Development of ¹⁷O NMR approach for fast imaging of cerebral metabolic rate of oxygen in rat brain at high field

Abstract

A comprehensive technique was developed for using three-dimensional ¹⁷O magnetic resonance spectroscopic imaging at 9.4T for rapidly imaging the cerebral metabolic rate of oxygen consumption (CMRO₂) in the rat brain during a two-min inhalation of ¹⁷O₂. The CMRO₂ value (2.19 ± 0.14 μmol/g/min, n = 7) was determined in the rat anesthetized with α-chloralose by independent and concurrent ¹⁷O NMR measurements of cerebral H₂¹⁷O content, arterial input function, and cerebral perfusion. CMRO₂ values obtained were consistent with the literature results for similar conditions. Our results reveal that, because of its superior sensitivity at ultra-high fields, the ¹⁷O magnetic resonance spectroscopic imaging approach is capable of detecting small dynamic changes of metabolic H₂¹⁷O during a short inhalation of ¹⁷O₂ gas, and ultimately, for imaging CMRO₂ in the small rat brain. This study provides a crucial step toward the goal of developing a robust and noninvasive ¹⁷O NMR approach for imaging CMRO₂ in animal and human brains that can be used for studying the central role of oxidative metabolism in brain function under normal and diseased conditions, as well as for understanding the mechanisms underlying functional MRI.

In the brain, the majority of energy consumption occurs by neuronal activity (1). This energy need is met predominantly through oxygen consumption mediated by the mitochondrial respiratory chain, coupled to oxidation of glucose as the main carbon source (2). The oxidative energy need of the human brain to sustain neuronal activity constitutes a large fraction of the energy requirements (≈20%) in the entire body. The central role played by oxygen consumption is also evident in pathologies associated with the brain; perturbations in brain oxidative metabolism have been closely linked to many brain diseases such as schizophrenia, Alzheimer's disease, Huntington's disease, Parkinson's disease, and mitochondrial dysfunction, as well as aging problems (3–7). One line of evidence linking these diseases and cerebral oxidative metabolism is the histopathological finding that the activity of cytochrome oxidase, the key mitochondrial enzyme that catalyzes the reduction of oxygen to form water, is impaired in patients with schizophrenia (3) and Alzheimer's disease (4, 5).

Basal cerebral oxygen consumption is not uniform across different regions in the brain. It is recognized that capillary density in the brain is inhomogeneous (ref. 8 and refs. therein), suggesting that mitochondrial density, and hence, oxygen consumption are also likely to be inhomogeneous because a good correlation exists between capillarity and mitochondrial density in tissues (9). Alterations in neuronal activity due to stimulation or performance of a task increases this spatial nonuniformity by inducing regional changes in oxygen utilization as well as in blood flow and glucose consumption (10–12). In addition, the impairment in cytochrome oxidase content in the diseased brain was also found to affect different regions in the brain selectively (3–5). These considerations imply that the ability to image the cerebral metabolic rate of oxygen consumption (CMRO₂) in vivo is essential for efforts aimed at investigating cerebral oxidative metabolism under normal and pathological conditions.

The earliest approach of CMRO₂ measurement is based on the Kety-Schmidt method for measuring cerebral blood flow (CBF) together with arteriovenous differences of oxygen content (13). This method, thought to be the most accurate one in the literature, provides only global CMRO₂ information without spatial differentiation within the brain. Positron emission tomography (PET) has been widely used for imaging CMRO₂ in humans (14–16). By introducing oxygen gas enriched with the isotope ¹⁵O into the human body and monitoring the spatial distribution and accumulation rate of metabolic water (H₂¹⁵O), PET can determine regional CMRO₂ and provide a CMRO₂ image. However, the ¹⁵O-PET is unable to distinguish the radioactive signals between the ¹⁵O₂ molecule bound to hemoglobin and the ¹⁵O atom incorporated into H₂¹⁵O. To overcome this drawback, a measurement of cerebral blood volume using inhalation of C¹⁵O has to be performed in addition to the CBF measurement based on i.v. injection of H₂¹⁵O. These PET requirements make CMRO₂ measurements difficult to execute experimentally and complicate the calculation of CMRO₂ values.

