Quantitative and simultaneous measurement of oxygen consumption rates in rat brain and skeletal muscle using 17O MRS imaging at 16.4T

Oxygen‐17 (17O) MRS imaging, successfully used in the brain, is extended by imaging the oxygen metabolic rate in the resting skeletal muscle and used to determine the total whole‐body oxygen metabolic rate in the rat.


| INTRODUCTION
Noninvasively measuring cellular oxygen metabolism using oxygen-17 ( 17 O) tracer and in vivo 17 O spectroscopic imaging ( 17 O MRSI) at ultrahigh field is a promising tool for studying cellular energy metabolism and physiology. 1 The 17 O imaging approach allows to quantify the cerebral metabolic rate of oxygen (CMRO 2 ) in human 2 and animal brain. [3][4][5] Although imaging studies have been performed to differentiate the cerebral metabolic rates between gray and white matter with 17 O, 2,6,7 few measurements were done outside the brain, [8][9][10] and they often focused on aerobic organs with a high metabolic rate. Obviously, lower metabolic rates, such as in resting muscle tissue, result in slower turnover rates from 17  in modeling with resulting longer inhalation durations is affected by blood perfusion and recirculation. This unmet challenge motivates our investigation of the feasibility of imaging the low oxygen metabolic rate in resting skeletal muscle using 17 O MRSI in simultaneous comparison to the brain oxygen metabolism rates in the same subject. Using significantly longer 17 O 2 inhalation times, the amount of generated H 2 17 O in biological tissues and the resulting 17 O MR signal are largely increased, even more so by multiple inhalations, 7,11 allowing to reliably observe metabolic and perfusion 12 parameters at increased sensitivity. However, in advantage against 15 O PET,no subtraction scans have to be performed in 17 O imaging for the exclusion of vascular or gaseous oxygen signal. [13][14][15][16] In this study, we simultaneously acquired dynamic time courses of H 2 17 O signals in brain and muscle tissue in rats during 17  before, during, and after inhalations to determine the metabolic rates of oxygen consumption 2 ; and (2) a washout model applied to the postinhalation brain data to estimate CBF. 12 The 3-phase metabolic model, which was previously used in human gray and white matter, was modified to obtain the low resting-state metabolic rate of oxygen consumption in the rat skeletal muscle (muscle RMRO 2 ) for exploring the feasibility of using the 17 O MRSI method in other tissues. The washout technique, allowing for estimation of CBF as previously validated in rodent brain, was used to investigate 2 groups of rats ventilated with different blends of gases (oxygen with N 2 or N 2 O). Finally, recirculation of H 2 17 O leading to a new equilibrium at the end of the postinhalation period was observed and employed to estimate the organism's global metabolism rate (ie, total body oxygen consumption (VO 2 )), which was then compared to the regional metabolic rates of oxygen consumption.

| Three-phase model adaptation
A previously published model for determining the human brain oxygen metabolism rate, 2  The 3 terms on the right side of Equation (1) can be separated into: (1) the regional metabolic activity producing application of 17 O MRSI to a wider range of organs, although further validation is advised.

K E Y W O R D S
cerebral blood flow (CBF), cerebral metabolic rate of oxygen (CMRO 2 ), mitochondrial water, muscle resting metabolic rate of oxygen consumption, 17  O signal for each imaging voxel to derive the oxygen metabolic rate (MRO 2 ) in brain or muscle tissue (as CMRO 2 or RMRO 2 ). We propose herein that, in principle, for any sufficiently perfused organ, oxygen consumption rates even below the systemic global aerobic rate (VO 2 ) can be measured. The quantification is simplified if the water content of the imaged tissue, which can be calibrated by the H 2 17 O natural abundance concentration and the 17 O signal measured in phase 1, is known. [3][4][5] The water content of muscle and brain can be approximated by assuming comparability to humans (ie, mice 17 : 74.4% weight [wt] in muscle vs. human 18 : 79.5% wt in striated muscle and 73.3% wt in brain). Furthermore, the tissue density for rodents (1.06 kg/L for skeletal muscle 19 ) was employed for unit conversion. Equation (1) can then be used to determine the oxygen metabolic rates of the rodent muscle and brain.

