in vivo chlorine-35, sodium-23 and proton magnetic resonance imaging of the rat brain

Research Article

592

Received: 18 August 2009, Revised: 3 December 2009, Accepted: 4 December 2009, Published online in Wiley InterScience: 15 March 2010

(www.interscience.wiley.com) DOI:10.1002/nbm.1500

In vivo chlorine-35, sodium-23 and protonmagnetic resonance imaging of the rat brainStefan Kirscha*, Mark Augathb, David Seiffgec, Lothar Schillingc

and Lothar R. Schada

In this study we demonstrate the feasibility of com

NMR Biom

bined chlorine-35, sodium-23 and proton magnetic resonanceimaging (MRI) at 9.4 Tesla, and present the first in vivo chlorine-35 images obtained by means of MRI. With theexperimental setup used in this study all measurements could be done in one session without changing the setup ormoving the subject. The multinuclear measurement requires a total measurement time of 2 h and providesmorphological (protons) and physiological (sodium-23, chlorine-35) information in one scanning session. Chlorine-35,sodium-23 and high resolution proton imageswere acquired from a phantom, a healthy rat and from a rat displaying afocal cerebral infarction. Compared to the healthy tissue a signal enhancement of a factor of 2.2W 0.2 in thechlorine-35 and a factor of 2.9W 0.6 in the sodium-23 images is observed in the areas of infarction. Exemplaryunlocalized measurement of the in vivo longitudinal and transversal relaxation time of chlorine-35 in a healthy ratshowed multi-exponential behaviour. A biexponential fit revealed a fast and a slow relaxing component withT1,a¼ (1.7W 0.4) ms, T1,b¼ (25.1W 1.4) ms, amplitudes of A¼ 0.26W 0.02, (1–A)¼ 0.74W 0.02 and T2,a¼ (1.3W 0.1)ms, T2,b¼ (11.8W 1.1) ms, A¼ 0.64W 0.02, (1–A)¼ 0.36W 0.02. Combined proton, sodium-23 and chlorine-35 MRI mayprovide a new approach for non-invasive studies of ionic regulatory processes under physiological and pathologicalconditions in vivo. Copyright � 2010 John Wiley & Sons, Ltd.

Keywords: magnetic resonance imaging; in vivo; chlorine; sodium; 23Na mri; 35Cl mri; chloride

* Correspondence to: Dr S. Kirsch, Department of Computer Assisted ClinicalMedicine, Medical Faculty Mannheim, University of Heidelberg, Theodor-Kutzer-Ufer 1–3, D-68167 Mannheim, Germany.E-mail: [email protected]

a S. Kirsch, L. R. Schad

Department of Computer Assisted Clinical Medicine, Medical Faculty

Mannheim, University of Heidelberg, Mannheim, Germany

b M. Augath

Department of Physiology of Cognitive Processes, Max Planck Institute for

Biological Cybernetics, Tubingen, Germany

c D. Seiffge, L. Schilling

Division of Neurosurgical Research, Medical Faculty Mannheim, University of

Heidelberg, Mannheim, Germany

Abbreviations used: EEC, European Economic Community; EUVD, Euro-

pean Union Vendor Declaration; FID, free induction decay; FOV, field of view;

LDF, laser Doppler flow; MCA, middle cerebral artery; MR, magnetic resonance;

MRS, magnetic resonance spectroscopy; MSME, multi slice multi echo;

Q-factor, quality factor; rf, radio frequency; ROI, region of interest; SNR, signal

to noise ratio; TA, total measurement time; USR, ultra shield refrigerated; UTE,

ultra short echo time.