The ¹³C NMR methods based on monitoring the cerebral metabolism of intravenously infused [1-¹³C] glucose (e.g., (10, 11, 17, 18)) provide an alternative approach for CMRO₂ measurements. The labeling kinetics for several intracellular metabolites, most importantly glutamate, can be unequivocally measured by these methods; extraction of CMRO₂ from experimental data relies on extensive modeling (10, 11, 18, 19). In addition, the ¹³C NMR methods require a long measurement time (60–120 min) because of relatively slow turnover rates of cerebral metabolites, and it is difficult, although not impossible, to achieve three-dimensional (3D) CMRO₂ imaging because of a relatively long repetition time required for signal acquisitions.

The possibility of using the ¹⁷O NMR spectroscopy/imaging techniques for monitoring labeled H₂¹⁷O as a tracer of oxygen utilization or tissue perfusion have been introduced a decade ago (20–30). These early studies can be divided into two distinct groups. The first group involved the simple ¹⁷O NMR approach for detecting H₂¹⁷O directly (20, 21, 24–27). Nonlocalized ¹⁷O NMR was used for most studies to estimate CMRO₂ in the whole brain of experimental animals as well as in the human occipital lobe (31) after inhalation of oxygen gas enriched with the ¹⁷O isotope (¹⁷O₂). Few studies attempted to obtain a coarse CMRO₂ image (≥800 μl voxel size) in the cat brain with a long measurement time (25, 27). These previous efforts, all of which were conducted at relatively low fields compared with what is available currently, demonstrated the dramatic limitations imposed by the low inherent sensitivity of ¹⁷O NMR due to its low gyromagnetic ratio. These sensitivity limitations have hindered further developments of the ¹⁷O NMR approach and lead to the exploration of a second group of methods that attempt to circumvent the sensitivity limitations of ¹⁷O NMR by using indirect detection through ¹⁷O-coupled protons.

The indirect ¹⁷O detection approach relies on the effect of ¹H-¹⁷O scalar coupling on the transverse relaxation time (T₂) of the water protons, which is considerably shortened by the presence of the ¹⁷O nucleus in the same water molecule (22). It has also been shown that the water proton T₂ of H₂¹⁷O can be restored to that of H₂¹⁶O when ¹⁷O decoupling is applied; in this case, the difference between proton signals obtained with and without ¹⁷O decoupling is related to the ¹⁷O content of the water sample (28–30). However, the indirect ¹⁷O measurements present difficulties for quantitatively correlating the proton signal changes to the concentrations of H₂¹⁷O because the water T₂ is sensitive to many physiologic parameters such as pH and temperature. In addition, the fractional change in ¹H signal induced by variations in metabolic H₂¹⁷O concentration is small at basal conditions. These indirect ¹⁷O detection approaches have so far not been successful in imaging CMRO₂ quantitatively.

Recently, we have investigated the longitudinal and transverse relaxation times [T₁, T₂, and apparent T₂ (T₂*)] as well as the signal-to-noise ratio (SNR) of ¹⁷O spins in water by using the ¹⁷O magnetic resonance spectroscopic (MRS) imaging method at 4.7T and 9.4T field strengths (32). Our results showed that the relaxation times of ¹⁷O water are field-independent, and the ¹⁷O NMR sensitivity increased approximately fourfold at 9.4T compared with a field strength of 4.7T. With this sensitivity gain, we were able to obtain two-dimensional ¹⁷O MRS images of natural abundance H₂¹⁷O in the rat brain at 9.4T with a temporal resolution of 12 s (32). These results demonstrated the potential of ¹⁷O MRS imaging at ultra-high field for mapping the spatial distribution of metabolic H₂¹⁷O dynamically in the rat brain.

In this study, we investigated the feasibility for developing the ¹⁷O NMR approach for imaging CMRO₂ in a small animal model during a brief inhalation of ¹⁷O₂ gas. We have successfully applied the 3D ¹⁷O MRS imaging method to determine CBF after a bolus injection of H₂¹⁷O through an internal carotid artery and the accumulation rate of H₂¹⁷O generated by oxidative metabolism in the rat brain during a two-min ¹⁷O₂ inhalation with superior spatial (27 μl) and temporal (11 s per 3D image) resolution. Furthermore, we performed experiments where such images were obtained together with continuous experimental monitoring of the arterial input function related to recirculated H₂¹⁷O via an implanted carotid artery MR coil. These measurements were used to calculate CMRO₂ values rigorously in image voxels by using a complete modeling that accounts for all contributions to the cerebral H₂¹⁷O content accumulated during an inhalation of ¹⁷O₂ and demonstrate the feasibility for imaging CMRO₂. Based on these results, we present high-resolution 3D CMRO₂ images obtained with ¹⁷O NMR under the basal condition in the rat brain anesthetized with α-chloralose.