| Systemic oxygen expenditure VO 2
The total body oxygen consumption (VO 2 in the unit of μmol/g body weight/min) can be defined as the cumulative amount of metabolic H 2 17 O added to the organism by inhalation and metabolism of 17  ) at equilibrium (in this study, ~3 inhalation durations after 17 O 2 inhalation) when its pre-inhalation level is set to 0; and the conventional format of VO 2 in volume of oxygen gas is usually given in mL/kg body weight/min and requires a unit conversion f by division of 0.0446 μmol −1 mL. 20 3 | METHODS

| Simulation of circulation impact and metabolic rate on the H 2 17 O time courses
Inhalations with a 17 O 2 enrichment of 70% were simulated using the previously outlined 3-phase model for 2 settings: (1) simulation with a fixed high metabolic rate (ie, isometabolic CMRO 2 = 2 μmol/g/min) varying only the circulatory parameters (K G and K L ) in ranges reported in the literature 2,6,7,9,21 (phase 2 with 15.25 min inhalation duration); and (2) simulation of varying metabolic rate and corresponding changes in perfusion. In the second-stage simulation, different levels of local oxygen metabolic rates were set (MRO 2 = 2; 1; 0.5; 0 μmol/g/min) with fixed parameters K L = 0.2 and K G = 0.3, unless otherwise noted, during an inhalation using Equation (1) to qualitatively assess the time dependence of the tissue H 2 17 O signal. Specifically, this allowed investigation under idealized conditions of the transitions between phases of the model. 2 Furthermore, we adapted and evaluated the rodent specific systemic recirculation parameter as detailed in the Supporting Information Figure S1.

| Animal preparation and physiology monitoring
All procedures and experiments were approved by the local authorities (Regierungspräsidium [German for government council], Tübingen, Germany) and were in compliance with the guidelines of the European Community (EUVD 86/609/EEC) for the care and use of laboratory animals. A total of 8 male Wistar rats (Charles River Laboratories, Sulzfeld, Germany) were used in this study (Table 1) (mean body weight 312 ± 93 g). Artificial ventilation and maintenance of physiological stability is described in detail in the Supporting Information.
Ventilation mixtures with enriched 17 O 2 gas (oxygen gas fraction ~25%-35% with 70% enriched 17 O 2 Nukem GmbH, Germany) were prepared in nondiffusive gas bags (Hans Rudolph, Inc., Shawnee KS). Oxygen was mixed with N 2 in 1 group (Table 1, animals A-D; n = 4) and with N 2 O in a second group of animals ( Table 1, animals E-H; n = 4). At the end of each experiment, the animals were euthanized followed by postmortem imaging as previously reported. 22

| MRI instrumentation and data acquisition
MRI was performed on a BioSpec Avance III system (Bruker Biospin MRI GmbH, Ettlingen, Germany) using a 26 cm bore 16.4 T magnet and gradients with 12-cm inner diameter, In the majority of animals (Table 1, animals A-F, referred to as high-resolution protocol), the FOV was scanned by an acquisition matrix of 15 × 7 × 7, resulting in a voxel volume of 43.1 μL as defined by the width of the spatial response function. [24][25][26] Each 3D 17 O CSI volume was acquired within 30.2 s, with a maximum number of averages nt max = 74 at the k-space center (a total of 6144 FIDs or 735 k-space points per CSI volume). Fifty natural abundance H 2 17 O CSI were acquired within ~25 min at baseline, and a total of 109 volumes per full inhalation, started shortly before it, were collected within 54 min 57 s (see Table 1 for individual inhalation durations), including a ~38 min long postinhalation acquisition (ie, the H 2 17 O washout period).
In a subgroup of 2 animals (Table 1) (animals G and H, referred to as low-resolution protocol), the same FOV was scanned with an acquired matrix of 9 × 7 × 7, leading to a voxel size of 77.3 μL by spatial response function adjustment with nt max = 45 averages at the k-space center (a total of 2,048 FIDs or 441 k-space points per CSI volume) and 10.1 s acquisition per 3D CSI volume. Natural abundance H 2 17 O CSI volumes (n = 50) were acquired within ~8 mins 24 s, and the same acquisition duration of 54 min 57 s was used to acquire 327 volumes of inhalation data. Other acquisition parameters remained the same. Postmortem CSI-acquisitions were performed without k-space weighting (12 ms TR and 70° flip angle) and with a pulse length of 400 μs. A FOV of 27.5 × 12.5 × 25 mm 3 was sampled with a matrix of 41 × 19 × 25 voxels (nominal voxel size 0.44 μL). Approximately a total of 2.5 million FIDs, with 1000 points each and a spectral bandwidth of 100 kHz, were acquired in 8 h 18 min.

