In Vivo ¹⁷O MRS Imaging – Quantitative Assessment of Regional Oxygen Consumption and Perfusion Rates in Living Brain

Abstract

In the last decade, in vivo oxygen-17 (¹⁷O) MRS has evolved into a promising MR technique for noninvasively studying oxygen metabolism and perfusion in aerobic organs with the capability of imaging the regional metabolic rate of oxygen and its changes. In this chapter, we will briefly review the methodology of the in vivo ¹⁷O MRS technique and its recent development and applications; we will also discuss the advantages of the high/ultrahigh magnetic field for ¹⁷O MR detection, as well as the challenges and potential of this unique MRS method for biomedical research of oxygen metabolism, mitochondrial function and tissue energetics in health and disease.

Introduction

Compared to well-established in vivo magnetic resonance (MR) spectroscopy (MRS) methodologies such as phosphrous-31 (³¹P) or carbon-13 (¹³C) MRS [1], in vivo oxygen-17 (¹⁷O) MRS has had a relatively short history and is less commonly employed in biomedical research. As one of the most abundant elements on earth, molecules containing oxygen exist in all forms and levels of life. Thus, efforts to develop and utilize ¹⁷O-based MR technology to obtain valuable information of biological systems are ongoing. Perhaps the most intriguing use of in vivo ¹⁷O MRS is to study the cellular oxygen metabolism in living organs. The original idea to use ¹⁷O MRS for in vivo studies appeared in the late 1980s [2–7], and it was likely inspired by the ¹⁵O-based positron emission tomography (PET) technique developed for imaging the cerebral metabolic rate of oxygen (CMRO₂), a critically important physiological parameter of the brain [8, 9].

Among all neuroimaging modalities capable of providing CMRO₂ information, ¹⁵O-PET has been regarded as a gold standard for directly imaging the oxygen metabolic rate in the living brain. However, two major drawbacks associated with ¹⁵O-PET imaging have seriously limited its availability and applicability: one is the very short half-life of ¹⁵O (2.04 min), which requires expensive equipment such as a cyclotron to produce the ¹⁵O-tracers on site; and the other is the inability of the ¹⁵O-PET in distinguishing the ¹⁵O signals of the oxygen substrate from that of metabolically produced water, so additional measurements and complicated mathematical models are needed to calculate the CMRO₂ values [8, 9]. Magnetic resonance technology is a widely used neuroimaging modality in present-day medical research and clinical practice. ¹⁷O MRS (or imaging) and ¹H MRI-based approaches have been developed for assessing CMRO₂. Even though the ¹⁷O-MR approach is technically much simpler than ¹⁵O-PET since it only detects the metabolically generated water signal (see details below) and does not require the injection of large radiation doses to the subject, its initial utilization was slow, with only a small number of publications in the first decade after inception that mainly focused on feasibility assessment [10–16]. The major challenge in advancing this technology can be attributed to the extremely low intrinsic sensitivity of ¹⁷O as compared to that of ¹H (~1×10⁵ fold differences) at natural abundance [17]. On the other hand, ¹H MRI detects abundant water signals with great sensitivity, although the water signals in general cannot be directly related to oxygen metabolism. To specifically assess cerebral oxygenation, several methods based on ¹H MRI have been introduced [18–23]. They include (i) the blood oxygenation level dependent (BOLD) functional magnetic resonance imaging (fMRI) in combination with hypercapnic and hyperoxic respiratory challenges to calibrate the fMRI signals [19, 21]; (ii) susceptometry-based or T₂-based methods to determine the cerebral venous oxygen saturation (Yv) [22, 23]; and (iii) a quantitative BOLD approach to extract the venous cerebral blood volume (CBV) and deoxyhemoglobin concentration from the transverse relaxation time (T₂ or T₂^*) [18, 20]. However, these indirect methods have their own technical challenges and limitations; and various assumptions, predetermined parameters and different models must be applied in estimating CMRO₂.