Theory

The dynamic change of cerebral H₂¹⁷O concentration accumulated during an inhalation of ¹⁷O₂ is determined by three processes: oxygen consumption generating metabolic H₂¹⁷O in the brain, perfusion resulting in H₂¹⁷O washout from the brain, and flow recirculation bringing extra H₂¹⁷O into the brain. The mass balance equation of labeled H₂¹⁷O concentrations during an inhalation of ¹⁷O₂ is given by

1

where C_a(t), C_b(t), and C_v(t) are the time-dependent H₂¹⁷O concentrations expressed as [H₂¹⁷O] in excess of the natural abundance [H₂¹⁷O] level in the arterial blood, brain tissue, and venous blood, respectively. The constant α is the enrichment fraction of ¹⁷O-labeled O₂ gas. Proper use of Eq. 1 requires consistencies of units among all terms used in the equation. The determinations of C_b, C_a, and C_v are based on the ¹⁷O NMR measurements and calibrated by using the natural abundance H₂¹⁷O concentration (20.35 μmol/(g of brain water) for brain tissue, and μmol/(g of blood water) for blood, calculated from natural abundance H₂¹⁷O enrichment of 0.037% and the molecular weight of H₂¹⁷O = 19.0. Therefore, we prefer to use convenient units of μmol/(g of brain water) for C_b(t) and μmol/(g of blood water) for C_a(t) and C_v(t). CMRO₂ is expressed in the conventional unit of μmol/min/(g brain tissue) and is converted from the unit of μmol/min/(g brain water) by using the conversion constant f₁ given by f₁ = (g of brain tissue)/(g of brain water) = 1/β_brain = 1.266, where β_brain is 0.79 (14, 27). The constant f₂ is also a unit conversion factor and is given by f₂ = ρ _blood⋅β_blood/β_brain = 1.05 × 0.81/0.79 = 1.077, where ρ_blood = (g blood)/(ml blood) = 1.05 (33), and β_blood = (g blood water)/(g blood) = 0.81 (14). If water in brain tissue is in equilibrium with water in venous blood (i.e., fast water exchange across capillaries), f₂C_v(t) = C_b(t)/λ where λ = 0.90 is the brain/blood partition coefficient with the unit of (ml blood)/(g of brain tissue) (34). Substituting this relation and introducing two new correction parameters (n and m) into Eq. 1 leads to

2

The solution of Eq. 2 is

3

where m is a correction factor accounting for the limited permeability of water because water is not freely diffusible across the brain-blood barrier; the value of m is 0.84 for the CBF range that we studied herein (35). The constant n is another correction factor that accounts for the additional restriction on the permeability of the H₂¹⁷O generated through oxidative metabolism in the mitochondria (“metabolic” H₂¹⁷O). This additional restriction is included in this modeling because we have observed that the washout rate of the metabolic H₂¹⁷O after the cessation of ¹⁷O₂ inhalation was significantly slower than the washout rate of the H₂¹⁷O that permeates brain tissue subsequent to a bolus injection of H₂¹⁷O through the internal carotid artery in the rat.‡ The cause of this difference could be permeability restrictions imposed by the inner and outer mitochondrial membranes that must be traversed by all water molecules generated by oxygen consumption. The ratio of the washout rate of the metabolic H₂¹⁷O after the cessation of ¹⁷O₂ inhalation vs. the cerebral H₂¹⁷O after a bolus injection of H₂¹⁷O through the internal carotid artery gives the value of n. Limited permeability for the metabolic water has been reported in the human brain (36).

The CMRO₂ value can be precisely calculated by using Eq. 3 if all parameters of C_b(t), CBF, and C_a(t) can be experimentally measured for the same animal brain. The C_b(t) and CBF values can be determined by using ¹⁷O NMR approaches. To measure C_a(t), we designed an implanted ¹⁷O rf coil that permits the continuous measurement of C_a(t) in a rat carotid artery. This measurement can be simultaneously performed with C_b(t) measurement during an inhalation of ¹⁷O₂ by using separate rf channels on the MR spectrometer equipped with multiple receivers. In principle, the C_a(t) function is determined by the metabolic H₂¹⁷O generated in the entire animal body and the equilibration of this H₂¹⁷O with the blood compartment. Considering that the oxygen consumption rate is constant (i.e., time independent) in all tissues during the short inhalation time, one can approximate C_a(t) as a linear function of time [i.e., C_a(t) = At, where A is a constant]. This approximation was supported by our experimental results (see Fig. 5 presented later) and the literature. Then, the solution of Eq. 3 for CMRO₂ is

