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J Magn Reson. Author manuscript; available in PMC 2016 November 01. Published in final edited form as: J Magn Reson. 2015 November ; 260: 1–9. doi:10.1016/j.jmr.2015.09.001.

Micron-scale magnetic resonance imaging of both liquids and solids Eric Moore and Robert Tycko* Laboratory of Chemical Physics, National Institute of Diabetes and Digestive and Kidney Diseases, National Institutes of Health, Bethesda, Maryland 20892-0520, U.S.A

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Abstract

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We describe and demonstrate a novel apparatus for magnetic resonance imaging (MRI), suitable for imaging of both liquid and solid samples with micron-scale isotropic resolution. The apparatus includes a solenoidal radio-frequency microcoil with 170 μm inner diameter and a set of planar gradient coils, all wound by hand and supported on a series of stacked sapphire plates. The design ensures efficient heat dissipation during gradient pulses and also facilitates disassembly, sample changes, and reassembly. To demonstrate liquid state 1H MRI, we present an image of polystyrene beads within CuSO4-doped water, contained within a capillary tube with 100 μm inner diameter, with 5.0 μm isotropic resolution. To demonstrate solid state 1H MRI, we present an image of NH4Cl particles within the capillary tube, with 8.0 μm isotropic resolution. High-resolution solid state MRI is enabled by frequency-switched Lee-Goldburg decoupling, with an effective rotating frame field amplitude of 289 kHz. At room temperature, pulsed gradients of 4 T/m (i.e., 170 Hz/μm for 1H MRI) are achievable in all three directions with currents of 10 A or less. The apparatus is contained within a variable-temperature liquid helium cryostat, which will allow future efforts to obtain MRI images at low temperatures with signal enhancement by dynamic nuclear polarization.

Graphical Abstract

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*

corresponding author: Dr. Robert Tycko, National Institutes of Health, Building 5, Room 112, Bethesda, MD 20892-0520, phone 301-402-8272; fax 301-496-0825; [email protected]. Publisher's Disclaimer: This is a PDF file of an unedited manuscript that has been accepted for publication. As a service to our customers we are providing this early version of the manuscript. The manuscript will undergo copyediting, typesetting, and review of the resulting proof before it is published in its final citable form. Please note that during the production process errors may be discovered which could affect the content, and all legal disclaimers that apply to the journal pertain.

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Keywords MRI; solid state NMR; Lee-Goldburg decoupling; constant-time imaging; microcoil

Introduction

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Magnetic resonance imaging (MRI) is an established technique in both the clinic and the research laboratory. Much of the utility of MRI as an imaging technique stems from the wide range of contrast mechanisms available. In principle, any measurement that is possible in conventional nuclear magnetic resonance (NMR) spectroscopy can be used to provide spatial contrast. Although MRI is most widely known for its clinical applications, where the image resolution is typically near 1 mm, MRI images with resolution in the micron regime are also possible [1–7].

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The highest-resolution images reported to date, consisting of 1H NMR images of water in “phantom” samples, have voxel volumes of approximately 30 μm3 [3, 5]. As is well known [8–10], when MRI is based on NMR signals from liquids, the required pulsed field gradient strengths for a given image resolution are dictated by translational diffusion rates. The achievable resolution is then limited by the achievable gradient strengths, and also by the inherently low sensitivity of conventional NMR (since signals from each individual voxel must be detectable above the noise level). MRI images with the highest resolution have been obtained with solenoidal or planar radio-frequency (RF) microcoils, which improve the inherent signal-to-noise ratio (SNR) of the NMR signals [3, 11–15], and are appropriate for samples with dimensions on the order of 1 mm or less. Pulsed field gradients exceeding 5 T/m (i.e., greater than 250 Hz/μm for 1H spins) were used in the highest-resolution experiments, with total measurement times exceeding two days [3, 5]. MRI images of real samples, as opposed to phantoms, have also been obtained with microcoils, with voxel volumes of approximately 250–15000 μm3. Images of unusually

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large single cells have been reported [16–20], as have images of tissue sections [1, 6, 14, 21]. Images with voxel volumes below 1000 μm3 generally require total measurement times on the order of 24 h or more [1, 3, 6, 9, 20]. Based on the existing literature, it appears impractical to obtain MRI images with voxel volumes much less than 30 μm3 with established methods, unless RF microcoils with effective sample volumes much less than 0.001 mm3 are used, along with very large pulsed field gradients and very long measurement times. One alternative is to use magnetic resonance force microscopy [22–28] (MRFM) or NV centers in diamond particles as magnetometers [29–33] to image nuclear spin or electron spin densities. These approaches may be best suited for structural studies of individual objects with sub-micron overall dimensions, rather than for imaging of internal structure within objects with diameters in the 1–100 μm range (such as most biological cells).