High/ultrahigh field MR scanner technology emerged in the 1990s and the extraordinary sensitivity gain of ¹⁷O detection at higher field observed in 2001 have stimulated new interests and efforts in the field of in vivo ¹⁷O MRS research [24, 25]. After more than ten years of technology development and validation, the in vivo ¹⁷O MRS/MRI methodology has now been established as a valuable tool for studying cerebral oxygen metabolism in preclinical animal models and in human brains under various physiopathological conditions [25–39]. Details of the technology development can be found in several earlier review articles [40–43]. In the present review, we will briefly describe the methodology and its utility for non-invasive investigation and quantitative understanding of the roles of oxygen metabolism in health and disease. The main scope of this review will be on the new advancements and recent applications in this research field.

Methodology Overview

Among the three stable isotopes of oxygen (i.e., ¹⁶O, ¹⁷O and ¹⁸O), ¹⁶O represents over 99.7% of the composition, but unfortunately does not possess a nuclear spin; only ¹⁷O with a spin number of I=5/2 can be detected by MR. In this section, we will introduce different in vivo ¹⁷O MR methods according to their signal sources and/or related uses with either a spectroscopic or imaging approach,.

Steady-state detection of the natural abundance ¹⁷O signal

The relative amount of ¹⁷O in oxygen in nature is very rare (only 0.037% of total oxygen). Because the concentration of major metabolites is in the range of few millimolar to tens millimolar, long ¹⁷O MR acquisition times are required to accumulate reasonable signal for meaningful detection or adequate signal-to-noise ratio (SNR) of oxygen-containing metabolites in tissue samples. Figure 1 illustrates natural abundance in vivo ¹⁷O MR spectra obtained in live and postmortem rat brains with a 11.7T scanner that required several million scans over 6 and 14 hours, respectively, in which a number of broad resonances from oxygen in phosphate, sulfonic and carbonyl groups covering a wide chemical shift range are detected [44]. It is clear that the SNR of these natural abundance ¹⁷O signals is insufficient for practical applications aiming to monitor oxygen-containing metabolites. The only molecule that can be easily detected by in vivo ¹⁷O MRS techniques is water, a major component of the tissue. However, with known water content and molar concentration in different tissues, the natural abundance water signal in ¹⁷O MR spectra can serve as an internal reference for quantification purposes. This is an important advantage of the in vivo ¹⁷O MR methodology and is crucial for its application.

Figure 1.

Pulse-acquired natural abundance ¹⁷O MR spectra of rat brain: live (top, 1.5 million averages) and post mortem (bottom, 3.5 million averages) brains obtained with 6 and 14 hours data acquisition, respectively (chemical shifts are referenced to the water signal at 0 ppm). Adapted from reference [44]. Copyright © 2008 Elsevier Inc.

Dynamic monitoring of the metabolism of ¹⁷O-labled oxygen gas

A unique feature of the in vivo ¹⁷O MR technique is its ability to directly and quantitatively monitor the production of metabolic water after the introduction of ¹⁷O-labeled oxygen (¹⁷O₂) gas to an animal or human. It is well known that when oxygen ¹⁷O₂ molecules enter the blood stream, they are brought to the organ or tissue of interest via effective blood circulation and perfusion. At the cellular level, the oxygen consumption occurs in the mitochondria through cellular respiration that generate ¹⁷O-labeled water (H₂¹⁷O); the rate of H₂¹⁷O production reflects the oxygen consumption rate. Since the ¹⁷O₂ molecules that are either bound to hemoglobin or dissolved in blood are invisible to ¹⁷O MR detection, only H₂¹⁷O molecules, the final product of the oxygen metabolism, has an ¹⁷O signal; therefore, dynamic in vivo ¹⁷O MR measurements can be used to monitor the production of H₂¹⁷O and determine the metabolic rate of oxygen in the targeted tissue or organ. The principle of this technique, the quantification model and corresponding mass balance equation, as well as the simplified method for determining CMRO₂ and cerebral blood flow (CBF) in small animal brains via a brief (few minutes) inhalation of ¹⁷O₂ gas, are schematically illustrated in Figure 2; the detailed methodology and its validity, which has been experimentally verified in rat brains, can be found elsewhere [26, 29, 31, 34, 38–42].

