardized phantoms have been developed for these purposes in clinical MRI; such phantoms have not previously been implemented for ULF MRI. There are many advantages of ULF scanners, such as greater T1 contrast, lack of susceptibility artifacts, low cost, and light weight, which compensate for low SNR. Given the difference in longitudinal relaxation times between cancerous and healthy tissue at ULF, there is a clear potential clinical application for ULF MRI. ULF MRI may provide a simpler, more versatile, and noninvasive method of imaging for diagnosis and treatment monitoring of certain cancers, and potentially other pathologies. It is desirable to characterize ULF system
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performance quantitatively by use of a standardized phantom. The prototype phantoms proved to be well-matched to superconducting quantum interference device-based ULF MRI systems and can already be used to characterize their T1 sensitivity and spatial resolution. Their use facilitates comparison between different scanners at both ULF and clinical field strengths through common metrics such as SNR, imaging time, resolution, and spatial homogeneity of the signal detection sensitivity, and can guide scanner development. The next generation of phantoms will incorporate design changes based on lessons learned from the prototypes, improving stability, facilitating maintenance, providing traceability, and allowing for standardized comparisons of system performance over time and across scanners. ACKNOWLEDGMENTS The authors gratefully acknowledge the assistance of Dr. Hui Dong in identifying the source of the spatially inhomogeneous response of the ULF MRI system. Imaging at 3 T was performed at the University of Colorado Health Sciences Center (Aurora, CO). The 1.5T NMR and MRI data were acquired on a vertical bore system and a horizontal bore scanner, respectively, at the National Institute of Standards and Technology (Boulder, CO). Ultralow field NMR and MRI experiments were conducted at the University of California, Berkeley. Work at UC Berkeley was funded by the Donaldson Trust and the National Institutes of Health. K. B. gratefully acknowledges receipt of a fellowship from the Deutsche Forschungsgemeinschaft. REFERENCES 1. Gunter JL, Bernstein MA, Borowski BJ, Ward CP, Britson PJ, Felmlee JP, Schuff N, Weiner M, Jack CR. Measurement of MRI scanner performance with the ADNI phantom. Med Phys 2009;36:2193–2205. 2. Busch S, Hatridge M, M€ oßle M, Myers W, Wong T, M€ uck M, Chew K, Kuchinsky K, Simko J, Clarke J. Measurements of T1-relaxation in ex vivo prostate tissue at 132 mT. Magn Reson Med 2012;67:1138–1145. 3. Zotev VS, Matlashov AN, Savukov IM, Owens T, Volegov PL, Gomez JJ, Espy MA. SQUID-based microtesla MRI for in vivo relaxometry of the human brain. IEEE Trans Appl Supercond 2009;19:823–826. 4. Russek SE, Boss M, Jackson EF, Jennings DL, Evelhock JL, Gunter JL, Sorensen AG. Characterization of NIST/ISMRM MRI system phantom. In Proceedings of the 20th Annual Meeting of ISMRM, Melbourne, Australia, 2012. Abstract 2456. 5. Koenig SH, Brown RD. Relaxometry of solvent and tissue protons: diamagnetic contributions. Magnetic resonance imaging. Physical principles and instumentation. Vol. 2. Philadelphia: Saunders; 1987. 6. Lee SK, M€ oßle M, Myers W, Kelso N, Trabesinger AH, Pines A, Clarke J. SQUID-detected MRI at 132 mT with T1-weighted contrast established at 10 mT–300 mT. Magn Reson Med 2005;53:9–14. 7. Myers W, Slichter D, Hatridge M, Busch S, M€ oßle M, McDermott R, Trabesinger A, Clarke J. Calculated signal-to-noise ratio of MRI detected with SQUIDs and Faraday detectors in fields from 10 mT to 1.5 T. J Magn Reson 2007;186:182–192. 8. M€ oßle M, Myers WR, Lee SK, Kelso N, Hatridge M, Pines A, Clarke J. SQUID-detected in vivo MRI at microtesla magnetic fields. IEEE Trans Appl Supercond 2005;15:757–760. 9. Packard M, Varian R. Free nuclear induction in the earth’s magnetic field. Phys Rev 1954;93:941. 10. Qiu L, Zhang Y, Krause HJ, Braginski AI, Burghoff M, Trahms L. Nuclear magnetic resonance in the earth’s magnetic field using a nitrogen-cooled superconducting quantum interference device. Appl Phys Lett 2007;91:072505.
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