| Brain coregistration and tissue selection
The 3D 17 O-CSI data were coregistered with 1 H anatomic images and high-resolution (postmortem) H 2 17 O images with the same FOV as illustrated in Figure 1A-C. Equally sized regions of interest (ROI) were selected (−3 mm Bregma) 27 for brain and in lateral muscle compartments in the same coronal Pop. Mean ± SD: 312 ± 93 g 15 ± 0.6 min Each row represents 1 resting 17 O 2 inhalation measurement, which for rats A, C, D, and E was repeated multiple times within the same experimental session per animal. 17 O, oxygen-17; Inh, inhalation.
T A B L E 1 Summary of performed inhalation numbers and inhalation durations and weight for each animal slices. The topography of the temporalis muscle was verified anatomically, [28][29][30] and left and right lateral ROIs (42.4 μL, n = 40 voxels after zero-filling in animals A-F; 49.4 μL, n = 28 voxels after zero-filling in animals G-H) were chosen as a subset of the temporalis volume (0.422 mL), 31 carefully avoiding partial volume contamination from adjacent brain tissue. Separately, for each rat the data of 2 rats in the low-and high-resolution protocols were phased, and the localized semilogarithmic FIDs were fitted against time for in vivo T * 2 relaxation measurement as described in detail in Ref. 22. The metabolic model was fitted according to Equation (1) using a nonlinear least-squares algorithm (Curve Fitting Toolbox 3.5.6, MatLab; MathWorks, Natick, MA, version 9.3.0.713579 (R2017b)) to the H 2 17 O signal time courses of tissue signal (inhalation timeas independent variable; CMRO 2 for brain and RMRO 2 for muscle, K G , K L as dependent variables).  The primary decay constant, proportional to CBF/k 1 , can be converted by multiplication with 1.86 to absolute CBF units of mL/g/min (whereas k 3 and k 4 are scale factors). 12 Next, the 2 groups with different ventilation mixtures are compared (N 2 vs. N 2 O).

| Estimation of CBF
All results are reported in mean ± SD.

| RESULTS
Proton structural images showed a clear anatomical contrast between brain and muscle tissue ( Figure 1A). Coregistered geometry of 17 O contrast in both in vivo ( Figure 1B) and ex vivo 17 O high-resolution images ( Figure 1C) matched the anticipated intensity distribution of the 17 O surface coil, that is, stronger 17 O water signal at the surface and in the quadrature B 1 field overlap region in the brain.

O dynamics
The simulation results shown in Figure 3A Figure 3B for 4 different metabolic rates, exemplifying representative values for the brain and the muscle. The simulated metabolic rates at different levels showed a qualitatively distinct shape of the H 2 17 O signal dynamics at low metabolism (ie, sigmoidal). Despite significant differences in the early phase 2 (phase 2A), the slopes converge in a nonlinear way during the late phase 2 (phase 2B), as shown in Figure 3B. The simulation results indicate that the early dynamic change of the tissue H 2 17 O signal after inhalation of 17 O 2 gas is more sensitive to the local metabolic rate than that of late phase 2.
A novel observation from this simulation was that the same K G /K L ratio leads to the same equilibrium level of H 2 17 O signal at the end of phase 3 ( Figure 3A for brain and 3B for muscle at K G /K L ratio = 1.5). This suggests that even if the oxygen metabolic rates vary greatly in different tissues (eg, brain vs. muscle), the relative contributions of the H 2 17 O signal gain and signal loss due to recirculation and perfusion in different voxels remain the same. Thus, the voxels containing different tissue types will eventually reach the same H 2 17 O concentration level.

