Gradient-echo (GRE) blood oxygen level-dependent (BOLD) effects have both intra-and extravascular contributions. To better understand the intravascular contribution in quantitative terms, the spin-echo (SE) and GRE transverse relaxation rates, R 2 and R * 2 , of isolated blood were measured as a function of oxygenation in a perfusion system. Over the normal oxygenation saturation range of blood between veins, capillaries, and arteries, the difference between these rates, R 2 ‫؍‬ R * 2 ؊ R 2 , ranged from 1.5 to 2.1 Hz at 1.5 T and from 26 to 36 Hz at 4.7 T. The blood data were used to calculate the expected intravascular BOLD effects for physiological oxygenation changes that are typical during visual activation. This modeling showed that intravascular ⌬R* 2 is caused mainly by R 2 relaxation changes, namely 85% and 78% at 1.5T and 4.7T, respectively. The simulations also show that at longer TEs (>70 ms), the intravascular contribution to the percentual BOLD change is smaller at high field than at low field, especially for GRE experiments. At shorter TE values, the opposite is the case. For pure parenchyma, the intravascular BOLD signal changes originate predominantly from venules for all TEs at low field and for short TEs at high field. At longer TEs at high field, the capillary contribution dominates. The possible influence of partial volume contributions with large vessels was also simulated, showing large (two-to threefold) increases in the total intravascular BOLD effect for both GRE and SE. Most BOLD-fMRI studies employ R* 2 contrast changes, which, compared to R 2 changes, are larger and take place both inside the vasculature and in the tissue surrounding the vessels (1,2). Hoogenraad et al. (3) recently provided an overview of these BOLD R* 2 relaxation processes and their roles in different types of experiments and brain regions. Intravascular contributions consist of intrinsic blood R* 2 relaxation, as well as additional dephasing due to susceptibility differences between blood and tissue, or changes in local blood velocity. Extravascular vessel-size-dependent effects due to vessel-tissue susceptibility differences are expected around both capillaries and larger vessels. Susceptibility-based dephasing mechanisms have been modeled (4 -6), and based on these theoretical predictions, most in vivo BOLD quantification approaches focus on extravascular R* 2 effects. However, experiments at low and intermediate field strengths have shown that the intravascular contribution of large vessels to the fMRI signal is dominant (2,7-10). When motional weighting with magnetic field gradients is employed (2,11), approximately 60 -70% of the R* 2 effect can be removed. Based on flow and diffusion considerations, such experiments remove most signals originating from inside large vessels and from arterioles and venules above 10 -20 m. The total BOLD R* 2 effect is also expected to have a substantial intravascular contribution from venules and capillaries, because R 2 effects in pure blood are very large (12)(13)(14). To b...

(1) have shown that this oxygenation-dependent contrast affects both T 2 and T * 2 in blood, while extravascular effects are predominantly related to T * 2 . In addition, it was recently demonstrated that macrovascular signal changes dominate in both gradient-echo (GRE) and spin-echo (SE) BOLD experiments at low (4 -6) and intermediate (7,8) field strengths. It has been suggested that this situation can be exploited as a means to directly determine venous blood oxygenation by measuring intravascular T 2 (9,10), T * 2 (4,11,12), or signal phase (13-15). These relaxation approaches find their origin in the literature data for pure blood, showing a direct relationship between oxygenation and transverse relaxation times (16 -21). We recently extended the well-known Luz-Meiboom equation (17,22,23), often used to describe blood relaxation, by expressing the erythrocyte relaxation rate (R 2,ery ) and the plasma-erythrocyte susceptibility shift difference in terms of the fractional concentration of deoxyhemoglobin, x deoxy (24,25). For venous blood, this deoxygenation fraction is determined by the oxygen saturation fraction (Y v ), which is a function of the arterial oxygen saturation fraction (Y a ) and the oxygen extraction ratio (OER) of the tissue from which the venous blood is draining (24,25):This OER, also called the oxygen extraction fraction (OEF), describes the balance between oxygen delivery and consumption in a tissue:in which [Hb tot ] is the total hemoglobin concentration in millimolar, CMR O2 the cerebral metabolic rate of oxygen in mol/g/min, and CBF the cerebral blood flow in ml/g/min. Thus, at constant hematocrit and arterial oxygenation, OER depends on the ratio of CMR O2 and CBF. Under these conditions, the intravascular venous BOLD effect on any physiological alteration (i.e., brain activation, hypoxia, hypercarbia) directly reflects the mismatch of changes in oxygen metabolism and blood flow (25). Using the theory outlined in the Materials and Methods section, we recently quantified absolute OER values in the human visual cortex during baseline activity and visual stimulation using T 2 measurements inside veins draining from the parenchyma (25). Quantification of absolute T 2 s was achieved from the echo time dependence of the venous signal intensity obtained using single spin-echo (SE) experiments. However, these so-called Hahn spin-echo experiments are time-consuming and have several disadvantages. First of all, the theory predicts that the apparent T 2 values measured with this kind of sequence are TE-dependent, and quantification of OER in the previous article could only be achieved by minimizing the difference between exponential signal-intensity fits through experimen-