Figure 2.

Schematic illustration of in vivo ¹⁷O MR technique for quantifying CMRO₂ and CBF in rodents via a brief inhalation of ¹⁷O-labeled oxygen gas: the principle (a), quantification model and mass balance equation (b), as well as the simplified method for calculating CMRO₂ and CBF based on the brain H₂¹⁷O time course (c) are displayed. Detailed description and explanation of the method can be found in references [38–39 & 40–43].

Blood flow measurement with exogenous tracer of H₂¹⁷O

Introducing ¹⁷O-labeled water as an exogenous tracer and using dynamic in vivo ¹⁷O MR technique to monitor the evolution of the H₂¹⁷O signal in targeted tissue or organ enables the determination of the washout rate of the H₂¹⁷O tracer, which represents the blood flow or perfusion rate [26, 31]. This approach is the same as conventional tracer techniques used in biomedical research or in the clinic. The ideal method of introducing H₂¹⁷O tracer is via a bolus injection to a feeding artery (e.g. the internal carotid artery) as shown in Figure 3 for CBF measurement, which is straightforward for CBF quantification but requires invasive procedures to gain access to the desired artery [26, 31]. Alternatively, intravenous injection is less invasive and more common for human application, but the CBF quantification is more complicated and requires the tracer input function.

Figure 3.

¹⁷O-MR based CBF measurement: (A) stacked plots of the cerebral H₂¹⁷O spectra after a bolus injection of H₂¹⁷O into an internal carotid artery of a rat; (B) exponential decay fitting of the H₂¹⁷O washout curve for calculating CBF. Adapted from reference [26]. Copyright © 2002, National Academy of Sciences, U.S.A.

The use of metabolic H₂¹⁷O as an endogenous tracer for CBF measurement

The ¹⁷O MR signal of the metabolic water generated in the mitochondria following a ¹⁷O₂ inhalation can serve as a perfusion tracer, and its decay rate (k) can be used to determine the CBF value with an experimentally derived correction factor (i.e., CBF ≈ 1.86 × k). This empirical approach was demonstrated in rat brains across a wide range of conditions [39]. This finding suggests that the dynamic in vivo ¹⁷O MR technique is capable of simultaneously and noninvasively determining both CMRO₂ and CBF as schematically shown in Figure 2. Thus, ¹⁷O MR provides a new method for studying the cerebral oxygen metabolism and perfusion underlying brain function and disease. In addition, the oxygen extraction fraction (OEF) that reflects the balance between oxygen supply and usage can be calculated from the corresponding CMRO₂ and CBF values using the relationship of OEF = CMRO₂/(C_a,O2×CBF), where C_a,O2 is a constant representing the arterial oxygen concentration of the brain [38]. Therefore, by combining the dynamic in vivo ¹⁷O MR technique with a short ¹⁷O₂ gas inhalation, it is possible to quantify three important physiological parameters, namely CMRO₂, CBF and OEF in a completely non-invasive manner. A similar approach and quantification model in human brain has recently been developed [45].

In vivo ¹⁷O MRS imaging (or MRI)

Another important feature of the in vivo ¹⁷O MR technology is its capability of detecting not only the dynamic change of the ¹⁷O signal but also the signal distribution in space via three-dimensional (3D) ¹⁷O MRS imaging [41, 42] or MRI approaches, although conventional MRI sequences are not applicable due to the extremely short T₂ relaxation time of the H₂¹⁷O water (in the range of few milliseconds) [25, 28, 46, 47]. Thus, with this technique, it is possible to assess regional oxygen consumption and perfusion rates in the living brain or in other organs. Furthermore, after introducing an exogenous ¹⁷O-tracer, the ¹⁷O signal can reach a new steady state within a reasonable time window, e.g. ~10–20 minutes in the brain, which permits repeated measurements with subsequent tracer administration under different conditions [34]. This approach enables the quantification of the same parameter with absolute values at different states, as well as the changes due to altered physiological conditions or brain states [34, 35, 48].