INTRODUCTION

In conventional clinical nuclear magnetic resonance imaging(MRI) the measured signals and the resulting images originatefrom the protons (1H) of the tissue-water molecules. Since thepresentation of the first in vivo sodium-23 (23Na) magneticresonance (MR) images by Hilal et al. (1) significant progress hasbeen made to establish in vivo imaging of sodium ions as animportant modality in human MRI research (2–12). In a recentstudy we demonstrated the feasibility of combined in vivo23Na and potassium-39 (39K) MRI (13). The MRI of 23Na and39K may provide a potent and worthwhile tool to study in anon-invasive manner the activity of the cellular Na/K pump underphysiological and pathological conditions in vivo.In addition to the cations Naþ and Kþ, chloride (Cl�) is the most

abundant anion in themammal organismplaying an important rolein many cellular processes. For instance, plasma membrane Cl�

currents are important for the regulation of excitability in nerve andmuscle. Moreover, Cl� ions play a crucial role in controlling the ioniccomposition of the cytoplasm and the volume of cells (14). Severalchlorine-35 (35Cl) magnetic resonance spectroscopy (MRS) studiesfocusing on different biochemical aspects of Cl� interactions withproteins (15–19), on membrane transport systems in living cells(20,21), and on the ex vivo perfused rat kidney (22) and rat heart (23)have been reported.In order to investigate the feasibility of combined in vivo

1H, 23Na and 35Cl MRI we developed a rf coil setup to measure1H, 23Na and 35Cl signals in one scanning session without movingthe subject or changing the setup. The performance of the rf coilsetup and the MRI pulse sequence is demonstrated on aphantom, a healthy rat and on a rat displaying a focal cerebral

ed. 2010; 23: 592–600 Copyright � 2010

infarction. To the best of our knowledge here we present the firstin vivo 35Cl images acquired with MRI and exemplarymeasurement of the T1 and T2 relaxation time in vivo.

THEORY

Chlorine occurs as the two isotopes 35Cl and 37Cl, with a naturalabundance of 75.78% and 24.22%, respectively (24). The nuclei ofthe 23Na, 35Cl and 37Cl isotopes have a spin of 3/2 and therefore

John Wiley & Sons, Ltd.

IN VIVO CHLORINE-35, SODIUM-23 AND PROTON MRI OF THE RAT BRAIN

they exhibit an electrical quadrupole moment. MRI ofquadrupolar nuclei suffers from low sensitivity due to theirphysical properties, e.g. the low gyromagnetic ratio and fastrelaxation.A special characteristic of spin-3/2 nuclei is their multi-

exponential relaxation. Theoretical treatment of relaxation ofspin-3/2 nuclei reveals that the longitudinal and the transverserelaxations produced by a quadrupolar interaction are the sumsof two or more decaying exponentials (27). Assuming a uniformlymagnetized sample, the signal decay in the rotating frame ofreference following an ax rf pulse can be written as (27)

Mz tð Þ ¼ M0 1þ cosa� 1ð Þ 0:8 exp �a1tð Þ þ 0:2 exp �a2tð Þð Þ½ �;(1)

Mþ tð Þ ¼ Mx þ iMy

¼ iM0 sina 0:6 exp �b1tð Þ þ 0:4 exp �b2tð Þ½ �: (2)

HereM0 denotes the equilibriummagnetization just before theax rf pulse, a¼ flip angle and x¼phase of the rf pulse,Mz¼ longitudinal and Mþ¼ transversal component of themagnetization. The relaxation rates a1, a2 and b1, b2 dependon the molecular mobility via the spectral density functions andthe correlation time tc [see Ref. (27) Eq. 7 and 9aþb]. Smallmolecules in low-viscosity solutions typically have correlationtimes of a few tens of pico seconds or less. In this so-calledextreme narrowing regime Eq. (1) and Eq. (2) reduce tomonoexponential decays (27).Typical reported spin-lattice relaxation times T1 and transversal

