Cerebral blood flow (“CBF”) is among the most important physiological variables in understanding local brain metabolism. CBF provides the physiological basis for blood oxygen level dependent (BOLD) contrast, the most frequently used imaging technique for estimating changes in neural activation1 (all references cited to in this application are listed in Appendix A hereto, which is incorporated herein by reference). In addition to its fluctuations during normal brain function, CBF also changes during many of the pathological events that lead to acute or chronic brain dysfunction. Because these pathological changes typically have an earlier onset than structural biomarkers, measurement of CBF can be a useful tool in diagnosing stroke, ischemia, brain tumors, and dementia.
The earliest measurements of CBF used radioactive tracers.2 In fact, positron emission tomography (“PET”) measurements using radioisotopes such as O15 are still regarded as the most accurate3. However, the invasiveness and expense of radioactive tracers significantly inhibit the widespread use of PET for CBF measurement.
Developments in MRI have led to an alternate method for measuring CBF, one using dynamic contrast enhancement following a rapid bolus injection of MR contrast agent. Although less invasive and expensive than radioactive tracers, this technique is still problematic, particularly in research settings. The need for venipuncture discourages some potential subjects, while others are excluded due to contraindications of the MRI contrast agent.
Arterial Spin Labeling (“ASL”) techniques replace the need for an injected contrast agent by using water protons in the plasma as an endogenous MRI tracer. ASL is both non-invasive and cost effective, and has higher spatial resolution than PET. Additionally, its clinical utility is enhanced by the fact that ASL images can be routinely acquired during the same imaging session as structural or other MRI scans and can be directly compared with the anatomical and pathological features they reveal.
In ASL an endogenous tracer (arterial water) is used instead of an exogenous tracer. Flowing spins are inverted (labeled) at a plane in the main arteries which is proximal to an imaging volume. The labeled spins then flow through the arterial tree into the capillary bed arriving at a particular location in the tissue with a distribution of arrival or transit times. Once in the capillary bed, the arterial water molecules then exchange, one for one, with the extravascular water molecules. Due to the accumulation of inverted spins in the extravascular component of the tissue there is a reduction in the total magnetization of the tissue. This, in turn, causes a reduction in the imaged signal in that region of the imaged plane. The reduction in signal of the tissue in the imaged plane is thus proportional to the amount of labeled spin which flowed into that plane.
Using ASL, in general CBF can be measured by the subtraction of two images. The first is an unlabeled or control image in which there has been no inversion of the spins. The subsequent image is labeled, i.e., the spins have been inverted in an arterial plane proximal to the imaged slice. The resultant difference between the control and labeled images is directly proportional to the flow.
Two common methods of conducting ASL are Continuous ASL (“CASL”) and Pulsed ASL (“PASL”). In PASL, spatially broad inversion pulses can be used to invert all of the spins in a region next to the slice of interest and the exchange between the inverted region and the imaged slice can be subsequently observed.9 In most cases PASL measures the ratio of CBF at various locations because it is difficult to precisely define the amount of inflowing spins.
In CASL, incoming arterial water can be continuously inverted by an adiabatic inversion pulse. Although CASL can be used to measure CBF, this process is problematic in that it relies on certain assumptions regarding the transit times and the relative T1's of blood and tissue to do so. If an image is acquired rapidly following an RF irradiation the labeled arterial water will dominate the difference signal such that any signal reduction from the water molecules that have exchanged into the tissue will not be visible. While this can be addressed by imposing a post labeling delay (PLD), for this to be successful the PLD needs to be longer than the longest transit time. This tends to make PLD's long, which results in lesser signal detection inasmuch as the difference signal continually decreases as a function of PLD length.
In CASL, a separate acquisition is needed to obtain a control image. Flow can then be calculated from the difference between the control image and the labeled images. Transit times between the labeling and imaging planes are not measured in CASL. Nonetheless, transit time is an important parameter in the calculation of quantitative CBF.14 To minimize the role of this unknown variable in flow quantification, post-labeling delays (PLD) between the end of labeling and the start of acquisition have been utilized, as noted. A PLD allows unlabeled spins which start flowing into the arterial tree sufficient time to wash out the labeled spins so that any labeled spins left in the arteries will not be confused with those that have exchanged into the tissue. Thus, if the PLD is longer than the transit time, the observed results are independent of transit time provided that the T1's in the blood and the tissue are the same, which is approximately true for gray matter. By varying a PLD one can investigate the effects of various transit times since an image acquired after a given PLD corresponds to integration over all longer transit times.15 
However, the insertion of a PLD has several disadvantages. First, it is costly in SNR terms. Furthermore, it is vulnerable to changes in transit during activation as well as the potentially longer transit times that often occur in stroke and other cerebral vascular diseases (which can exceed the PLD). For example, a standard PLD sufficiently small not to degrade SNR beyond acceptable limits is 1 to 2 seconds, and the transit times common in stroke and other cerebral vascular diseases can be as large as 4 to 5 seconds.
In white matter, however, quantification of CBF using CASL encounters additional unique difficulties. The use of a PLD for measuring CBF in gray matter depends upon the assumption that the T1gray=T1blood, which, while acceptable for gray matter, is not valid for white matter. CASL also suffers from the lengthy transit time required for blood to reach white matter after being labeled in the artery. The relaxation of the labeled spins during this transit results in a much lower SNR for white matter relative to gray matter.
In ASL, the spins of protons in arterial water are typically inverted (labeled) at a plane in a cerebral artery. CBF then can be quantified by measuring the effect of the labeled spin when the protons enter a nearby imaging plane. However, when the transit time between the labeling and imaging planes cannot be measured, a major uncertainty is introduced into this calculation—it is difficult to discern how much of the detected signal is due to label accumulated in the tissue or label remaining in the arterial space in the tissue. While CASL attempts to reduce this uncertainty by introducing a PLD between the end of the labeling pulse and the start of data acquisition, this solution is only partially effective; there is uncertainty regarding the relative relaxation times, and it imposes an approximately three-fold cost in signal-to-noise ratio due to the longitudinal relaxation of the label during the PLD interval.
Moreover, current ASL methods use a continuous or square waveform for spin labeling. This can be difficult for the RF amplifiers on most commercial MR scanners to generate because they are generally designed for pulsed, as opposed to continuous, applications.
Finally, CASL uses lengthy (e.g., 1-2 second) RF pulses for spin labeling. The length of such pulses tends to place strain on the RF transmitters in most MRI scanners that is typically relieved through the use of transmit-receive head coils. However, receive-only coils offer greater sensitivity and are more widely available making this solution non-optimal. Additionally, long RF pulses introduce confounds from magnetization transfer (MT) effects whereby the spins in the imaged plane are saturated by the long RF pulse. To address this, CASL requires that the acquisition of a control image be done in the presence of a RF pulse which mimics the MT effects of labeled images but does not invert the arterial spins. However, while this approach eliminates MT effects, it introduces additional sources of error, particularly motion artifacts.
Thus, what is needed in the art is a method of blood flow and perfusion measurement that can overcome the problems of conventional methods, and that can allow for the direct measurement of transit time distribution while acquiring ASL data.
What is further needed in the art is a method of blood flow and perfusion measurement that can add certainty to CBF quantification and obviate the need for PLD's control image acquisition or reliance on less sensitive signal receiving hardware.