| Metabolic rate estimates for brain and resting skeletal muscle tissue
As shown in Figure 4, O signal was heterogeneously affected in SNR due to the B 1 sensitivity profile of the two 17 O surface coils being used) brain ROIs ( Figure 4B). Reproducible time courses were observed during 3 repeated inhalation measurements in the same animal and MR imaging session ( Figure 4C).
Fitting the metabolic rates of brain ROIs, an overall average of CMRO 2 = 1.97 ± 0.19 μmol/g/min (n = 26 ROIs from all 8 rats) was determined. For the 2 subgroups consisting of 4 rats each, a CMRO 2 of 2.07 ± 0.15 (n = 14 ROIs) and slightly lower 1.84 ± 0.14 μmol/g/min (n = 12 ROIs) were estimated with N 2 and N 2 O, respectively (Table 2); and no significant differences between the 2 hemispheres were detected. In O water content in low metabolizing tissue increases much slower than that in the higher ones (phase 2A); signals approach a similar slope during phase 2B; and at the end of the inhalation, the H 2 17 O in low metabolizing tissue continues to rise with gradually decreased slope. Both high and low metabolizing tissue approach the same equilibrium due to recirculation of body water at a new steady-state level determined by the global metabolic rate (VO 2 , according to Equation (2)). Unless otherwise stated in the legend, all time courses in (B) had K G = 0.3 and K L = 0.2. The dashed and dotted lines indicate the beginning and end of the phase 2, respectively. K G , K Gain parameter; K L , K Loss parameter | 2239 WIESNER Et al.
muscle ROIs, an average RMRO 2 of 0.32 ± 0.12 μmol/g/min (n = 23 ROIs) was determined, with some notable intrasubject left and right lateral differences. The estimated muscle oxygen metabolic rates were only one-sixth that of the brain. The perfusion-and diffusion-related parameter K G was higher than the parameter K L for both tissue types (for brain: averaged K G = 0.34 ± 0.05, K L = 0.22 ± 0.03, n = 26; and for muscle: K G = 0.63 ± 0.33, K L = 0.40 ± 0.17, n = 23). Group averages for brain tissue were K G = 0.34 ± 0.04 (n = 14) for N 2 and K G = 0.34 ± 0.07 (n = 12) for N 2 O. without O concentration in both tissue types was confirmed by matching 99% prediction bounds of the fit (dashed surrounding lines). The baseline H 2 17 O concentration was set to 0, which was above natural abundance level due to a prior 17 O 2 inhalation. Finally, a comparable fit quality was shown by the very similar spread of the confidence intervals despite their significantly different metabolic estimates.
(C) Normalized display of multiple inhalations showing the reproducibility of the technique in the same animal. Note the reproducible convergence at the end of each inhalation time course despite the differences in metabolic rate. With longer experimental duration, maintaining stable anesthetic conditions becomes more challenging as also reflected in higher signal fluctuations a statistically significant difference. In contrast, K L = 0.20 ± 0.02 (n = 14) for brain within the N 2 group increased by +22% to K L = 0.24 ± 0.02 (n = 12) in the N 2 O group, with statistical significance (2-sided unpaired t test at P < .005). The overall ratio of K G /K L determined within sessions was 1.51 ± 0.23 (n = 23 ROIs) for muscle and 1.58 ± 0.23 (n = 26 ROIs) for brain tissue, respectively; no statistically significant difference between the 2 tissue types was observed. The same K G /K L ratios between the brain and muscle converged to the same level of equilibrium H 2 17 O signal at the later phase 3 ( Figure 4B) despite > 6 times of difference in the metabolic rate between the 2 tissues. This finding is in agreement with the prediction from the simulations shown in Figure 3

| CBF and VO 2 in N 2 O versus N 2 ventilated animals
The estimated average oxygen metabolic rate of the entire body per gram tissue VO 2,average according to Equation (2) was 1.08 ± 0.20 μmol/g/min (n = 13) and 1.08 ± 0.16 μmol/g/min (n = 13), as inferred from right and left averaged muscle and brain ROI time courses, respectively (Table  3). Consistent with the conventional unit commonly used in the literature, VO 2 was converted to 24.2 ± 3.6 mL/kg body weight/min, derived from the steady-state H 2 17 O signals from both tissue types. CBF in the N 2 ventilated animal brain from average k 1 = 0.15 ± 0.01 (n = 14) resulted in CBF 0.28 ± 0.02 mL/g/min, and significantly elevated CBF (+21%, P < .005 with 2-sided unpaired t test) was observed in the N 2 O ventilated group with a mean of 0.34 ± 0.06 mL/g/min (n = 12) (Table 3). Figure 5 shows the lower T * 2 of H 2 17 O in muscle tissue and ~40% higher T * 2 in brain tissue, which are correlated against the independent metabolic rates in the 2 types of tissues. In the same rats, a more than fivefold difference in metabolic rate between muscle and brain is apparent.

| DISCUSSION
This study demonstrates 3 perspectives about the utility of the noninvasive and quantitative 17 (Figures 3, 4) (Supporting Information Figure S1B). The contribution of recirculating water increased with inhalation time and gradually dominated the H 2 17 O signal in the later phase of the inhalation, resulting in converging slopes between high-and low-activity tissues as observed in experimental data.
T A B L E 3 Summary of VO 2,average (μmol/g tissue/min) and cerebral blood flow (mL/g tissue/min) results based on washout in brain *This subgroup of 2 animals was acquired at a higher temporal resolution (10 s per 3D CSI volume) and with lower spatial resolution protocol. **P < .05 significant population difference between N 2 (rats A-D) and N 2 O (rats E-H) groups (unpaired t test).