relaxation times T2 for 23Na at 8.4 Tesla are T1¼ 23ms(monoexponential signal decay) and T2¼ 2ms and 17ms(biexponetial fit on measured signal decay) (25). In the rat brainat 4.7 Tesla T1 and T2 values of T1¼ (40� 5) ms and T2¼ (34� 2)ms are reported (31). For 35Cl monoexponential signal decays andtransversal relaxation times of T2¼ 12.5–100ms for Cl� insolutions with arginine, histidine, lysine and alanine are reported(17). Furthermore, monoexponential signal decays and T2 times inthe range of 22–30ms are obtained for Cl� ions bound toproteins (cytochrome c) (19). In addition to the fast relaxation thein vivo concentration of Cl� and Naþ is much lower compared towater protons. The interplay of the different physical propertiesresults in in vivo 23Na MR signals which are 20 000 times lowerthan 1H signals (26). An estimate of the signal intensity expectedfrom 35Cl can be derived by comparing the receptivity in naturalabundance Dp of 23Na¼ 9.27 � 10�2 and Dp of 35Cl¼ 3.58 � 10�3

(24). Assuming a linear frequency dependence of the noise, i.e.increasing noise with increasing frequency, usingg23Na¼ 11.262MHz/T and g35Cl¼ 4.1717MHz/T the in vivo signalintensity of 35Cl is expected to be approximately 9.6 times lowerthan the 23Na signal intensity. Note, this estimation assumes samesensitivity of the 23Na and 35Cl receiving hardware (rf coils,transmit lines etc.).

5

MATERIAL AND METHODS

Experimental setup

The imaging experiments were performed on a 9.4 T BrukerBiospec 94/20 USR small animal system equipped with 740mT/mx, y, z-gradients and a laser controlled positioning system for theanimal cradles. The animal cradles provide stereotactic fixation

NMR Biomed. 2010; 23: 592–600 Copyright � 2010 John Wiley

with bite bar and ear pins and several mounting points for surfacecoils.In order to test the coil setup and optimize the imaging

sequence a phantom consisting of two plastic flasks (60ml and2ml) was built. The 2ml flask filled with a solution containing154mmol/l NaCl was attached inside the 60ml flask filled with a154mmol/l KCl soultion.In vivo MRI was performed on Sprague-Dawley rats

(290–350 g). All experiments were approved by the localauthorities and performed in full compliance with the guidelinesof the European community (EUVD 86/609/EEC) for the care anduse of the laboratory animals. Twenty four hours before the NMRmeasurement a rat underwent right-sided middle cerebral artery(MCA) occlusion to induce a focal brain ischemia. Surgery wasperformed under isoflurane anesthesia (2–2.5%) in a 80/20%air/O2 mixture as described previously (28). In brief, the animal’shead was fixed in a stereotactic head holder and a small burr holedrilled 0.5mm posterior and 4mm lateral to the bregma leavingthe inner bone layer intact. A small laser Doppler flow (LDF) probecoupled to a LDF monitor (Periflux system 5000, Perimed, Jarfalla,Sweden) was placed on the craniotomy to assess cortical bloodflow changes. After attaching the LDF fibre to the skull the animalwas turned into a supine position and the neck opened by anincision in the midline. The right common carotid artery wastemporarily ligated and a nylon suture (4–0) carrying a silicone tip(diameter, 450mm) inserted into the external carotid artery. Thefilament was advanced into the intracranial circulation until asharp drop of the LDF signal indicated occlusion of theMCA. Afterfixing the filament and removing the LDF probe all wounds wereclosed. The animal received a subcutaneous injection of 0.3mg/100 g body weight buprenorphin (Temgesic1 Essex Pharma,Munich, Gemany) to reduce postoperative pain before anesthesiawas ceased, 100min after MCA occlusion the animal was againanesthetized as described above and the LDF probe againattached. Two hours after MCA occlusion the filament waswithdrawn and the opening of the MCA verfied by a steepincrease of the LDF signal. All wounds were re-closed, a seconddose of buprenorphin (0, 15mg/100 g body weight) injected andthe animal allowed to recover.Twenty-four hours after MCA occlusion the animal was

anesthetized with 1.2–1.5% isoflurane and positioned in theNMR. The respiratory frequency and the body temperature weremonitored throughout the experiment and the latter wasmaintained with a water heating pad.