The aforementioned methodological aspects of the in vivo ¹⁷O MR technology reveal its unique utility for biomedical research in the area of oxygen metabolism, energetics and vascular-metabolic relationships in a living organ. It should be noted that the spatial and temporal resolution of dynamic ¹⁷O MR imaging are dependent upon the available ¹⁷O signal; thus, ¹⁷O sensitivity is the key limiting factor that determines the spatial/temporal resolution and the reliability of the ¹⁷O MR imaging.

Advantage of high/ultrahigh fields for in vivo ¹⁷O MR imaging

¹⁷O has a roughly 7 times lower gyromagnetic ratio and 2700 times lower natural abundance than ¹H, which leads to 10⁵ times lower sensitivity compared to ¹H [17, 42]. The advantage of the high/ultrahigh magnetic field strength for improving the sensitivity of the ¹⁷O MR imaging is essential for its potential applications. Such an advantage stems from the quadrupolar relaxation mechanism of the ¹⁷O-water. The nearly field-independent relaxation times of the ¹⁷O-water eliminates the potential signal loss commonly occurring in the ¹H MRS or MRI at higher field because increasing field strength will not prolong the longitudinal relaxation time (T₁) and/or shorten the T₂ (or T₂^*) of the ¹⁷O-water. Thus, the acquisition efficiency of the ¹⁷O MR signal will not be compromised [25, 28, 46, 47]. The extremely short T₁ of the ¹⁷O-water permits fast acquisition and more signal averaging in unit sample time, thus enhancing the apparent ¹⁷O SNR. Also, the effect of the magnetic field (B₀) inhomogeneity on the ¹⁷O-water linewidth, which is inversely proportional to the T₂^*, is very small (negligible) compared to the effect of the quadrupolar relaxation [25, 28, 46, 47]. Therefore, in vivo ¹⁷O MR measurement is virtually insensitive to the B₀ shimming, a challenging procedure faced by many other in vivo MRS methodologies.

The field-dependent relationship between the apparent ¹⁷O-SNR (normalized to that of 9.4T) and the field strength across a wide B₀ range from zero up to 16.4T that was obtained from rat brains is displayed in Figure 4. The reported SNR increase with field strength is based on the empirical relationship of SNR ≈ C × B₀^β with C = 0.015 and β =1.9. This value of β is slightly higher than the theoretical prediction (β =1.75) based on the work of Hoult et al. [49]; this could be due to imperfect control of the experimental variables (e.g. the Q factor (quality factor) of the radio frequency (RF) coils is not identical at different fields) used in the two studies. The finding of field-dependent ¹⁷O-SNR enhancement suggests that compared to in vivo ¹H, ³¹P and ¹³C MRS methodologies, the in vivo ¹⁷O MRS can benefit the most from the sensitivity gain at high/ultrahigh field strengths [46].

Figure 4.

Field dependence of the relative ¹⁷O SNR in the rat brains (filled circles) normalized to the SNR value measured at 9.4. The equation represents the fitting result (solid line) of the experimental data covering the B₀ fields of zero to 16.4 T. Adapt from reference [46]. Copyright © 2012 Wiley Periodicals, Inc.