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A second alternative [34] is to use RF microcoils, sample temperatures below 100 K, and dynamic nuclear polarization (DNP). In principle, the large enhancements in SNR that can be achieved with low-temperature DNP [35–39] can permit the detection of 1H NMR signals from voxel volumes of 1 μm3 or less [34]. Of course, at low temperatures, the sample to be imaged becomes a solid. 1H signals from water or organic materials are then broadened by strong 1H-1H dipole-dipole couplings, requiring the use of homonuclear decoupling techniques [34, 40–47]. Fortunately, the large RF magnetic fields that can be produced in microcoils facilitate homonuclear decoupling, so that pulsed gradients on the order of 10 T/m are still sufficient for solid state 1H MRI imaging with micron-scale resolution [34].

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In this paper, we describe a novel MRI system based on RF microcoils and planar gradient coils that is capable of both liquid state and solid state imaging. This system has been designed for eventual application in DNP-enhanced, low-temperature MRI. Samples are contained within capillary tubes with 100 μm inner diameters and 170 μm outer diameters, with total image volumes up to 0.003 mm3. We demonstrate the performance of this system at room temperature with three-dimensional (3D) images of two types of phantom samples: (i) polystyrene beads in CuSO4-doped water, for which we report a liquid state 1H image with 6.0 μm isotropic resolution; (ii) NH4Cl particles, for which we report a solid state 1H image with 9.7 μm isotropic resolution. At room temperature, the performance of this system in liquid state imaging is comparable to that of the highest-resolution microcoilbased MRI systems described previously. The performance in solid state imaging is substantially better than any previously described system, in terms of the achievable resolution for small samples.

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Methods MRI probe head Fig. 1a shows a photograph of the MRI probe head, which is mounted on the temperaturecontrolled copper block of a commercial cold finger cryostat (SuperTran-B, Janis Research Co.). The cryostat fits into the 89 mm bore of the superconducting magnet used for imaging experiments at 399.25 MHz (9.38 T). The same cryostat has been used in previous lowtemperature solid state NMR experiments on a variety of chemical and physical systems [48,

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49]. Figs. 1b and 1c show the details of the MRI probe head, which is designed for compatibility with low temperature operation (to be demonstrated in future work) and is therefore constructed from a set of sapphire plates (obtained from O’Keefe Ceramics). Sapphire was chosen as the construction material to provide strong, rigid, thermally conductive, electrically insulating, and cryo-compatible substrates for the magnetic field gradient coils and radio-frequency (RF) sample coil. Even at room temperature, the sapphire plates have the important advantage of conducting heat away from the gradient coils during current pulses, thus preventing the gradient coils from being destroyed during MRI experiments.

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Five square sapphire plates (3.15 cm X 3.15 cm X 0.51 cm) act as a 2.54 cm standoff from the copper block, to minimize eddy currents in the copper that would otherwise interfere with the rapid switching of magnetic field gradients discussed below. The gradient coils are supported above the standoff (as depicted in Fig. 1) by a total of eight plates, four of which are rectangular (6.10 mm X 31.50 mm X 1.52 mm) and four of which are square (6.10 mm X 6.10 mm X 1.52 mm). The lower half of the z gradient (Gz) coil is mounted on the top of rectangular sapphire plate I. Rectangular sapphire plate II is separated from plate I by the two pairs of square sapphire plates. The y gradient (Gy) coil fits into machined grooves in the upper surface of plate II. Rectangular plate III (plastic in experiments described below, but sapphire in future low-temperature experiments) mounts directly on top of plate II. The x gradient (Gx) coil fits into grooves in the lower surface of rectangular sapphire plate IV, which is rotated 90° about z relative to plates I, II, and III. The upper half of the Gz coil is mounted on the top of plate IV. Dimensions of all sapphire plates are shown in Fig. S1. Dimensions of gradient coils are shown in Fig. S2.

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In prior work, Thurber and Tycko estimated the magnetic field gradient necessary to achieve 1 μm resolution in 1H MRI of protonated solids to be approximately 34 T/m, based on the observation of 1H NMR linewidths of 830 Hz under frequency-switched Lee-Goldburg (LG) decoupling [34]. Somewhat smaller linewidths are observed in experiments described below. Using this estimate, the required coil dimensions, turn count, and spacing were determined from Biot-Savart law simulations. Field gradient simulations for idealized representations of the gradient coils are shown in Fig. S2. According to these simulations, pulsed field gradients are predicted to be 0.414 T/m-A, 0.566 T/m-A, and 0.712 T/m-A for Gx, Gy, and Gz coils, respectively. Field gradients are predicted to vary by less than 10% over the sample volumes in our experiments, with only Gx having significant variations.