4

Therefore, according to Eq. 4, the CMRO₂ value can be calculated by experimental data on CBF, A, n, C_b(t), and other known constants (f₁, f₂, m, α, and λ) for each data point measured at different inhalation times for each voxel. All CMRO₂ values reported (except for those in Fig. 5) are averages of the values calculated as a function of time, excluding the first two time points because of their relatively large measurement errors and the fact that they may represent transient values during the approach to a steady-state rate of H₂¹⁷O production by oxidative phosphorylation. Such transient values are expected during the build-up of ¹⁷O₂ in the brain at values that exceed the K_m of cytochrome oxidase for oxygen. Ability to obtain measurements as a function of time and the use of only the time independent values avoids this complication.

Figure 5.

(A) Natural abundance ¹⁷O spectrum of H₂¹⁷O in the rat carotid artery blood (7 μl) obtained before inhalation of ¹⁷O₂. (B) Time course of ¹⁷O MR signals of C_a(t) in one carotid artery (●) and C_b(t) from a representative voxel in the rat brain (□) during inhalation of ¹⁷O₂. (C) Plot of the calculated CMRO₂ values using the complete modeling as described by Eq. 4 as a function of time.

CBF values can be determined from the “washout” rate of the tracer H₂¹⁷O in the brain tissue following a rapid bolus injection of H₂¹⁷O through one carotid artery (32) according to Eq. 5:

5

where C_b(0) is the cerebral H₂¹⁷O concentration after the arrival of all labeled water after the bolus injection. Fitting the exponential decay of the H₂¹⁷O washout curve according to Eq. 5 gives the CBF value.

Materials and Methods

NMR Methods.

All NMR experiments were performed on a 9.4T/31-cm bore magnet (Magnex Scientific, Abingdon, U.K.) interfaced to a Unity INOVA console (Varian). A multinuclear head surface-coil probe consisting of a four-turn oval-shaped ¹⁷O coil (≈ 1 cm × 2 cm; 54.25 MHz) and a large butterfly-shaped ¹H coil (400 MHz) was used.

A TURBOFLASH (fast low-angle shot) image sequence was used to acquire scout images [TR/TE = 8 ms/4 ms, field of view (FOV) = 3 cm × 3 cm, and image matrix size = 128 × 128]. The spatial localization of ¹⁷O MRS was achieved by using the 3D Fourier Series Window MRS imaging technique (32, 37). In this method, the k-space sampling is weighted according to the Fourier coefficients of a predetermined voxel shape to achieve an optimal filter applied at the acquisition stage. A cylindrical voxel shape (circular shape on the coronal orientation) was used in this study (37). Short T₁ value (4.5 ms) of cerebral H₂¹⁷O in the rat allows rapid signal acquisition, and ultimately, more signal averages within the same sampling time (32). Therefore, a short TR of 12 ms (≈3T₁) was used for gaining SNR. The acquisition time for each 3D ¹⁷O-MRS image was 11 s (total scan number = 948; rf pulse width = 50 μs, spectral width = 30 kHz; FOV = 18 × 18 × 15 mm³, 9 × 9 × 5 phase encodes). Phase-encoding gradient was 0.35 ms long and had a half-sine waveform, which allowed for a short gradient echo time (TE = 0.4 ms). This short TE is critical for minimizing signal loss because of the relatively fast transverse relaxation of the ¹⁷O magnetization (T₂ = 3.0 ms at 9.4T; ref. 32). The peak strength of the phase-encoding gradient increment was 0.09–0.11 G/cm. The image voxel size was 27 μl. A 17 × 17 × 9 matrix of free induction decay (FIDs) were generated from the original 9 × 9 × 5 phase encode data for each 3D ¹⁷O image. The FIDs were zero-filled, and a 100-Hz line broadening (LB) was used before fast Fourier transformation for SNR enhancement. The ¹⁷O NMR signals from each voxel were quantified by measuring the H₂¹⁷O resonance peak intensities.

General Animal Preparation.

