Hyperpolarized 129Xe MRI of the human lung

Abstract

By permitting direct visualization of the airspaces of the lung, magnetic resonance imaging (MRI) using hyperpolarized gases provides unique strategies for evaluating pulmonary structure and function. Although the vast majority of research in humans has been performed using hyperpolarized (3)He, recent contraction in the supply of (3)He and consequent increases in price have turned attention to the alternative agent, hyperpolarized (129) Xe. Compared to (3)He, (129)Xe yields reduced signal due to its smaller magnetic moment. Nonetheless, taking advantage of advances in gas-polarization technology, recent studies in humans using techniques for measuring ventilation, diffusion, and partial pressure of oxygen have demonstrated results for hyperpolarized (129)Xe comparable to those previously demonstrated using hyperpolarized (3)He. In addition, xenon has the advantage of readily dissolving in lung tissue and blood following inhalation, which makes hyperpolarized (129)Xe particularly attractive for exploring certain characteristics of lung function, such as gas exchange and uptake, which cannot be accessed using (3)He. Preliminary results from methods for imaging (129) Xe dissolved in the human lung suggest that these approaches will provide new opportunities for quantifying relationships among gas delivery, exchange, and transport, and thus show substantial potential to broaden our understanding of lung disease. Finally, recent changes in the commercial landscape of the hyperpolarized-gas field now make it possible for this innovative technology to move beyond the research laboratory.

Figure 1

Optical pumping and spin exchange. (a) Simplified conceptual diagram of the optical-pumping and spin-exchange process. In optical pumping (left), circularly-polarized laser light at the appropriate wavelength polarizes the spins of the valence electrons (shown in green) of rubidium atoms by preferentially populating one of the two spin states for the valence electron. In this diagram, electron-spin polarization is illustrated by a change in the electron-spin orientation (symbolized by the small black arrow through the electron) from down to up. Spin-exchange collisions (right) transfer this polarization from rubidium electrons to noble-gas nuclei. After spin exchange, the spins of the valence electrons of rubidium atoms are polarized again and participate in subsequent spin-exchange collisions, further increasing the noble-gas polarization. (Diagram adapted from Fig. 1 of ref. (99).) (b) Photograph of an early prototype commercial gas-polarization system (model IGI 9600, Magnetic Imaging Technologies, Inc. [MITI], Durham, NC), showing some of the system’s major functional aspects. This system, the first manufactured by MITI, has an exchangeable central “cartridge” to permit operation with either ¹²⁹Xe or ³He.

Figure 2

State-of-the-art gas-polarization system for ¹²⁹Xe. These photographs show a prototype commercial fourth-generation gas-polarization system optimized for ¹²⁹Xe (model XeBox-E10, Xemed, LLC, Durham, NH). This polarizer is based on the flowing approach for hyperpolarized-gas production as discussed in the text. The polarizing column, which is the functional analog of the optical cell shown in Fig. 1, is indicated on the right. Additional technical details can be found in ref. (28).

Figure 3

Effect of T1 relaxation on thermally-polarized versus hyperpolarized magnetization. These plots compare the evolution of the longitudinal component of thermally-polarized magnetization (left column) to that of hyperpolarized magnetization (right column) for the first several repetitions of either a single spin-echo (upper row) or low-flip-angle gradient-echo (lower row) pulse sequence. Due to the non-equilibrium nature of the hyperpolarized magnetization, pulse sequences using repeated, high-flip-angle RF pulses, such as a spin-echo method as illustrated in the upper row, are not useful for hyperpolarized-gas imaging of the lung. In contrast, for appropriate pulse-sequence parameter values, low-flip-angle gradient-echo imaging yields similar relative magnetization evolutions for both thermally-polarized and hyperpolarized magnetization as illustrated in the lower row. For the single spin-echo pulse sequence a 90° excitation RF pulse is applied every 0.5 s (i.e., TR = 500 ms) starting at time zero, and for the low-flip-angle gradient-echo pulse sequence a 10° excitation RF pulse is applied every 20 ms starting at time zero. For thermally-polarized and hyperpolarized magnetization, the T1 values are representative of biological tissue (1 s) and ¹²⁹Xe in the lung (20 s), respectively, and the initial values of the magnetization are denoted M_o and M_h, respectively. Ideal spoiling of transverse magnetization between pulse-sequence repetitions is assumed.

Figure 4

Ventilation imaging of the healthy human lung. Coronal ventilation images were acquired using hyperpolarized ¹²⁹Xe (upper and lower rows) or ³He (middle row). In both the ³He images shown in the middle row and the ¹²⁹Xe images shown in the lower row, hyperpolarized gas distributed generally uniformly throughout the ventilated airspaces of the lung. The tremendous improvement in the quality of ¹²⁹Xe images between 1996 and 2009 is attributable primarily to an increase in gas polarization; gas polarizations were approximately 1% for the images in the upper row and 35% for those in the middle and lower rows. The images in the upper and middle rows were acquired at 1.5T using a single-channel chest RF coil, while those in the lower row were acquired at 3T using a 32-channel array chest RF coil (33). All images were obtained using a low-flip-angle gradient-echo (FLASH) pulse sequence. (Images in the upper row are reproduced with permission from Fig. 1 of ref. (12), and those in the lower row are from the healthy subject shown in Fig. 1 of ref. (48).)