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Gradient coils are hand-wound from 30 AWG copper magnet wire (0.254 mm diameter) and glued to the sapphire plates with Stycast 2850-FT epoxy. Leads to the gradient coils are 22 AWG wire, tightly twisted and jacketed with heat shrink tubing to prevent their flying apart when current pulses are delivered in the 9.38 T field. In practice, 1 ms pulses of 100 A can be delivered indefinitely with repetition rates of at least 1 Hz without damage to the gradient coils or leads. In experiments described below, peak currents were 20 A, with a 4 Hz repetition rate. Experimental calibrations of the gradient coils, shown in Fig. S3, indicate that the Gx, Gy, and Gz coils actually produce gradients of 0.399 T/m/A, 0.479 T/m/A, and 0.510 T/m/A, respectively. Ideally, the gradient coils produce zero magnetic field in the z direction at the center of the sample, and only the Gx coil produces a transverse field at the J Magn Reson. Author manuscript; available in PMC 2016 November 01.

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center of the sample (which does not affect the NMR measurements). From the onedimensional (1D) images in Fig. S3, we estimate fields in the z direction at the center of the sample to be 2.4 X 10−5 T/A for Gx, 2.6 X 10−5 T/A for Gy, and 2.5 X 10−5 T/A for Gz. RF microcoils are seven-turn solenoids, made from 40 μm diameter bare copper wire (California Fine Wire, Inc.), wound by hand around quartz capillary tubes with 170 μm outer diameter and 100 μm inner diameter (Vitrocom, Inc., part number CV1017Q), and held in place with cyanoacrylate glue. Capillary tubes are cut to lengths of 10–15 mm after coil winding. For liquid state MRI, both ends of the capillary tube are plugged with petroleum jelly after filling with sample. Plugs are not used for solid state MRI. The capillary tube fits into a groove in rectangular plate II, between the grooves for the y gradient coil.

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The microcoil is connected to the rest of the RF circuitry by clamping its two ends against 0.02 mm thick copper leads (yellow pieces in Fig. 1c). Use of this mechanical connection, rather than solder joints, facilitates changing of samples and avoids heating of samples during soldering. A 15 pf “pretuning” capacitor (American Technical Ceramics part number 100B150JMN500X) is also clamped across the copper leads, as close to the microcoil as possible. One lead is soldered to the inner conductor of a semi-rigid coaxial cable, while the other lead is clamped to the outer conductor. The coaxial cable serves as a half-wavelength transmission line to variable tuning and matching capacitors (Polyflon part number RPVC10-6-05), which form the remainder of the parallel-tuned, series-matched RF tank circuit. The variable tuning and matching capacitors are mounted within the cryostat, below the section shown in Fig. 1a. Tuning rods for these capacitors are connected to rotary feedthroughs, so that the RF circuit can be tuned and matched while the cryostat is in the 9.38 T magnet.

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With the design in Fig. 1, the stack of sapphire plates can be disassembled, RF microcoils and samples can be changed, and the stack can be reassembled without soldering or unsoldering individual components. The gradient coils remain properly aligned with one another and with the RF microcoil after disassembly and reassembly. Spectrometer and gradient pulses

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Experiments described below were performed with a Varian InfinityPlus NMR spectrometer console. This console has three RF channels that are controlled by pulse programs, but no separate pulsed field gradient channels. Therefore, RF channel 1 is used to produce RF pulses and detect 1H NMR signals, while RF channels 2 and 3 are used to generate gradient current pulses. Circuitry for converting RF pulses to gradient pulses is depicted in Fig. S4. This circuitry takes advantage of an 80 MHz reference RF signal that is available from the NMR console. The reference signal is split into three lines, with relative phases of 0°, 0°, and 90°. Voltage pulses for Gz are produced by mixing RF pulses from channel 2 (operating at 80 MHz carrier frequency) with the 0° reference signal and sending the product signal through a low-pass filter and amplifier, then to the gradient power amplifier. Voltage pulses for Gx and Gy are produced by sending RF pulses from channel 3 (also operating at 80 MHz carrier frequency) through a 0°/0° power splitter, then mixing the two outputs with 0° and 90° reference signals and sending the product signals through low-pass filters and amplifiers, then to the gradient power amplifiers. By varying the amplitude and phase of the J Magn Reson. Author manuscript; available in PMC 2016 November 01.

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RF pulses from channel 3, any desired combination of x and y gradient pulse amplitudes can be produced. Gradient pulses with amplitudes in the 0–5 V range are used to drive three independent gradient power amplifiers (AE Techron model 7548), operating in controlled voltage mode, which are capable of providing current pulses up to 100 A. In practice, current rise and fall times are approximately 30 μs for the 10–15 A pulses used in experiments described below.