Male Sprague–Dawley rats (250–300 g body weight) were anesthetized with ≈2% (vol/vol) isoflurane in a mixture of O₂ and N₂O gases (2:3) during surgery. After oral intubations, femoral artery and vein were catheterized for physiological monitoring, blood sampling, and chemical administration. After surgery, anesthesia was switched to α-chloralose by using continuous infusion of 0.4 ml/hr of 15 mg/ml α-chloralose solution. Blood gases were sampled for monitoring physiological conditions [pH = 7.35 ± 0.03, pCO₂ = 42.1 ± 3.0 mmHg, pO₂ = 124 ± 16 mmHg (1 mmHg = 133 Pa)]. The arterial blood pressure (80–110 mmHg) and EtCO₂ level (≈3%) were monitored throughout the experiment. The rectal temperature of the rats was maintained at 37 ± 1°C by using a heated water blanket. All animal surgical procedures and experimental protocols were conducted under the guidelines of the National Institutes of Health and the Institutional Animal Care and Use Committee of the University of Minnesota.

CBF Measurements.

For CBF measurements, one external carotid artery was catheterized with a PE-50 tubing for gaining access to the internal carotid artery without interrupting blood circulation into the brain (32). A 0.05–0.1 ml 50%-enriched H₂¹⁷O was rapidly injected into the brain through the internal carotid artery for CBF measurements using ¹⁷O MRS imaging (32).

C_a(t) Measurements.

An implanted vascular rf coil was designed and constructed for continuously detecting the ¹⁷O NMR signal changes of H₂¹⁷O in the rat carotid artery at 9.4T. This coil was based on a modified solenoid coil design combined with an rf shielding.§ The rf shielding ensured that the NMR signal detected by the implanted coil was only attributed by the artery blood (≈7 μl) without contaminations from surrounding tissues. Therefore, additional spatial localization was not necessary for determining C_a(t). In addition, the rf shielding minimized the electromagnetic coupling between the implanted ¹⁷O coil and the head ¹⁷O surface coil tuned at the same operating frequency (two-coil arrangement), which allowed simultaneous measurements of both C_a(t) and C_b(t) by using two receiver channels with the same temporal resolution (11 s).

¹⁷O₂ Inhalation Experiments.

For ¹⁷O₂ inhalation studies, the ¹⁷O-labeled O₂ gas with 58.2–72.1% enrichment (Isotec) was mixed with N₂O gas (≈2:3) and then stored in a cylindrical gas reservoir. The 3D ¹⁷O MRS imaging acquisitions were started with the normal gas mixture first for 2 min; these natural abundance 3D ¹⁷O MRS images were used as references for quantifications of metabolic H₂¹⁷O concentrations in the rat brain during and after the ¹⁷O₂ inhalation. Subsequently, the respiration gas was quickly switched to the ¹⁷O₂-labeled gas mixture. After 2 min of ¹⁷O₂ inhalation, the gas line was switched back to the normal gas mixture, and the 3D ¹⁷O MRS imaging acquisitions were continued throughout the procedure; a total of 100 images were collected in about 19 min.

Determination of Constant n.

The ratio between the washout rate of the metabolic H₂¹⁷O after the cessation of ¹⁷O₂ inhalation and the washout rate of the H₂¹⁷O trace from the H₂¹⁷O bolus injected through the rat internal carotid artery gave the value of constant n for each voxel.

All results are presented as mean ± SD.

Results

The multinuclear surface-coil probe was used to collect a series of global ¹⁷O and ¹H spectra from the natural abundance water in the rat brain to estimate the relative signal fluctuations between ¹⁷O and ¹H NMR. Two stacked plots of 25 ¹⁷O and ¹H spectra acquired consecutively by using the ¹⁷O coil and the much larger ¹H coil, respectively, are shown in Fig. 1. Although the number of acquisitions (NT) for each spectrum was different for the ¹⁷O and ¹H nuclei, the repetition time (TR) was ≈3T₁ in both cases; the total acquisition time was 6.2 s for all spectra. Under this condition, we found that the signal fluctuations in the consecutively acquired spectra for the two nuclei were comparable (SD = 0.20% for ¹H spectra; SD = 0.35% for ¹⁷O spectra). This result demonstrates that with respect to the capability for detecting fractional changes in NMR signal intensity, the ¹⁷O method is comparable to the ¹H method.

Figure 1.

(A) Stacked plot of 25 ¹H NMR spectra of natural abundance water in the rat brain acquired and processed with the following parameters: NT = 1, TR = 6.2s, LB = 25Hz (≈ half linewidth). The SD of the ¹H water peak height in these spectra is 0.20%. (B) Stacked plot of 25 ¹⁷O NMR spectra of natural abundance water in the rat brain acquired and processed with the following parameters: NT = 512, TR = 12ms for each scan, LB = 100Hz (≈ half linewidth); the SD of the ¹⁷O water peak height in these spectra is 0.35%. The total acquisition times for each ¹H and ¹⁷O spectra were the same (6.2 s).