RF coil setup

For the 1H and 23Nameasurements we used a linear double tunedvolume resonator with an inner diameter of 7 cm from Bruker(Ettlingen, Germany). Additionally, we placed a surface coil thatoperated at the resonance frequency of 35CI at 9.4 T of 39.2 MHzon the head of the animal. This coil was constructed from silverwire with 3mm diameter. A single loop of 35mm diameter wasbent on a 35mm plexiglass half cylinder to achieve an optimalfilling factor for a rat head. The 35Cl coil was built large enough tocover the whole brain for sure and glued to the plexiglass halfcylinder to which it was bent. That fact and a mark on the animalcradle together with laser controlled stereotactic positioning canensure a repeatable placement. A variable capacitor of 1–40 pFand two fixed capacitors of 100 pF each in parallel tuned the coilto the resonance frequency of 39.2MHz. A balanced network oftwo capacitors of 22 pF and one variable capacitor of 1–40 pF

& Sons, Ltd. www.interscience.wiley.com/journal/nbm

93

S. KIRSCH ET AL.

594

connected the circuit to the signal line. All variable capacitorsused were purchased from Voltronics1 (Denville, NJ, USA), whilethe fixed capacitors were purchased from American TechnicalCeramics1 (Huntington Station, NY, USA). Figure 1 shows thecircuit diagram of the 35Cl coil.The unloaded Q factor, measured as the resonance frequency

divided by the difference between the frequencies of theresonance curve at �3 dB, was 138. The loaded Q-value with aglass vial filled with 12ml of 0.9% NaCl solution was 71. Nocoupling with the 1H-23Na volume resonator was observed. All RFcoils were operated in transmit/receive mode.

MR imaging

Proton imaging was performed using a double contrast multislice multi echo (MSME) sequence with TR¼ 2000ms, TE1¼ 13msand TE2¼ 65ms (two images per slice). The field of view (FOV)was 64� 64mm2 at a matrix of 256� 256 with 9 coronal slices of3mm thickness and an inter-slice distance of 3.5mm. The totalmeasurement time (TA) was 6min 24 s.The 23Na and 35Cl imaging was done using a slice selective ultra

short echo time (UTE) pulse sequence with radial k-spaceacquisition (11,29). For both nuclei 3 coronal slices withFOV¼ 64� 64mm2, matrix of 64� 64, slice thickness¼ 3mmand an inter-slice distance of 3.5mm were measured. The 3 sliceswere matched with the slice positions of the corresponding1H images by means of the scanner software Paravision1 5. Theparameters for the 23Na imaging were TR¼ 40ms, TE¼ 0.321ms,0.3ms gauss excitation rf pulse, readout bandwidth¼ 25 kHz/FOV,number of averages¼ 225 and TA¼ 30min 17 sec. For the35Cl imaging the following parameters were used: TR¼ 40ms,TE¼ 0.448ms, 0.24ms gauss excitation rf pulse, readout band-width¼ 25 kHz/FOV, number of averages¼ 455 and TA¼ 1h 1min.The optimal transmit power for the excitation rf pulse was

determined by manually adjusting the attenuation of a referencerf pulse in a FID experiment with same TR as used in the 23Na and35Cl MRI sequence in order to obtain maximum signal. The signalto noise ratio (SNR) was calculated by dividing the differencebetween the mean signal of a region of interest (ROI) and the

Figure 1. Surface rf coil for the 35Cl MRI.

www.interscience.wiley.com/journal/nbm Copyright � 201

mean magnitude of the background noise by the standarddeviation of the noise.