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With this scheme for converting RF pulses to gradient pulses, gradient pulse amplitudes can be sensitive to drift in the phase of the reference RF signal relative to the overall phases of the RF pulse channels. In practice, phase drifts on the order of 10° are observed over 24 h periods. Therefore, a feedback system was devised to adjust the overall RF phases of channels 2 and 3. As depicted in Fig. S4, the feedback system consists of a multifunction data acquisition card (DAQ), which can digitize input signals from the three gradient channels, an independent laptop computer that is connected to the DAQ, and voltagecontrolled phase shifters on spectrometer channels 2 and 3, which are connected to outputs from the DAQ. During image acquisition, after each NMR signal is stored, the imaging pulse program sends a trigger pulse to the DAQ, which then samples 800 points at the three inputs over a 2.4 ms period. During this period, the pulse program also issues 0.5 ms pulses with x and y RF phases on channels 2 and 3. A Python script running on the laptop computer determines the average voltage at each DAQ input during these RF pulses, and calculates the changes at DAQ outputs that are required to adjust RF phases appropriately. Imaging Pulse Sequences

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The pulse sequence used for liquid state imaging is shown in Fig. 2a. To limit the imaged length in the direction parallel to the capillary tube that contains the sample (and thereby reduce the total time required to obtain an image), slice selection is first performed using a truncated sinc-shaped π/2 pulse in the presence of an x gradient, followed by reversal of the gradient to refocus the resulting transverse magnetization along y. (Note that we use x, y, and z to identify both axes of field gradients and axes of the nuclear spin rotating frame, which are not the same. In each case, the meaning should be clear from the context.) Phase encoding gradients with amplitudes Gx, Gy, and Gz are then switched on simultaneously for the first half of the encoding time τph. A hard π pulse is applied at τph/2 to refocus spin precession due to resonance offsets and static field inhomogeneities that do not arise from field gradient pulses. The gradients are then switched on again for the second half of τph with a sign reversal. A hard π/2 pulse then stores either the y or x (i.e., real or imaginary) component of transverse magnetization along the ±z axis and a second hard 90° pulse, applied 3.0 μs later, converts this to ±y magnetization for NMR signal acquisition. Signals are acquired under pulsed spin locking (PSL), consisting of a train of 512 π pulses, separated by 500 μs delays during which NMR signals are digitized [4, 5, 50–53]. The pulse sequence used for solid state imaging is shown in Fig. 2b. Solid state 1H imaging of samples with abundant 1H spins requires a homonuclear decoupling technique [34, 40– 47], i.e., an RF pulse sequence that averages out 1H-1H dipole-dipole couplings, which otherwise broaden 1H NMR lines to 20–50 kHz full-width-at-half-maximum (FWHM). In Fig. 2b, homonuclear decoupling is accomplished with frequency-switched LG irradiation J Magn Reson. Author manuscript; available in PMC 2016 November 01.

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[34, 40, 54] during the phase encoding period. This consists of periods of RF irradiation with the RF carrier frequency shifted by +fLG or −fLG from the on-resonance condition and with RF phases of +y or −y, respectively (denoted by +LG and −LG in Fig. 2b). Due to limitations on the frequency switching times of our spectrometer hardware, carrier frequency offsets are implemented as a windowless train of on-resonance RF pulses with lengths of 150 ns and with RF phase steps of ±φLG between each pulse. The effective frequency offset is then (φLG/360°)×6.6667 MHz, or fLG = 166.67 kHz with φLG = ±9°. Homonuclear decoupling is optimized when the effective field direction in the nuclear spin rotating frame (with z component fLG and y component ν1 = γB1/2π, where γ is the nuclear gyromagnetic ratio and B1 is the rotating RF field amplitude) is tilted from z by the magic angle θm = cos−1(1/√3). If fLG = 166.67, this implies ν1 = 235.70 kHz. The magnitude of the effective field is then 288.68 kHz, which greatly exceeds both the strength of 1H-1H dipoledipole couplings and the magnitudes of NMR frequency offsets induced by field gradient pulses, as required for efficient homonuclear decoupling.

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As shown in Fig. 2b, the pulse sequence begins with a train of nine hard π pulses, separated by dephasing periods τ = 1.0 ms and followed on alternate scans by a narrowband composite π pulse [55] and an additional dephasing period, which serve to suppress 1H NMR signals from material outside the region of homogeneous RF fields in the RF microcoil (e.g., the glue holding the coil to the quartz capillary). A pulse with phase x and flip angle θ = π/2 + θm then rotates the 1H magnetization from ±z to a direction perpendicular to the LG effective field. The ensuing phase encoding period is divided into four periods of length τph/4, with alternation between the +LG and −LG conditions to cancel the net precession around the LG effective field for on-resonance spins. As in Fig. 2a, a hard, on-resonance π pulse in the middle of the phase encoding period refocuses net contributions to spin precession in τph due to resonance offsets and static field inhomogeneities that do not arise from field gradient pulses. Field gradient pulses are applied during both halves of τph, with sign reversal. Real or imaginary components of transverse magnetization (relative to the LG effective field direction) are then stored along ±z by a pulse with phase −x and flip angle θ, or by a pulse with phase −x and flip angle θ followed by a π/2 pulse with phase with phase −y. After a dephasing delay, NMR signals are acquired under PSL, using 4000 pulses with 1.0 μs lengths, 371 kHz RF field strengths, and 10.0 μs separations.