Recent development and applications of in vivo ¹⁷O MRS and imaging

Preclinical study in animal brain

As noted above, currently, the in vivo ¹⁷O MR methodology has been established for the simultaneous dynamic imaging of three important physiological parameters, namely CMRO₂, CBF and OEF in preclinical animal brain. An example of such a study is mouse brain, in which the animals have undergone transit middle cerebral arterial occlusion (tMCAo) (Figure 5). The figure displays anatomic regions of post-stroke brains with voxels located in ischemic and intact hemispheres, corresponding to dynamic H₂¹⁷O signals before, during and after 2-min ¹⁷O₂ inhalation; reproducibility between two repeated imaging measurements in the same animal; and the results of CMRO₂, CBF and OEF imaging indicate significant changes in the ischemia affected brain region [38]. This study demonstrates a promising ¹⁷O-MR based imaging technique with the merits of robustness, simplicity, noninvasiveness and reliability, features that are essential for imaging abnormal oxygen metabolism and perfusion in diseased brains.

Figure 5.

In vivo ¹⁷O MR application in stroke mice: anatomic image with ischemic and intact voxels identified; corresponding dynamic H₂¹⁷O signals before, during and after a 2-min ¹⁷O₂ inhalation in a mouse model of tMCAo; reproducibility of repeated imaging measurements; and CMRO₂, CBF and OEF maps indicating significant changes in ischemia affected brain region. Adapted from reference [38]. Copyright © 2012 Elsevier Inc.

Several animal species commonly used in preclinical studies have been tested, which include mouse, rat, cat and swine at field strengths from 3T up to 16.4T; the results confirm the feasibility of dynamic in vivo ¹⁷O MRS imaging for studying oxygen metabolism in these preclinical models [35, 38, 39, 48, 50–52]. This novel ¹⁷O-MR based CMRO₂ imaging technique has been successfully applied to Huntington disease mouse brains and used as a metabolic imaging tool to assess the mitochondrial function in diseased brain in vivo by directly determining the cerebral oxygen consumption deficit under metabolic stress [51]. It is expected that more studies of this kind will be forthcoming once the research community recognizes the promise of this new technology.

Translational study in human brain

The dynamic in vivo ¹⁷O MRS imaging approach can be easily applied to the human brain to monitor the change of H₂¹⁷O signals during and after a short ¹⁷O₂ inhalation [32]. However, to calculate the oxygen consumption rate from the dynamic H₂¹⁷O signals, the simplified model used in small animals is not suitable due to the slower exchange of the ¹⁷O-labeled oxygen gas in the human body.

A three-phase model has been proposed, in which the time-dependent function of the ¹⁷O-labeled oxygen gas fraction in arterial blood is modeled based on an estimated replacement rate of fresh ¹⁷O-oxygen in the blood. In addition, two rate constants of K_L and K_G representing the loss and gain of H₂¹⁷O within the imaging voxel, respectively, are included in the proposed quantification model, whereas the CBF is not presented in the model since its influence on the measured signal is accounted for by the rate constants K_L and K_G [36]. Figure 6 shows results of human brain CMRO₂ mapping at 9.4T and two representative H₂¹⁷O time-courses in voxels with low and high metabolic rates. The model provides reasonable fitting to the ¹⁷O MR data yielding regional CMRO₂ values with corresponding K_G and K_L constants [36].

Figure 6.

¹⁷O MR based CMRO₂ imaging in human brain: CMRO₂ maps in mol/g brain/min unit (top) and representative H₂¹⁷O time-courses for single voxels corresponding to low (bottom left, white matter) and high (bottom right, gray matter) metabolic rates. The three-phase metabolic model of water production (thick line) accurately describes the ¹⁷O MR data (thin line) to yield CMRO₂ values with K_G and K_L constants shown. Adapted from reference [36]. Copyright © 2010 Elsevier Inc.