The 3D ¹⁷O MRS imaging technique was applied to measure the natural abundance H₂¹⁷O distribution in the rat brain with high temporal resolution of 11 s and spatial resolution (27 μl voxel size). Fig. 2A shows representative data obtained from three adjacent coronal images in a rat brain showing anatomical images acquired by the ¹H coil and the ¹⁷O MRS images of natural abundance H₂¹⁷O acquired by the ¹⁷O coil. Fig. 2B displays one ¹⁷O chemical shift image from the middle image slice as shown in Fig. 2A. The spatial distributions of ¹⁷O NMR signal intensity are not uniform because of the inhomogeneous B₁ of the ¹⁷O surface coil. Nevertheless, excellent SNR was achieved with a relatively short acquisition time at 9.4T, especially for the central voxels where SNR was optimized by the B₁ profile of the ¹⁷O coil (SNR ≈ 40:1).

Figure 2.

(A) 3D ¹⁷O brain images of natural abundance H₂¹⁷O from three adjacent slices (Left, color images) and corresponding anatomical images (Right, gray images) in the coronal orientation from a representative rat. (B) Chemical shift image of natural abundance H₂¹⁷O from Middle as shown in A.

The large sensitivity gained at ultra-high field revealed the potential of ¹⁷O MRS imaging for detecting small dynamic changes of cerebral H₂¹⁷O in small animal brains during a short inhalation of ¹⁷O₂. Fig. 3 demonstrates stacked plots of ¹⁷O spectra of cerebral H₂¹⁷O from one representative voxel before (natural abundance), during, and after a 2-min inhalation of ¹⁷O₂ (72.1% ¹⁷O enrichment). It is clearly evident that the ¹⁷O signal intensity of cerebral H₂¹⁷O in the entire measurement is characterized by three distinct phases: (i) constant before the ¹⁷O₂ inhalation; (ii) approximately linearly increasing during the ¹⁷O₂ inhalation; and (iii) decreasing after the cessation of ¹⁷O₂ inhalation, approximately exponentially, and reaching a new steady state within a short time (<10 min).

Figure 3.

Stacked plots of the cerebral H₂¹⁷O spectra from one representative voxel (27 μl voxel size) as indicated in the anatomical image (Left Inset) acquired before (natural abundance), during (as indicated by the dark bar under the stacked plots) and after a 2-min ¹⁷O₂ inhalation.

When the cerebral H₂¹⁷O concentrations reached a new steady state after the ¹⁷O₂ inhalation experiment, the CBF measurement was performed by a rapid bolus injection of H₂¹⁷O into one internal carotid artery. The dynamic distribution of cerebral H₂¹⁷O signal during the washout period was monitored by using 3D ¹⁷O MRS imaging. The washout curves excluding the first data point were fitted to an exponential decay function according to Eq. 5, and these fits yielded CBF values. Fig. 4 demonstrates one example of CBF measurement obtained from the same rat and image voxel used in Fig. 3 showing the stacked plot of ¹⁷O spectra and the exponential fit. Because the H₂¹⁷O bolus was injected into one of the internal carotid arteries, and the H₂¹⁷O tracer was mainly distributed into the ipsilateral hemisphere, CBF and, subsequently, CMRO₂ quantifications were performed only for the ¹⁷O image voxels in the ipsilateral hemisphere of the rat brain.

Figure 4.

(A) Stacked plots of the cerebral H₂¹⁷O spectra from one voxel (used in Fig. 3) after a fast bolus injection of H₂¹⁷O into the rat brain through an internal carotid artery. (B) Exponential decay fitting of the H₂¹⁷O washout curve for calculating CBF (0.56 ml/g/min for this voxel).

Fig. 5A illustrates the natural abundance H₂¹⁷O spectrum from the blood in the rat carotid artery. Fig. 5B plots the results of simultaneous measurements of C_a(t) change in one carotid artery and C_b(t) change in a voxel in the ipsilateral hemisphere brain during an inhalation of ¹⁷O₂. The C_a(t) data were fitted by a linear function for calculating the value of constant A. Finally, the values of C_b(t), CBF, and n from each voxel and the value of A measured from each ¹⁷O inhalation measurement were used to calculate CMRO₂ according to Eq. 4. Fig. 5C demonstrates one example of CMRO₂ calculations as a function of time from a single voxel. It is evident in Fig. 5C that the CMRO₂ value is independent of the inhalation time if the initial two points are excluded. The averaged CMRO₂ values (excluding two initial points) and CBF values from seven measurements are 2.19 ± 0.14 μmol/g/min and 0.53 ± 0.07 ml/g/min, respectively, in the rat brains anesthetized with α-chloralose.