In the present study blood T 1 was determined as a function of hematocrit and oxygen saturation. T 1 showed a significant linear dependency on both of these parameters. In addition, oxygen dissolved in blood plasma in hyperoxygenated blood resulted in relaxation enhancement, comparable in size to that due to the change in oxygenation state of hemoglobin. As blood T 1 is a key factor for quantification of flow with arterial spin labeling methods, the influence of T 1 variation in the physiological range of hematocrit and oxygen saturation to flow determination is discussed. Several physiologic variables are amenable to quantification by MRI in perfused tissue through contrast generated by T 2 , T * 2 , or T 1 in blood. A blood oxygenation leveldependent (BOLD) effect (1) on the transverse relaxation has recently been exploited for determination of oxygen extraction in the brain from T * 2 (2) or T 2 (3). Flow-related T 1 effects are exploited for quantification of perfusion using arterial spin labeling (ASL) methods. The models of cerebral hemodynamics used for blood flow quantification (4,5) generally consider that T 1 is a B 0 -dependent constant, independent of both blood oxygenation and hematocrit (Hct). However, in light of T 1 data from protein phantoms it appears conceivable that the blood longitudinal relaxation would be affected by these factors.The rate of energy transfer between spins and lattice in pure water is inefficient, resulting in low relaxation rates. The presence of macromolecules provides more relaxation pathways via a water-protein interaction, including exchange of water and protons as well as diffusion-mediated relaxation. The efficiency of energy transfer from water to proteins depends strongly on the motional state of the protein and, indeed, longitudinal relaxation in immobilized or cross-linked protein samples resembles that determined in vivo (6). It should be noted that the macromolecules of blood, especially albumin in the plasma and also hemoglobin inside erythrocytes, are quite mobile relative to cross-linked proteins, even if the viscosity slows down the intracellular motion in erythrocytes (7). Paramagnetic materials in the system provide an additional relaxation pathway through nonzero electronic spin of the solute. In the case of strongly paramagnetic contrast agents, like Gd-complexes, this effect is certainly dramatic, but it can also be detected with weakly paramagnetic material, such as molecular oxygen (8).In blood, there are several factors that may contribute to R 1 . First, increased macromolecule content is known to accelerate relaxation (9,10) in protein solutions. As intracellular hemoglobin (Hb) is the dominant blood macromolecule, alterations in blood Hct (i.e., erythrocyte volume fraction) are expected to influence the relaxation rate (10). In vivo the physiologic macrovascular Hct of individuals varies from 0.34 to 0.50 and, furthermore, microvascular Hct in the brain is only 85% of that in macrovasculature (11). Second, deoxygenated Hb is moderately paramagnet...