Recently, a modified quantification model was proposed for human application. In this model, the exchange function of the ¹⁷O-labeled oxygen in the human lung is accounted for by its influence on the fractional change of ¹⁷O₂ in the arterial blood. The ¹⁷O₂ gas exchange rate can be experimentally determined from a simple breathing test using non-labeled oxygen gas [45, 53]. This additional step is necessary for short ¹⁷O₂ inhalation because the transition of the oxygen gas in human lung takes ~2 min or longer to reach a steady state. More importantly, not only is it possible to obtain the CMRO₂ value, but the CBF value can also be obtained from the model fitting of the dynamic H₂¹⁷O signals in the human brain tissue. Consequently, all three parameters of CMRO₂, CBF and OEF can be simultaneously determined from the non-invasive ¹⁷O MRS imaging measurement in humans with a ¹⁷O₂ inhalation as short as 2–3 min [45, 53].

Feasibility assessments in normal human brain at resting or functional stimulated state, and in diseased human brains were performed at a field strength of 3T up to 9.4T; the results are encouraging, indicating that the in vivo ¹⁷O MRS imaging technique has the potential to become a promising neuroimaging modality for the translational study of oxygen metabolism in humans [45, 53–56]. It is worth pointing out, however, that future prospects of this methodology will be largely affected by the challenges encountered in expanding this research field for routine applications at clinical settings (see more discussion later).

Study of oxygen metabolism in extracerebral tissues

Besides the brain, the heart is another highly aerobic organ consuming a large amount of oxygen since the oxygen metabolism is essential for supporting the mechanical work of myocyte contractions. The feasibility of applying the in vivo ¹⁷O MR technique to hearts has been tested in rats at 9.4T [37, 57] and in human hearts at 3T [54]. The preliminary results as displayed in Figure 7 suggest that it is feasible to establish a noninvasive imaging modality with in vivo ¹⁷O MR detection and appropriate quantification modeling for simultaneously imaging the myocardial oxygen metabolic rate (MVO₂) and myocardial perfusion [37].

Figure 7.

In vivo ¹⁷O MR application in heart: (a) ¹H MRI of rat heart in axial (left panel) and sagittal (right panel) orientation showing the ¹⁷O surface coil location; (b) stack plot of the heart ¹⁷O-water signals acquired before, during (grey bar) and after the inhalation of ¹⁷O₂ gas.

Challenges and the potential of the in vivo ¹⁷O MRS and imaging

The most important application of the in vivo ¹⁷O MRS is to non-invasively study the oxygen metabolism in living animals or humans with introduction of ¹⁷O₂ gas. A major obstacle in this regard is the relatively high cost due to the limited supply of the ¹⁷O-enriched oxygen gas. Currently, only few vendors can provide ¹⁷O-isotope-labeled oxygen gas in large quantities with high enrichment, while the demand for this oxygen isotope is also limited. Increasing the demand may stimulate ¹⁷O-oxygen production and lower its cost; but at present, saving the cost by reducing or improving the efficiency of the ¹⁷O-oxygen gas usage seems to be the main focus, which is particularly crucial for human studies. Rebreathing pure ¹⁷O₂ with the CO₂ removed from the expired gas [36], and using a demand oxygen delivery system to reduce the total ¹⁷O₂ gas supply [55, 56] and/or recycling the expired ¹⁷O₂ gas for animal studies [32, 45, 53] are a few approaches being explored for this purpose. An alternative approach that administers ¹⁷O₂-enriched blood substitutes instead of ¹⁷O₂ inhalation has been suggested with the assumption that the high oxygen affinity of the artificial blood would improve its delivery efficiency [43, 58]; although the effectiveness and ability of this approach for accurately quantifying the metabolic rate of oxygen is yet to be determined. On the other hand, minimizing the ¹⁷O₂ gas inhalation time is a more effective way to lower the cost, although it requires a higher temporal resolution of the dynamic ¹⁷O measurement.