The mean CBF and n values obtained from the entire hemisphere ipsilateral to the internal carotid artery where a H₂¹⁷O bolus was injected were subsequently used to calculate CMRO₂ values for both ipsilateral and contralateral hemispheres approximately and to obtain 3D CMRO₂ images. This approximation was made based on the fact that the CBF distribution in the hemisphere where it was measured was relatively uniform and that two hemispheres should be equivalent with respect to CBF. Fig. 6 illustrates three adjacent CMRO₂ images in the coronal orientation from a representative rat brain. The inhomogeneous signal intensity caused by the surface coil sensitivity profile (Fig. 2) is not reflected in the CMRO₂ maps, as expected.

Figure 6.

3D coronal CMRO₂ images of rat brain measured during a 2-min ¹⁷O₂ inhalation.

Discussion

The ¹⁷O T₁ value in H₂¹⁷O is less than 5 ms in the rat brain at 9.4T (32). This short T₁ permits much faster data acquisitions, more signal averages, and, consequently, a high SNR per unit of time. This advantage is critical for the success of CMRO₂ measurement using direct ¹⁷O detection. The comparable fractional signal fluctuations between consecutively acquired ¹H and ¹⁷O NMR spectra (Fig. 1) indicated that direct ¹⁷O detection of changes in metabolic H₂¹⁷O signal is likely to be more sensitive than the indirect ¹⁷O detection through ¹H NMR. This notion is based on the fact that the fractional change of water proton signal caused by a 2-min ¹⁷O₂ inhalation at basal condition is estimated to be 0.15–0.30% (29), which is difficult to detect reliably because this change is comparable to the signal fluctuations in consecutively acquired ¹H spectra or images. These signal fluctuations are not dominated by the inherent SNR of a single spectrum or image but are mainly determined by the physiological processes of breathing, heart pulsation, and vasomotion, as well as NMR scanner instability. In contrast, ¹⁷O NMR has a much larger dynamic range for detecting the changes of metabolic H₂¹⁷O (20–40%) during the same ¹⁷O₂ inhalation. This difference is the major merit that makes ¹⁷O NMR more suitable for CMRO₂ imaging.

The ¹⁷O approach has several major advantages for CMRO₂ imaging in comparison with other existing methods. First, ¹⁷O NMR can specifically detect metabolic H₂¹⁷O without confounding signals from the ¹⁷O₂ bound to hemoglobin or ¹⁷O₂ dissolved in tissue space. When bound to hemoglobin, the ¹⁷O₂ resonance is broadened beyond detection because of the slow rotational motion for the large molecular weight oxyhemoglobin complex. ¹⁷O₂ as an unbound molecule in gas phase or dissolved in water is strongly paramagnetic because of its unpaired electrons and, hence, undetectable because of the dipolar coupling between the electrons and the nucleus. As a result, ¹⁷O NMR avoids the complication in the calculation of CMRO₂ encountered in the PET methodology. In addition, unlike PET, the ¹⁷O isotope is stable and nonradioactive, and the experimental procedure of ¹⁷O₂ inhalation does not introduce any physiological perturbation and complexities in the animal. These merits can significantly simplify the modeling of CMRO₂ calculation as well as provide an approach for imaging CMRO₂ in living organs with minimal biological risk. Second, the natural abundance signal of H₂¹⁷O, which can be accurately imaged, provides an internal reference for rigorously calibrating and calculating the absolute H₂¹⁷O concentration changes for both C_b(t) and C_a(t) during ¹⁷O₂ inhalation; such a quantification is crucial for determining absolute CMRO₂ values and is independent of the inhomogeneities of ¹⁷O coil sensitivity. Third, there is only a single and well defined ¹⁷O resonance peak of H₂¹⁷O, thus making spectral analysis simple. Fourth, the ¹⁷O resonance peak is relatively broad and has a low resonance frequency (only 14% of the ¹H resonance frequency); hence, it is significantly less sensitive to static magnetic field inhomogeneities. Fifth, after the cessation of ¹⁷O₂ inhalation, the cerebral H₂¹⁷O concentration reaches a new steady state within a short time (6–10 min); this fast recovery should allow repeated CMRO₂ measurements in the same subject and same experimental session. This capability is essential for studying the oxidative metabolism changes related to perturbations in physiology and function, where at least two measurements are required under control and perturbation conditions. Finally, at ultra-high fields the ¹⁷O NMR sensitivity is adequate to obtain relatively high-resolution images, even relative to the small size of the rat brain, with an 11-s acquisition time; this sensitivity makes the ¹⁷O approach possible for imaging CMRO₂ within a short measurement duration (2 min).

