In pathologies in which slow or collateral flow conditions may exist, conventional arterial spin labeling (ASL) methods that apply magnetic tags based on the location of arterial spins may not provide robust measures of cerebral blood flow (CBF), as the transit delay for the delivery of blood to target tissues may far exceed the relaxation time of the tag. Here we describe current methods for ASL with velocity-selective (VS) tags (termed VSASL) that do not require spatial selectivity and can thus provide quantitative measures of CBF under slow and collateral flow conditions. The implementation of a robust multislice VSASL technique is described in detail, and data obtained with this technique are compared with those obtained with conventional pulsed ASL (PASL). The technical considerations described here include the design of VS pulses, background suppression, anisotropy with respect to velocity-encoding directions, and CBF quantitation issues. In conventional arterial spin labeling (ASL) techniques, including both pulsed ASL (PASL) (1-5) and continuous ASL (CASL) (6 -8), arterial blood is tagged by magnetic inversion or saturation proximal to the region of interest (ROI). Tagged blood then flows into the ROI, and the inflow is detected as a modulation of the longitudinal magnetization. In these techniques there is necessarily a spatial gap between the tagging location and the ROI. This gap results in a transit delay (␦t) for the delivery of tagged blood to the ROI. The gap (and hence the delay) can be small for single-slice imaging, but is larger for multislice or volume acquisitions. The magnitude and variability of the transit delay in relation to the T 1 decay of the tag is one of the largest potential sources of errors in the quantitation of perfusion using ASL in the normal human brain (5,6,9). In stroke and other pathologies in which flow may be slow or may follow circuitous collateral routes of delivery, ␦t can be much larger than T 1 (10), which makes conventional ASL an impractical method for obtaining accurate measures of CBF.We recently introduced a new ASL method in which the tag pulse is purely velocity-selective (VS) and not spatially-selective. This allows for the tagging of all flowing spins within a specified velocity range, regardless of location, and can in principle eliminate the problem of transit delays. We refer to this technique as VSASL, and in this work we describe the implementation of VSASL, as well as some of the considerations involved in the design of VSASL pulse sequences and the quantitation of perfusion using this technique. VSASL was introduced in abstract form in Ref. 11, and some of the issues addressed herein were described in suggested the use of VS pulses in ASL, but did not present an implementation of VSASL. THEORYIn principle, the elements of a VSASL pulse sequence are similar to those of a conventional ASL experiment and include a tagging pulse that modifies the magnetization of inflowing arterial spins, followed by a delay (TI) to allow for inflow, and a rapid image...

Specific ventilation (SV) is the ratio of fresh gas entering a lung region divided by its end-expiratory volume. To quantify the vertical (gravitationally dependent) gradient of SV in eight healthy supine subjects, we implemented a novel proton magnetic resonance imaging (MRI) method. Oxygen is used as a contrast agent, which in solution changes the longitudinal relaxation time (T1) in lung tissue. Thus alterations in the MR signal resulting from the regional rise in O(2) concentration following a sudden change in inspired O(2) reflect SV-lung units with higher SV reach a new equilibrium faster than those with lower SV. We acquired T1-weighted inversion recovery images of a sagittal slice of the supine right lung with a 1.5-T MRI system. Images were voluntarily respiratory gated at functional residual capacity; 20 images were acquired with the subject breathing air and 20 breathing 100% O(2), and this cycle was repeated five times. Expired tidal volume was measured simultaneously. The SV maps presented an average spatial fractal dimension of 1.13 ± 0.03. There was a vertical gradient in SV of 0.029 ± 0.012 cm(-1), with SV being highest in the dependent lung. Dividing the lung vertically into thirds showed a statistically significant difference in SV, with SV of 0.42 ± 0.14 (mean ± SD), 0.29 ± 0.10, and 0.24 ± 0.08 in the dependent, intermediate, and nondependent regions, respectively (all differences, P < 0.05). This vertical gradient in SV is consistent with the known gravitationally induced deformation of the lung resulting in greater lung expansion in the dependent lung with inspiration. This SV imaging technique can be used to quantify regional SV in the lung with proton MRI.