Figure 5

Ventilation imaging of the diseased human lung using ¹²⁹Xe. Coronal ventilation images were acquired in subjects with asthma (upper row), COPD (middle row) or cystic fibrosis (lower row). Numerous ventilation defects are seen in each of the images secondary to airflow obstruction caused by the underlying diseases. Gas polarizations were between approximately 25% and 35%. The images in the upper and middle rows were acquired at 3T using a 32-channel array chest RF coil, while those in the lower row were acquired at 1.5T using a single-channel chest RF coil. All images were obtained using a low-flip-angle gradient-echo-based (FLASH or TrueFISP) pulse sequence. (Images in the middle row are from the corresponding subject shown in Fig. 1 of ref. (48). Images in the bottom row are from the same subject shown in the first row of Fig. 1 in ref. (100).)

Figure 6

Comparison of ventilation imaging using ¹²⁹Xe and ³He in cystic fibrosis. Coronal ventilation images were acquired in two subjects with cystic fibrosis; imaging for each subject was completed on the same day. While, in subject A, the ventilation defects in ¹²⁹Xe images appear very similar to those in ³He images, in subject B the ventilation defects in ¹²⁹Xe images appear more severe than those in ³He images. This difference observed in subject B is possibly attributable to the different physical properties of ¹²⁹Xe and ³He gases. All images were acquired at 1.5T using a single-channel chest RF coil and a low-flip-angle gradient-echo-based (FLASH or TrueFISP) pulse sequence. (Images in the top and bottom rows are from the same subjects shown in the second and first rows, respectively, of Fig. 1 in ref. (100).)

Figure 7

Comparison of diffusion imaging using ¹²⁹Xe and ³He. Coronal ADC maps were obtained in two healthy subjects and in two subjects with COPD using ¹²⁹Xe (left two images) or ³He (right two images). For both gases, the ADC values are relatively low and spatially uniform in the healthy subjects. In contrast, the ADC values for the COPD subjects are markedly inhomogeneous and generally elevated compared to those for the healthy subjects, particularly in the apices. All images were obtained using a low-flip-angle gradient-echo (FLASH) pulse sequence that included a bipolar gradient waveform, applied between the excitation RF pulse and the spatial encoding gradients, to impart diffusion sensitization (b values: 0 and 10 s/cm² [¹²⁹Xe] or 0 and 1.6 s/cm² [³He]). (¹²⁹Xe ADC map for COPD is from the same subject shown in Fig. 4 of ref. (91).)

Figure 8

Directional gradients in ¹²⁹Xe ADC values. Coronal ADC maps obtained in a healthy subject demonstrate gradients of the ADC values in both the anterior-posterior and superior-inferior directions. The anterior-posterior gradient is attributed to the well-known gravity-dependent gradient in lung tissue density from anterior to posterior in the supine position. Kaushik et al suggest that the superior-inferior gradient may be associated with regional variation in concentration of the inhaled ¹²⁹Xe (36), although similar variations in ADC values have also been observed using ³He (71). (Fig. 5 from ref. (36), reproduced with permission.)

Figure 9

Comparison of oxygen-concentration imaging using ¹²⁹Xe and ³He. Coronal maps of the partial pressure of oxygen (pO₂) were obtained in healthy subjects using either ¹²⁹Xe (left) or ³He (right). For both gases, the pO₂ values are generally uniform and in a range consistent with normal lung physiology, except for artifactually high values at the base of the lung caused by slight motion of the diaphragm during the acquisitions. Both images were obtained using a gradient-echo-based, cardiac-triggered, short-breath-hold implementation of pO₂ mapping (101). (¹²⁹Xe pO₂ map is from healthy subject #1 in ref. (101).)

Figure 10

¹²⁹Xe in the lung. Inhaled xenon gas distributes among the airspaces of the lung and also dissolves in the lung parenchyma and blood. Because of the large chemical-shift, and hence frequency, difference between ¹²⁹Xe in the airspaces (alveoli) and that in the parenchyma/blood, an MR spectrum of ¹²⁹Xe in the lung reveals distinct spectral peaks corresponding to the gas and dissolved components. For clarity, the dissolved-phase peaks, associated with roughly 2% of ¹²⁹Xe in the lung, are shown several times larger than they would appear if the same flip angle was applied to both the gas-phase and dissolved-phase compartments. (Lung rendering adapted from: dir.niehs.nih.gov/dirlrb/mcb/proj-cox.htm. Micrographs reproduced with permission from Albertine KH. Structural organization and quantitative morphology of the lung. In: Cutillo AG, ed. Application of Magnetic Resonance to the Study of Lung. Hoboken, NJ: Wiley-Blackwell, 1996; 73–114.)