The effects of intracellular pH (pH i ), paramagnetic macroscopic, and microscopic susceptibility on T 1 in the rotating frame (T 1 ) were studied in rat brain. Intracellular acidosis was induced by hypercapnia and pH i , T 1 , T 2 , diffusion, and cerebral blood volume (CBV) were quantified. Taking into account the CBV contribution, a prolongation of parenchymal T 1 by 4.5% was ascribed to a change in tissue water relaxation caused by a one unit drop in pH i . Blood T 1 was found to prolong linearly with blood oxygenation saturation (Y). The macroscopic susceptibility contribution to parenchymal T 1 was assessed both through BOLD and an iron oxide contrast agent, AMI-227. The T 1 data from these experiments could be described by intravascular effects with insignificant effects of susceptibility gradients on tissue water. Tissue oxygen tension (PtO 2 ) was manipulated and monitored with microelectrodes to assess its plausible contribution to microscopic susceptibility and relaxation. Parenchymal Acute ischemia results in subtle and almost instantaneous changes in cerebral MR relaxation times, including prolongation of T 1 (1,2) and T 1 in the rotating frame (T 1 ) (3,4), shortening of T* 2 (5), and T 2 (1,6,7). These relaxation changes are expressed in global forebrain ischemia when there is no possibility for edema formation (2) and also almost immediately after the occlusion of the middle cerebral artery (3,8). The time-dependency of these relaxation changes reveals that net water accumulation to the ischemic brain tissue is not responsible for the depicted MR changes. Early T* 2 and T 2 reductions in ischemic brain have been ascribed to local increases in the level of deoxyhemoglobin, i.e., a blood oxygenation level-dependent (BOLD) phenomenon (5,6,9) and, furthermore, T 2 effects have been quantitatively modeled as intravascular susceptibility effects reflecting elevated oxygen extraction in the flow-compromised tissue (7). The presence of negative BOLD in flow-compromised brain is an index of ongoing oxidative metabolism; thus, it can reveal potentially viable tissue (10).While the transverse relaxation changes in the early moments of ischemia can be explained by both intra-and extravascular susceptibility effects, the longitudinal relaxation alterations might have a different pathophysiological background. From a general principle, tissue R 1 relaxation rate (R 1,obs ) has been proposed to receive contributions from the following physical factors:where subscripts exch, rot, and diff indicate exchange, rotational, and diffusion contributions to R 1 , respectively, and b is the fraction of bound water (11). The same relation has commonly been regarded as being applicable throughout the resonance frequency range, thus also including T 1 relaxation in the rotating frame (11,12). Knispel et al. (11) quantitatively explored T 1 relaxation and relaxation dispersion, i.e., T 1 as a function of B 1 , in tissue phantoms and they concluded that three distinct correlation times, falling into the three physical categor...

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T 2 of cortical gray matter is generally assumed to be longer than that of white matter. It is shown here that this is not the case in the occipital lobe, but that this effect is often obscured at lower resolution and concealed in standard T 2 -weighted images. Using a high-resolution (1 ؋ 1.3 ؋ 2 mm 3 ) segmented EPI CarrPurcell-Meiboom-Gill sequence, T 2 relaxation times of the brain were measured at 1.5 T for eight healthy adult volunteers. The average T 2 values of cortical gray and white matter were found to be 88 ؎ 2 and 84 ؎ 3 msec in the frontal lobe, 84 ؎ 2 and 83 ؎ 3 msec in the parietal lobe, and 79 ؎ 1 and 87 ؎ 3 msec in the occipital lobe, respectively. This unexpected occipital T 2 contrast between gray and white matter is attributed to regional differences in iron concentration. The transverse relaxation time T 2 is one of the most commonly employed MRI contrast parameters. We have recently shown that absolute T 2 values can be used to quantify cerebral hemodynamics and oxygen extraction (1). As such, it is useful to accurately map T 2 of various normal and abnormal tissues. Several studies (2-11) have shown that T 2 varies across brain regions and during different stages of brain maturation and aging (5,10). Numerous studies in normal adults at multiple field strengths (2-4,9) have concluded that T 2 of cortical gray matter (GM) is longer than that of white matter (WM). On the contrary, gray matter T 2 was found to be shorter than that of WM in the brains of newborns (10) and in spinal cord (12). Several factors are likely to contribute to the heterogeneity of T 2 values. Among these may be regional variations in brain iron concentration, as well as variations in the microarchitecture of the cellular constituents and myelin structure (cytoarchitecture and myeloarchitecture, respectively). Recently, investigators examining the superior temporal gyrus and primary auditory cortex region have shown that the T 2 -weighted signal intensities can vary appreciably even between neighboring regions of the cortex (11). We found a similar pattern in our studies of the visual cortex, where typically an oblique slice through the calcarine fissure is studied. In order to better quantify the T 2 contrast between GM and WM in this location, we used very-high spatial resolution absolute T 2 measurements to minimize the likelihood of including WM and CSF (cerebrospinal fluid) in voxels nominally occupied by GM. MATERIALS AND METHODSMeasurements were performed on a 1.5 T ACS-NT clinical MRI scanner (Philips Medical Systems, Best, Netherlands). A body coil was used for RF transmission and a head coil was used for signal reception. Eight normal adult volunteers (four male, four female, age range 22-44 years, mean 32 Ϯ 8 SD years) were studied in accordance with institutional guidelines. A multiecho Carr-Purcell-Meiboom-Gill (CPMG) pulse sequence with 13-shot segmented EPI acquisition was utilized at an echo-spacing of 25 msec. A single slice was acquired with 30 averages at the level of the dorsal-most aspects of the ...

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