The temporal dynamics of blood flow in the human lung have been largely unexplored due to the lack of appropriate technology. Using the magnetic resonance imaging method of arterial spin labeling (ASL) with subject-gated breathing, we produced a dynamic series of flow-weighted images in a single sagittal slice of the right lung with a spatial resolution of ~1 cm(3) and a temporal resolution of ~10 s. The mean flow pattern determined from a set of reference images was removed to produce a time series of blood flow fluctuations. The fluctuation dispersion (FD), defined as the spatial standard deviation of each flow fluctuation map, was used to quantify the changes in distribution of flow in six healthy subjects in response to 100 breaths of hypoxia (FI(O(2)) = 0.125) or hyperoxia (FI(O(2)) = 1.0). Two reference frames were used in calculation, one determined from the initial set of images (FD(global)), and one determined from the mean of each corresponding baseline or challenge period (FD(local)). FD(local) thus represented changes in temporal variability as a result of intervention, whereas FD(global) encompasses both FD(local) and any generalized redistribution of flow associated with switching between two steady-state patterns. Hypoxic challenge resulted in a significant increase (96%, P < 0.001) in FD(global) from the normoxic control period and in FD(local) (46%, P = 0.0048), but there was no corresponding increase in spatial relative dispersion (spatial standard deviation of the images divided by the mean; 8%, not significant). There was a smaller increase in FD(global) in response to hyperoxia (47%, P = 0.0015) for the single slice, suggestive of a more general response of the pulmonary circulation to a change from normoxia to hyperoxia. These results clearly demonstrate a temporal change in the sampled distribution of pulmonary blood flow in response to hypoxia, which is not observed when considering only the relative dispersion of the spatial distribution.

This demonstrates a MR imaging method to measure the spatial distribution of pulmonary blood flow in healthy subjects during normoxia (inspired O2, fraction (FIO2) = 0.21) hypoxia (FIO2 = 0.125), and hyperoxia (FIO2 = 1.00). In addition, the physiological responses of the subject are monitored in the MR scan environment. MR images were obtained on a 1.5 T GE MRI scanner during a breath hold from a sagittal slice in the right lung at functional residual capacity. An arterial spin labeling sequence (ASL-FAIRER) was used to measure the spatial distribution of pulmonary blood flow 1,2 and a multi-echo fast gradient echo (mGRE) sequence 3 was used to quantify the regional proton (i.e. H2O) density, allowing the quantification of density-normalized perfusion for each voxel (milliliters blood per minute per gram lung tissue).With a pneumatic switching valve and facemask equipped with a 2-way non-rebreathing valve, different oxygen concentrations were introduced to the subject in the MR scanner through the inspired gas tubing. A metabolic cart collected expiratory gas via expiratory tubing. Mixed expiratory O2 and CO2 concentrations, oxygen consumption, carbon dioxide production, respiratory exchange ratio, respiratory frequency and tidal volume were measured. Heart rate and oxygen saturation were monitored using pulse-oximetry. Data obtained from a normal subject showed that, as expected, heart rate was higher in hypoxia (60 bpm) than during normoxia (51) or hyperoxia (50) and the arterial oxygen saturation (SpO2) was reduced during hypoxia to 86%. Mean ventilation was 8.31 L/min BTPS during hypoxia, 7.04 L/min during normoxia, and 6.64 L/min during hyperoxia. Tidal volume was 0.76 L during hypoxia, 0.69 L during normoxia, and 0.67 L during hyperoxia.Representative quantified ASL data showed that the mean density normalized perfusion was 8.86 ml/min/g during hypoxia, 8.26 ml/min/g during normoxia and 8.46 ml/min/g during hyperoxia, respectively. In this subject, the relative dispersion 4 , an index of global heterogeneity, was increased in hypoxia (1.07 during hypoxia, 0.85 during normoxia, and 0.87 during hyperoxia) while the fractal dimension (Ds), another index of heterogeneity reflecting vascular branching structure, was unchanged (1.24 during hypoxia, 1.26 during normoxia, and 1.26 during hyperoxia).Overview. This protocol will demonstrate the acquisition of data to measure the distribution of pulmonary perfusion noninvasively under conditions of normoxia, hypoxia, and hyperoxia using a magnetic resonance imaging technique known as arterial spin labeling (ASL).Rationale: Measurement of pulmonary blood flow and lung proton density using MR technique offers high spatial resolution images which can be quantified and the ability to perform repeated measurements under several different physiological conditions. In human studies, PET, SPECT, and CT are commonly used as the alternative techniques. However, these techniques involve exposure to ionizing radiation, and thus are not suitable for repeated measurements...

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