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For 3D images, each gradient amplitude Gq is varied between −nqΓq and +(nq−1)Γq, where Γq is the gradient increment along the q axis. The total number of time-domain PSL signals S(kx,ky,kz,t), with kq ≡ nqγΓqτph, is then 16nxnynz. In Fig. 2a, an eight-step phase cycle is applied to the hard 180° and 90° storage pulses to eliminate artifacts, requiring a minimum of 8 scans for each time-domain signal. In Fig. 2b, only the final 90° is phase-cycled, so that the minimum number of scans is two. PSL extends the NMR signals in time, increasing the total signal acquired in each scan and hence the image signal-to-noise ratio by a factor of approximately (T2/T2*)1/2, where T2 is the intrinsic transverse spin relaxation time and T2* is the apparent transverse relaxation time (including signal decay due to static field inhomogeneities and other sources that can be eliminated by appropriate RF pulse sequences). In practice, the signal-to-noise enhancement factor in our liquid state imaging experiments is approximately 2.5.

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In principle, a 3D image could be acquired with fewer scans (by a factor of 2nq) if gradient Gq was applied with a constant amplitude during NMR signal acquisition, PSL was omitted, and only the two remaining gradients were applied during τph. However, the signal-to-noise ratio of the resulting image would be lower by a factor of approximately [2nqT2/(αT2*)]1/2. This factor assumes ideal PSL, meaning that NMR signals are digitized at a constant rate without interference from the RF pulse train, that the audio filter bandwidth of the NMR receiver electronics is optimally matched to the digitization rate, and that signals are digitized for a total period of approximately T2. This factor also assumes that Gq is chosen so that the desired image resolution in the q direction is obtained by digitizing NMR signals for total period αT2*, with α ≤ 1. If 2nq images without PSL were added together, so that the total measurement times were the same with and without PSL, then the signal-to-noise ratio without PSL would still be lower by a factor of approximately [T2/(αT2*)]1/2. For liquid state imaging, this factor can be small if static field inhomogeneities are small. In our liquid state imaging experiments, this factor is approximately 2.5. For solid state imaging, this factor is necessarily large, because T2* is determined by NMR signal decay due to 1H-1H dipole-dipole couplings. PSL effectively attenuates these couplings, making T2 ≫ T2*. In our solid state imaging experiments, (T2/T2*)1/2 ≈ 7. Solid state imaging without PSL and with Gq applied during NMR signal acquisition would also require Gq to be approximately 50 times larger than in the experiments described below, which would be beyond the capabilities of our hardware.

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Data processing

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Data sets were processed off-line using custom Python scripts. The first two data points in each time-domain signal S(kx,ky,kz,t)R and S(kx,ky,kz,t)I were discarded (where subscripts R and I refer to the two magnetization components stored after the phase encoding period), then an exponential window (9 Hz for liquids, 35 Hz for solids), Fourier transformation with respect to t, and phase correction were applied to produce spectra S′(kx,ky,kz,ω)R and S′ (kx,ky,kz,ω)I. The center point in each PSL spectrum (proportional to the area of the timedomain PSL signal) was extracted and the complex k-space signal was constructed according to S″ (kx,ky,kz) = Re{S′ (kx,ky,kz,0)R} + iRe{S′ (kx,ky,kz,0)I}. Fourier transformation with respect to kx, ky, and kz to produce the final image S‴ (x,y,z) was then performed after baseline correction and zero-filling in each dimension, with no additional apodization. Images were circularly permuted in x, y, and z as needed to center the imaged objects.

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Liquid state imaging As an initial test of the performance of our MRI system, images were obtained of a capillary tube filled with water and 20.7 ± 0.3 μm diameter polystyrene beads (Polysciences, Inc.). The water was paramagnetically doped with 28 mM CuSO4 to reduce the 1H spin-lattice relaxation time (T1) to 0.2 s, permitting the recycle delay between NMR signal acquisitions to be 0.25 s. To load beads into the capillary tube, the capillary was inserted into a centrifuged pellet of beads in a 0.15 ml plastic tube. The imaged length along the x direction (i.e., the long axis of the capillary) was limited to approximately 100 μm by slice selection,

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as described above. The effectiveness of slice selection is illustrated by the 1D images in Fig. S5.