Figure 11

¹²⁹Xe exchange between lung airspaces and tissue. Following inhalation, a dynamic equilibrium is quickly established between xenon in the airspaces and xenon dissolved in the parenchyma and blood, resulting in diffusion-driven exchange of xenon between the airspaces (blue Xe atoms) and dissolved-phase compartments (red and gold Xe atoms). A fraction of dissolved xenon is transported to other organs by the bloodstream. (Diagram adapted from Fig. 2 of ref. (20). Micrograph reproduced with permission from Albertine KH. Structural organization and quantitative morphology of the lung. In: Cutillo AG, ed. Application of Magnetic Resonance to the Study of Lung. Hoboken, NJ: Wiley-Blackwell, 1996; 73–114.)

Figure 12

XTC: Xenon polarization transfer contrast. Following inhalation of hyperpolarized gas, a low-flip-angle, gradient-echo image of gas-phase ¹²⁹Xe is acquired (left). Next, image contrast reflecting the degree of gas exchange between the airspaces and tissue/blood is generated by applying a series of high-flip-angle RF pulses (green arrows in plots) at the frequency corresponding to the dissolved-phase compartments. During the delay between applications of these contrast-generating high-flip-angle RF pulses, exchange of xenon atoms between the airspaces and tissue/blood causes the gas-phase longitudinal magnetization to decrease (plot on right, blue curve) by an amount proportional to the degree of exchange. Finally, a second low-flip-angle, gradient-echo image of gas-phase ¹²⁹Xe is acquired (right). Analysis of the changes in signal intensity between the first and second images provides a regional map (top) reflecting the degree of gas exchange that occurred during the delay between contrast-generating RF pulses. For clarity, the dissolved-phase peaks (plot on left) and associated plots of longitudinal magnetization (plot on right) are shown several times larger than actual values.

Figure 13

XTC in a healthy human subject. Maps of fractional gas exchange were calculated from single breath-hold XTC acquisitions at 0.2T (25) performed at 47% (left) and 63% (right) of total lung capacity (TLC) in a healthy volunteer. The mean value of the fractional gas exchange decreased from 1.8% (47% of TLC) to 1.4% (63% of TLC) in response to the increase in lung volume. Although the fractional gas exchange appeared fairly uniform across the lung in both cases, summation of the values along the left-right direction revealed an increase in fractional gas exchange of roughly 0.5% from apex to base for 47% of TLC, whereas the corresponding values for 63% of TLC were uniform from apex to base (25). The time delay between applications of high-flip-angle RF pulses to the dissolved-phase compartments was 62 ms; other acquisition parameters are described in ref. (25). Note that the color scale for the map on the left has a maximum of 6%, while that for the map on the right has a maximum of 3%. (Maps of fractional gas exchange reprinted from European Journal of Radiology, 64(3), Patz S, Hersman FW, Muradian I, et al., Hyperpolarized ¹²⁹Xe MRI: a viable functional lung imaging modality?, 335–344, Copyright 2007, with permission from Elsevier.)

Figure 14

Lung microstructural parameters from XTC. Maps of the parameters MXTC-F (left half of figure), proportional to the tissue to alveolar-volume ratio, and MXTC-S (right half of figure), proportional to the mean septal wall thickness, were derived from multiple-exchange-time XTC acquisitions at 3T (91). In a healthy volunteer (upper row), both MXTC-F and MXTC-S clearly increased along the anterior-posterior direction, and were either generally uniform or showed relatively small variations within coronal sections. In contrast, for a subject with COPD (lower row), MXTC-F showed marked regional variations within sections with lower values in the apices. Further, MXTC-S did not show substantial variation along the anterior-posterior direction, but was generally higher than for the healthy subject. Acquisition parameters are described in ref. (91). (Maps reproduced with permission from Fig. 2 in ref. (91).)

Figure 15

Direct imaging of dissolved ¹²⁹Xe in the human lung. (a) Three-dimensional acquisitions depicting gas-phase ¹²⁹Xe in the lung airspaces (left side of each section) and dissolved-phase ¹²⁹Xe in the tissue and blood (right side of each section) from a healthy subject (top) and a subject with mild (GOLD stage 1) COPD (bottom) (95). While the signal-intensity distributions for the gas-phase and dissolved-phase components were generally uniform for the healthy subject, except for noticeably higher dissolved-phase signal intensity in the posterior portion of the lung, both gas-phase and dissolved-phase components showed non-uniform signal distributions in the subject with COPD, with some regions of mismatch between corresponding gas-phase and dissolve-phase signal variations. Acquisition parameters are described in ref. (95). (Adapted from Fig. 2 in ref. (95).) (b) Maps showing the ratio of dissolved-phase to gas-phase ¹²⁹Xe in the healthy (top row) and COPD (middle row) subjects, and showing the ADC for ¹²⁹Xe (b values of 0 and 10 s/cm²) for the COPD subject (lower row). In the anterior portion of the lung, the ratio of dissolved-phase to gas-phase ¹²⁹Xe for the COPD subject was substantially higher than that for the healthy subject. Regions of relatively low ratio values for the COPD subject were generally associated with elevated ADC values (e.g., arrowheads).