Because of the technical challenges faced for most existing CMRO₂ imaging methods, there is virtually no literature reporting direct measurements of CMRO₂ in the rat brain. For instance, the image voxel size of most conventional PET scanners is on the order of several milliliters (e.g., ref. 38); this size is already larger than the size of the entire rat brain. Nevertheless, we compared our result of CMRO₂ measurement by using the ¹⁷O NMR approach with the literature results obtained by indirect methods under similar condition. Based on an autoradiographic approach (39), cerebral metabolic rate of glucose (CMR_glc) was recently determined to be ≈0.37 μmol/g/min in the somatosensory and motor cortices of rats anesthetized with α-chloralose. By using the CMRO₂/CMR_glc ratio of 5.5 (2), a CMRO₂ value of 2.04 μmol/g/min is calculated from this CMR_glc measurement, which is in excellent agreement with our ¹⁷O-based CMRO₂ result (2.19 ± 0.14 μmol/g/min). The CBF value of 0.58–0.68 ml/g/min reported in the same autoradiographic study (39) is also close to the mean CBF value of 0.53 ± 0.07 ml/g/min measured in our study. The heteronuclear editing MRS and ¹H MRI methods covering a large range of cortical activity at different anesthesia conditions had been used to measure CMRO₂ and CBF values in the sensory motor cortex of mature rats, and a linear correlation between CMRO₂ and CBF was observed (CMRO₂ = 3.76⋅CBF + 0.18, R² = 0.99; ref. 40). The estimated CMRO₂ from this relation is 2.16 μmol/g/min if the CBF value observed in our study (0.53 ml/g/min) is applied. This CMRO₂ value is again in excellent agreement with our result. These comparisons provide mutual support between our direct CMRO₂ measurements and other indirect measurements reported in the literature and support the validity of the 3D CMRO₂ imaging using ¹⁷O NMR.

The CMRO₂ measurements presented in this paper involved two invasive procedures: one for determining C_a(t) by means of the implanted ¹⁷O rf coil, and the second for measuring CBF by an intra-arterial catheter and bolus injection of H₂¹⁷O. The CBF measurement can be performed noninvasively by using the arterial spin tagging MR approaches (41). The implanted rf coil, on the other hand, is not practical for routine measurements of CMRO₂ by using small experimental animal models (e.g., rat or mouse), nor is it suitable for human applications. Therefore, it would be important to explore further the feasibility of the ¹⁷O NMR approach for reliably imaging CMRO₂ without resorting to invasive C_a(t) measurements. A successful outcome from such a study will establish a completely noninvasive ¹⁷O approach for CMRO₂ imaging.

In comparison with most existing approaches, the high-field ¹⁷O NMR method requires a significantly short measurement time for imaging CMRO₂ and can be performed within a 2-min ¹⁷O₂ inhalation period. However, this method assumes that CMRO₂ is constant during the measurement. SNR considerations suggest that the CMRO₂ measurement time can further shorten at 9.4T or higher magnetic fields. Although rapid dynamic changes of CMRO₂ that occur faster than this time-scale will be beyond the reach of this method, this time-scale is faster than any other technique currently available for assessing CMRO₂, especially at the spatial resolution available from the ¹⁷O NMR approach.

Conclusion

We have successfully developed a comprehensive technique by using ¹⁷O MRS imaging at ultra-high field for fast imaging of CMRO₂ in the rat brain and validated it with complementary measurements. The overall results from our study provide an essential step toward the goal for developing a robust, fast, reliable, and noninvasive ¹⁷O NMR approach for imaging the rate of oxidative metabolism for both animal and human brains and, potentially, for other living organs. Such an approach would be useful for studying the central role of oxidative metabolism in brain function under normal and diseased conditions, as well as for understanding the mechanisms underlying functional MRI (42).

Abbreviations

CMRO₂

cerebral metabolic rate of oxygen

CBF

cerebral blood flow

PET

positron emission tomography

3D

three-dimensional

MRS

magnetic resonance spectroscopic

SNR

signal-to-noise ratio