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Fig. 3 shows selected two-dimensional (2D) planes and 1D slices from a 3D image that was obtained under conditions that produce nominal isotropic resolution Δ = 5.0 μm, as calculated from the expression Δ = (2γnqΓqτph)−1. If an additional factor of 1.21 is included, arising from the effect of k-space truncation on the MRI point spread function as discussed by Webb [10], the resolution is Δ′ = 6.1 μm. Additional 2D planes are shown in Fig. S6. Specifically, the 3D image was acquired with τph = 900 μs, Γx = Γy = Γz = 0.163 T/m, and nx = ny = nz = 16. As described above, RF phase cycling required 8 scans for each timedomain signal. Four images were added together to produce the final image in Fig. 3, which is therefore the result of 2097152 total scans, acquired in a total of 150 h. The k-space data were processed with zero-filling by factors of two in each dimension, producing a final image with 64 X 64 X 64 voxels, spanning a 160 μm X 160 μm X 160 μm volume. 1D slices in Fig. 3 show transitions from maximum image intensity to minimum image intensity within two points, verifying the expected 5.0 μm resolution. Further quantification of the image resolution, by comparison of experimental and simulated image slices, is given in Fig. S7. 2D planes show that the beads tend to associate with the inner walls of the capillary tube. The absence of significant image distortions or artifacts in Figs. 3 and S6 indicates that field gradients in our MRI system are nearly constant over the imaged volume and the MRI system is stable over long experiments. With a translational self-diffusion constant of 2.2 × 10−5 cm2/s, the root-mean-squared displacement of water molecules during τph is 1.4 μm. We therefore expect self-diffusion to have a negligible effect on the image resolution in Fig. 3 [8–10, 56].

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Solid state imaging

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Data in Fig. 4 demonstrate the performance of the frequency-switched LG technique for removing 1H-1H dipole-dipole couplings. These data were obtained with the pulse sequence in Fig. 2b, but without gradient pulses. Fig. 4a shows the dependence of the 1H NMR signal amplitude from NH4Cl particles on ν1 and τph. For these experiments, NH4Cl particles with appropriate diameters were selected from commercial NH4Cl powder under an optical microscope, placed into the capillary tube by hand, and positioned within the RF microcoil by pushing them with fine wire. Values of ν1 were determined from a 1D “nutation image” of the same sample (see Fig. S8), which shows that RF fields are homogeneous to within ±3% over a 250 μm long region within the capillary tube. The signal decay time during the LG period has a sharply defined maximum at ν1 = 235 ± 3 kHz, in good agreement with the expected value of 235.7 kHz discussed above. As shown in Fig. 4b, effective T2 values in the LG period were approximately 1.0 ms at the optimal value of ν1, for both NH4Cl particles and a powder of the tripeptide AlaGlyGly. The sharper dependence on ν1 for AlaGlyGly reflects stronger 1H-1H couplings. 1D 1H NMR spectra of solid NH4Cl and AlaGlyGly (Fig. S9) show FWHM linewidths of 25 kHz and 50 kHz, respectively, implying signal decay times of 10–20 μs in the absence of homonuclear decoupling. Based on the results in Fig. 4, a value of τph = 600 μs was chosen for 3D solid state imaging. We chose to

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use NH4Cl for 3D imaging, rather than AlaGlyGly, because of its shorter T1 value (0.5 s vs. 1.5 s). Fig. 5 shows selected 2D planes and 1D slices from a 3D image of NH4Cl particles. Additional 2D planes are shown in Fig. S10. The 3D image was acquired with τph = 600 μs, Γx = 0.132 T/m, Γy = Γz = 0.283 T/m, nx = 32, ny =15, and nz = 15. The effective gradient amplitudes are scaled down by the LG irradiation, ideally by a factor of cosθm = 0.5774, so that these gradients lead to an expected resolution of Δ = 8.02 μm (Δ′ = 9.71 μm) in the x direction and Δ = 7.99 μm (Δ′ = 9.67 μm) in the y and z directions. The 3D image was acquired with 4 scans per time-domain signal and a 0.5 s recycle delay, requiring a total measurement time of 64 h. The k-space data were processed with zero-filling by factors of two in each dimension, producing a final image with 128 X 60 X 60 voxels, spanning a 514 μm X 240 μm X 240 μm volume.

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1D slices in Fig. 5 show transitions from maximum image intensity to minimum image intensity within two pixels, verifying the expected 8.0 μm resolution. The 2D planes show that the particles have irregular shapes, with indentations on the 10 μm length scale. Gaps between the particles are clearly resolved, with widths of 8–12 μm. Although the RF microcoil is 450 μm long, image intensities fade beyond a length of 300 μm in the x direction. This observation indicates that RF fields within the microcoil are sufficiently homogeneous to produce effective homonuclear decoupling in the central 2/3 of the microcoil length, but that NMR signals from material near the edges of the coil or outside the coil are suppressed by the background suppression pulses shown in Fig. 2b and by ineffective homonuclear decoupling.