Limited detection sensitivity is another major challenge encountered in the in vivo ¹⁷O MR study. This is particularly critical for dynamic ¹⁷O MR imaging since both spatial and temporal resolutions are key factors demanding high ¹⁷O sensitivity. High/ultrahigh field scanners are employed in the ¹⁷O MRS imaging studies due to the advantages in gain of sensitivity. It has been shown that a few millimeter spatial resolution in preclinical studies [26, 34, 35, 37–39, 48, 50, 51] and close to a centimeter in human studies [32, 45, 53] with ~10–15 sec temporal resolution can be reached with B₀ ≥ 7T. Relatively lower resolutions are expected with lower field magnets, e.g., a 3T clinical scanner, or when large volume coils are employed [30, 36, 54, 56]. It is worth mentioning that while ¹⁷O MR benefits from the sensitivity gain at a higher field, it also requires a larger RF power to achieve the same RF pulse flip angle because of its low γ-ratio. Interestingly, a recent study has found that the high RF power demand in the ¹⁷O MR measurement is significantly compensated by its high B₁ efficiency as compared to that of ¹H MR at high/ultrahigh fields [59]. Of course, for human application, the specific absorption rate (SAR) has to be closely monitored to ensure it is below the FDA guideline, which could limit our ability to optimize the ¹⁷O MR acquisition parameters in the human studies.

Despite the limitation in the ¹⁷O detection sensitivity, the apparent SNR (i.e., the SNR in unit acquisition time) of the ¹⁷O MRS is still better than that of ¹H, ¹³C and ³¹P MRS, if the identical coil performance, acquisition scheme (e.g. at fully relaxed condition) and metabolite concentrations are used in comparison. Such unexpected superiority in sensitivity may be attributed to the relaxation times of the ¹⁷O nucleus, which are extremely short (in the order of few milliseconds), thus allowing efficient signal averaging during acquisition. To further improve the spatial resolution of the ¹⁷O-MR based CMRO₂ mapping, novel MR imaging techniques with efficient signal acquisition scheme and/or denoising capability are worthy of exploration [60].

In summary, after nearly three decades, in vivo ¹⁷O MRS and imaging is slowly emerging as a promising MR technique for noninvasively studying oxygen metabolism and perfusion in aerobic organs with the capability of imaging regional metabolic rates of oxygen and its change; thus providing a valuable tool for biomedical research of oxygen metabolism, mitochondrial function and tissue energetics in health and disease.

Abbreviations

α

¹⁷O enrichment fraction of inhaled ¹⁷O₂ gas

B₀

Magnetic field strength

BOLD

Blood oxygenation level dependent

C_a(t)

Time-dependent H₂¹⁷O concentration in excess of the natural abundance level in the arterial blood

C_b(t)

Time-dependent H₂¹⁷O concentration in excess of the natural abundance level in the brain tissue

C_v(t)

Time-dependent H₂¹⁷O concentration in excess of the natural abundance level in the venous blood

CBF

Cerebral blood flow

CBV

Cerebral blood volume

CMRO₂

Cerebral metabolic rate of oxygen

FDA

Food and Drug Administration (U.S.)

fMRI

Functional magnetic resonance imaging

H₂¹⁷O

Oxygen-17 labeled water

k

Decay rate of the endogenous H₂¹⁷O tracer

K_L

Rate constants reflecting the loss of H₂¹⁷O within the imaging voxel

K_G

Rate constants reflecting the gain of H₂¹⁷O within the imaging voxel

λ

Brain-blood partition coefficient

MR

Magnetic resonance

MRI

Magnetic resonance imaging

MRS

Magnetic resonance spectroscopy

MVO₂

Myocardial oxygen metabolic rate

OEF

Oxygen extraction fraction

¹⁷O₂

Oxygen-17 labeled oxygen gas

PET

Positron emission tomography

Q factor

Quality factor of the radio frequency coil

RF

Radio frequency

SAR

Specific absorption rate

SNR

Signal-to-noise ratio

T₁

Longitudinal relaxation time

T₂ (or T₂^*)

Transverse relaxation time (or apparent T₂)

tMCAo

Transit middle cerebral arterial occlusion

Yv

Cerebral venous oxygen saturation