F I G U R E 5
Scatter plot of in vivo tissue T * 2 versus metabolic rate showing high correlation in both tissue types of brain and muscle from clearly separated clusters of the independent properties of metabolic rate and relaxometric behavior inside ROIs. The distinction benefits from the fact of a stark difference in T * 2 , as also previously reported for both in vivo and postmortem tissues. Each cluster is based on the pooled tissue type of 2 representative rats (rat A and rat G), of high-resolution and low-resolution protocols, respectively. Note: this is only an observation to confirm the accurate selection and size (ie, partial volume contamination) of the ROIs and does not imply a causal relation between relaxation rate and metabolic rate in either direction 2242 | WIESNER Et al.

| Determining the oxygen metabolic rates in muscle and brain
Despite the limited spatial specificity, arteriovenous difference measurements can still be regarded as the gold standard for oxygen consumption measurements. However, due to their invasiveness, they are less convenient, and the variability of draining vascular territory effects on reproducibility motivates the use of noninvasive alternatives such as 17 O MRSI/MRI with 17 O 2 tracer inhalation, as in parallel has been attempted through 15 O PET. 37,38 By fitting the H 2 17 O signal dynamics of the rat muscle ROIs to the adapted 3-phase metabolic model, the resting-state metabolic rate of oxygen consumption in skeletal muscle (RMRO 2 ) was 0.32 ± 0.12 μmol/g/min. Comparing to the literature reports of oxygen metabolic rates in skeletal muscle from the earliest in vitro estimates 39 to more recent studies 40 in Wistar rats, the results of the present study show a good agreement with the literature values (Table 4). Perfused rat hindquarter muscle metabolic rate was reported similar (eg, 0.37 μmol O 2 /g/ min), 41 depending on modality. 42 Other differences could be inherent to the heterogeneity of muscle fibers, [43][44][45][46] which in the case of the temporalis muscle is low 30,47 compared to other muscles (eg, soleus or gastrocnemius) and in other species. 29,47,48 To the best of our knowledge, this study is the first to report measurements of oxygen metabolic rates using 17 O MR imaging for resting skeletal muscle, although working cardiac muscle with a high oxygen consumption rate has been shown before in isolated heart 8 as well as in vivo rat heart. 10 In muscle, alternative pathways (ie, fatty acids) are possible in contrast to the glucose-based metabolism of the brain. 30,31 However, both are based on oxygen as the substrate in the predominant mitochondrial electron transfer chain as origin of metabolic H 2 17 O. Therefore, this study is in agreement with previous measurements in the cardiac muscle, both perfused 8 and in vivo, 10 but extends to a much lower regime of metabolic rates in the immobilized, resting skeletal muscle, with very distinguishable characteristics. 2 The averaged CMRO 2 value (= 1.97 ± 0.19 μmol/g/min) as determined in this study is in agreement with the value (= 2.19 ± 0.14 μmol/g/min) from a literature report in the rat brain under relatively lower dose α-chloralose anesthesia obtained with a different modeling and experimental protocol. 3 These comparisons provide strong evidence to support the validity and reliability of the quantitative 17 O MRS imaging method, as described in this work for noninvasively imaging oxygen metabolic rates in the brain and resting muscle with a very low metabolic activity. Thus, we conclude that the same imaging approach should be applicable for most organs across a wide range of metabolic rates.