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Discussion The MRI apparatus described above has several advantageous features. The gradient design is simple, compact, effective, and relatively immune to damage from high current pulses. The fact that gradient coils and the RF microcoil are mounted on a series of stacked plates makes disassembly, sample changes, and reassembly relatively easy. If a gradient coil were to fail, which in fact did not occur in the course of the experiments described above, it would be a simple matter to replace the affected plate.

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The image resolution in Fig. 3 is comparable to the highest resolutions reported to date for liquid state MRI images obtained with conventional NMR signal detection. In 2002, Pennington and coworkers reported a 3D image of a similar phantom sample with 3.7 μm X 3.3 μm X 3.3 μm resolution [5], acquired in a 9.0 T field in approximately 30 h, using a solenoidal RF microcoil with 73 μm inner diameter and maximum pulsed field gradients of 4.6–5.8 T/m. The higher resolution in the experiments of Pennington and coworkers is largely attributable to the smaller microcoil, which leads to a higher SNR per sample volume but a smaller total volume (approximately 7 X 10−4 mm3, compared with the 1 X 10−2 mm3 microcoil volume in our experiments). In 2008, Weiger et al. reported a 3D image of glass fibers in water with 3.0 μm isotropic resolution [3], acquired in a 18.8 T field in 58 h, using a planar RF microcoil and maximum pulsed field gradients of 24.4 T/m. Based on the

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published 2D slices from this image, it appears that the effective microcoil volume in the experiments of Weiger et al. was approximately 6 X 10−4 mm3.

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To our knowledge, the image resolution in Fig. 5 is substantially higher than that of any previous 3D MRI image of abundant spins in a rigid solid [42–47], aside from images of much smaller sample volumes obtained with MRFM [27]. The quality of the image in Fig. 5 is a consequence of the efficient suppression of 1H-1H dipole-dipole couplings that we achieve with LG irradiation, which is aided by the large and homogeneous RF fields available within the microcoil. Although LG irradiation averages 1H-1H couplings to zero in lowest order, higher-order effects involving cross terms between different pairwise couplings can contribute significantly to signal loss during τph unless large RF fields are used. In addition, in our solid state MRI experiments, the pulsed field gradients produce 1H NMR frequency offsets up to approximately 40 kHz near the edges of the imaged volume. Large RF fields are therefore required to minimize cross terms between dipole-dipole couplings and frequency offsets. The overall stability of the apparatus over long experiments also contributes to the image quality in our experiments. As described above, our MRI probe head is contained in a cryostat that permits operation at temperatures in the 5–300 K range. For solid state MRI, low temperature operation is expected to enhance SNR and hence spatial resolution. Operation at low temperatures is not expected to impair the performance of the gradients or RF circuitry in our apparatus. In fact, higher gradient currents should be possible at lower temperatures, due to the reduced electrical resistance of the gradient wires and the more efficient heat dissipation of the sapphire pieces. The efficiency of the RF circuitry may also improve.

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With low-temperature DNP, isotropic resolution of 1 μm or better should be feasible [34], potentially enabling imaging of structures within typical eukaryotic cells and cell clusters. As depicted in Fig. 1c, a minor reconfiguration of the sapphire plates creates the necessary path for microwaves to the sample in future DNP-enhanced MRI experiments. Specifically, if plates I and IV are swapped, microwaves can travel through 2.90 mm diameter holes in plates I and III to irradiate the sample within the RF microcoil. Challenges that will be addressed as this work progresses include maximization of microwave field strengths within the sample, calibration of RF fields at low temperatures, and optimization of sample preparation and paramagnetic doping to maximize DNP-enhanced NMR signals and image contrast.

Supplementary Material Author Manuscript

Refer to Web version on PubMed Central for supplementary material.

Acknowledgments This work was supported by the Intramural Research Program of the National Institute of Diabetes and Digestive and Kidney Diseases, a component of the National Institutes of Health.

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Figure 1.

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MRI probe head. (a) Photograph of the assembled probe head, which is mounted on the temperature-controlled copper block of a Janis Supertran-B cryostat through a set of five sapphire plates with dimensions 3.15 cm X 3.15 cm X 0.51 cm. (b) Drawing of the assembled probe head. (c) Exploded drawing. Magnetic field gradient pulses for imaging are produced by x, y, and z gradient coils, which are supported by 1.52 mm thick sapphire plates and potted in black Stycast epoxy. The RF microcoil for excitation and detection of 1H NMR signals is wound around a quartz capillary tube with 170 μm outer diameter, which contains the sample to be imaged. The microcoil is tuned close to the 1H NMR frequency of 399.2 MHz by a 15 pf capacitor and connected to the remaining RF circuitry by a halfwavelength coaxial cable. Orange and pink pieces act as clamps to hold the capillary and RF components in place. Asterisks indicate pieces that are fabricated from Delrin, rather than sapphire.

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Figure 2.

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MRI pulse sequences. (a) Sequence for liquid state imaging. Slice selection is performed using a weak, sinc-modulated π/2 pulse in the presence of an x gradient, followed by reversal of the gradient to refocus nuclear spin magnetization. All three gradients are applied during the constant-time phase encoding period τph, with opposite signs before and after a π pulse that refocuses gradient-independent NMR frequency offsets and static field inhomogeneities. Gradient amplitudes in the three directions are varied independently as described in the text. A π/2 pulse with phase φ2 selects either the x or y component of nuclear spin magnetization for storage along z. Another π/2 pulse converts the selected component to transverse magnetization, and NMR signals are digitized in periods between PSL pulses. RF pulse phases are cycled according to φ1 = x, x, −x, −x, y, y, −y, −y; φ2 = x, −x (real) or y, −y (imaginary). Signals in successive scans are coadded with the pattern +, −, +, −, −, +, −, +. (b) Sequence for solid state imaging. Background signal suppression is accomplished with a train of nine π pulses, separated by delays with τ = 1 ms, followed by a composite narrowband π pulse on alternate scans (dashed lines). After preparation of transverse magnetization by a pulse with flip angle θ, 1H-1H dipole-dipole couplings and gradient-independent NMR frequency offsets are removed during τph by a frequencyswitched Lee-Goldburg sequence and a hard π pulse with phase x, as described in the text. The x component of transverse magnetization (relative to the Lee-Goldburg effective field direction) is stored along z by a second θ pulse with phase −x, or the y component is stored by a θ pulse with phase −x followed by a π/2 pulse with phase −y. A final π/2 pulse with phase φ3 = x, −x converts the selected component to transverse magnetization, and NMR signals are digitized in periods between PSL pulses and coadded with the pattern +, −.

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Figure 3.

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Liquid state image of a 100 μm inner diameter capillary tube containing CuSO4-doped water and 20 μm diameter polystyrene beads. The image is acquired with the pulse sequence in Fig. 2a under conditions that produce 5.0 μm isotropic resolution (see text). The imaged volume is limited to approximately 75 μm in the x direction by slice selection as in Fig. S5. (a–c) 2D planes at y = 27.5 μm, at z = −15.0 μm, and at x = −22.5 μm, respectively. (d–f) 1D slices along x at y = 27.5 μm, z = −15.0 μm, along y at x = −22.5 μm, z = −15.0 μm, and along z at x = −22.5 μm, y = 27.5 μm. Color-coded lines indicate the positions of these 1D slices in the 2D planes. Black circles indicate their point of intersection.

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Figure 4.

Performance of frequency-switched Lee-Goldburg decoupling. (a) 1H NMR signals from NH4Cl under the pulse sequence in Fig. 2b without field gradient pulses, as a function of τph (0.004 ms to 2.004 ms in 0.04 ms increments) and the RF field amplitude during ±LG blocks. (b) Signal decay time T2 during τph as a function of the RF field amplitude during ±LG blocks, for NH4Cl (filled symbols, red line) and AlaGlyGly powder (open symbols, blue line). Top axis indicates the angle between the Lee-Goldburg effective field and the z axis. T2 values were obtained by fitting experimental data, as in panel a, with exponential decay functions. (c) Signal amplitude as a function of the RF amplitude during ±LG blocks for NH4Cl at τph = 0.244 ms (blue), 0.524 ms (red), and 1.004 ms (green). (d) Signal amplitude as a function of the RF amplitude during ±LG blocks for AlaGlyGly at τph = 0.252 ms (blue), 0.500 ms (red), and 1.000 ms (green).

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Author Manuscript Author Manuscript Figure 5.

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Solid state image of NH4Cl particles in a 100 μm inner diameter capillary tube. The image is acquired with the pulse sequence in Fig. 2b under conditions that produce 8 μm isotropic resolution. The imaged volume is limited to approximately 300 μm in the x direction by the background suppression pulses in Fig. 2b and the sensitivity of Lee-Goldburg decoupling to RF field homogeneity. (a–c) 2D planes at z = 0.0 μm, at x = 44.0 μm, and at y =−4.0 μm, respectively. (d–f) 1D slices along x at y = −4.0 μm, z = 0.0 μm, along y at x = 44.0 μm, z = 0.0 μm, and along z at x = 44.0 μm, y = −4.0 μm. Color-coded lines indicate the positions of these 1D slices in the 2D planes. Black circles indicate their point of intersection. (g) Optical microscope image of the seven-turn microcoil wound around the capillary. The capillary has an outer diameter of 150 μm.

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