| Global systemic metabolic rates
It should be reasonable to assume that the metabolite pools are in equilibrium upon a stable physiological condition of the animal. 49 As observed in both simulation and experimental data, the post-inhalation H 2 17 O concentrations of different ROIs containing brain or muscle tissue eventually converged to the same steady-state level, which represented the new equilibrium H 2 17 O concentration after the 17 O 2 inhalation. Based on that information and Equation (2), we were able to derive the global systemic metabolic rate. Metabolic inter-or intrasubject fluctuations are likely caused by variations in the physiological animal condition (ie, ventilation parameters, anesthesia status, and body weight). Thus, in contrast to other studies, 7 our estimates of the average global oxygen metabolic rate (Table 3)  O in brain tissue during the post-inhalation period (ie, related to perfusion or CBF) has been established previously. 12 It reflects the dynamics of perfusion washout of the metabolically produced H 2 17 O in brain tissue and an inflow of global recirculating H 2 17 O. However, there is no observable "washout" in the lower metabolizing muscle tissue below the average body oxygen metabolic rate (VO 2,average ~1.1 μmol/g/ min), presumably due to a substantial inflow effect from recirculating H 2 17 O and low metabolic activity. Thus, in contrast to brain tissue, a significant extent of "wash-in" from systemic recirculation after the 17 O 2 inhalation was observed in muscle ( Figure 4, from t = 15 min onwards).
An increase in CBF through vasodilation has been observed and reported before with high percentage N 2 O administration. 58 Thus, the anesthetic properties and vasodilatory effects of N 2 O may reduce the global metabolism and possibly uncouple it partially from the narrowly regulated cerebral local oxygen metabolism. 59

| Validation of the 3-phase model in future research
Although the influence of recirculating metabolic water is substantial depending on the regional and global organism rates, the 3-phase model accounts accurately for the metabolic rate differences between tissue types. Our measurements used long inhalation times of over 15 min, thus, requiring a nonlinear metabolic model. 2 It can also be concluded that the longer duration of the inhalation phase does not linearly increase the CMRO 2 measurement sensitivity: it is limited by the accumulation of recirculating total body H 2 17 O.
An internal ROI validation confirmed whether the voxels that were selected truly reflected the chosen tissue type by assessment of T * 2 against metabolic rate in brain and muscle. Figure 5 shows a plot of the independent properties of tissue T * 2 and metabolic rate values for the ROIs taken from muscle and brain under the 2 different 17 O MRS imaging protocols (low-vs. high-resolution protocol). Two well-separated clusters associated with the 2 types of tissues due to stark difference in transverse relaxation between the tissues (a much longer T * 2 in brain than that of muscle) 22 confirm the ROI placements. Particularly, the muscle ROIs covered the temporalis muscle sufficiently accurately. It also demonstrates that the strong divergence in metabolic rate reflects an underlying tissue difference. However, this approximate separation is only possible due to the significantly shorter T * 2 values of H 2 17 O in muscle than in brain tissue. 22 It also should be noted that certain metabolic rate variability stems from tissue heterogeneity within ROIs. For example, in brain tissue estimates, despite low intrasession variance (eg, see rat A) a hemispheric difference was likely induced by ROI placement near the boundary between brain and muscle tissues, leading to partial volume effects. Another technical limitation is the relatively low SNR of 17 O signal detected in the muscle due to short T * 2 22 and lateral differences in B 1 , resulting in ~half SNR than that of brain tissue (see the 17 O spectra in Figure  2A,2B). Therefore, the fidelity in imaging muscle could be improved, for instance, using a coil array covering both brain and muscle with optimal detection sensitivity.
Finally, we would anticipate smaller variations of the 17 O MRSI approach when potentially activating the muscle by stimulation, as was done in a different paradigm during varying workload for instance, in cardiac muscle, 8,10 resulting in an elevated oxygen metabolic rate. In previous brain experiments, with an implantable 17 O RF coil, the measurement of an arterial input function and the measurement of blood flow through H 2 17 O bolus measurements was used for a detailed investigation, which also allowed the calculation of oxygen extraction fraction. 3 Thus, in future studies in other rat muscles (eg, in the leg, by implantation of an arterial 17 O RF coil on the femoral artery or separately on the tail artery), the metabolic rate could be validated after electrical stimulation over a wide range of metabolic rates and perfusion. Dynamically measuring the increased metabolic rate during 17 O 2 inhalations could give new insights to different muscle fiber types. Furthermore, we would expect a simultaneous measurement to be robust in consideration of systemic changes in animal physiology.

| CONCLUSION
In this study, we have extended the applicability of in vivo 17 O MR imaging to measure and image the resting skeletal muscle with a very low oxygen metabolic rate (~16% of the brain tissue). We have also confirmed the consistency of the CMRO 2 results measured during prolonged and repeated inhalations of 17 O 2 gas in this study with previous findings.
Because the brain has a very high metabolic rate of oxygen consumption, in contrast to the very low rate in the resting muscle, we anticipate that the same 17 O MR imaging approach and modeling will be useful for other organs such as liver and heart. Therefore, we expect a broad impact of using the 17 O MR imaging technology for metabolic rate measurements in normal and diseased organs beyond the brain.