Measurement of the transversal relaxation time

The transversal relaxation time T2 of 35Cl was measured on thephantom and in vivo using an unlocalized spin echo sequence ofthe type: (a/2)x - te/2 -Gs(t) - ax -Gs(t) - te/2 - acq. Here, a¼ flip angleand x¼ phase of the rf pulse, Gs(t)¼ spoiler gradients inx,y,z-direction with duration t, te¼ variable time delay, TE¼ echotime¼ teþ 2t and acq denotes the data acquisition. The a/2 rfpulse of 100ms durationwas adjustedmanually in a FID experimentin order to give maximum signal. The spoiler gradients Gs(t) with astrength of 148mT/m and a duration of 250ms dephasemagnetization components with �908<a< 908 and thereforeavoid non-echo signal contamination. In a first measurement33 increments of the echo time TE in the range of 2ms to 151mswere measured for the phantom and in vivo (n¼ 1). After receivingstimulating comments from the reviewers we performed a secondin vivo measurement on a different rat (n¼ 2) covering short TEvalues in the range of 0.65ms to 60ms with 40 increments. Furtherparameters for all measurements were TR¼ 200ms, number ofdummy scans¼ 32, number of averages¼ 600 (phantom) and 400(in vivo). After Fourier transformation of the time domain signal andmagnitude calculation, the integral of the resulting spectrum wasused as a measure of the signal intensity S. The signal intensity as afunction of the echo time TE was fitted by the monoexponentialfunction

SðTEÞ ¼ y0 þ S0 exp � TE

T2

� �(3)

and/or the biexponential function

SðTEÞ ¼ y0 þ S0 A exp � TE

T2;a

� �þ 1� Að Þ exp � TE

T2;b

� �� : (4)

Here, y0 denotes an offset in the signal intensity originatingfrom the integrated noise in the spectrum, S0 denotes the signalintensity at TE¼ 0 and A, (1–A) denote the amplitudes of the twosignal components.

Measurement of the longitudinal relaxation time

The unlocalized measurements of the T1 relaxation time wereperformed with the rf pulse sequences a) ax - Gs1(t) - tI - (a/2)x -acq and b) ax - Gs1(t) - tI - (a/2)x - te/2 - Gs2(t) - ax - Gs2(t) - te/2 - acq.Here, tI¼ variable time delay, TI¼ tþ tI denotes the inversiontime and Gs1¼ 148mT/m, Gs2¼ 111mT/m are the strengths ofthe spoiler gradients in x, y, z-direction with a duration of 250ms.Rf pulse sequence a represents an inversion recovery experimentwhereas rf pulse sequence b can be considered as an inversionrecovery experiment with a subsequent spin echo readout. Theadvantage of the rf pulse sequence b will be illustrated in thefollowing.With a surface coil it is not ensured that the magnetization is

fully inverted all over the sample by the a rf pulse. Therefore, theresulting signal progressionmay correspond to a superposition of(partial) inversion recoveries and saturation recoveries. As a resultof this superposition the initial data points of the resultingsignal progression can have values close to zero which is notdesirable when expecting a multi-exponential signal recoverywith very short T1 components. Considering a uniformlymagnetized sample with a single longitudinal relaxation time

0 John Wiley & Sons, Ltd. NMR Biomed. 2010; 23: 592–600

Figure 2. 1H, 23Na and 35Cl MRI on the phantom containing NaCl (small

flask) and KCl (surrounding flask). The images show the same slice. Note,the 1H and 23Na signal is measured with a volume rf resonator while the35Cl signal is measured with a surface rf coil.


T1 and assuming a homogenous rf excitation over the sample, thetransversal magnetization immediately after the (a/2)x rf pulse ofthe inversion recovery pulse sequence a and b is given by

Mþ TIð Þ ¼ Mx þ iMy

¼ iM0 sina

2

� �1� exp � TI

T1

� �� þ cos a sin

a

2

� �exp � TI

T1

� �� :

(5)

Equation (5) can be extended to the case of inhomogeneous rfexcitation by introduction of the local flip angle aj:

MþðTIÞ ¼ iM0 1� exp � TI

T1

� �� Xj

sinaj

2

� �"

þ exp � TI

T1

� �Xj

cosaj sinaj

2

� �#:

(6)

Assuming that in the most regions of the sample the conditionaj/2� 1808 is satisfied, the first term in the squared bracketscorresponds to a saturation recovery and the second term to anexponential decay with positive or negative sign depending onthe degree of inversion over the sample. If we can ensure thatmagnetization components with �908< aj< 908 do not con-tribute to the measured signal the sum over the j elements in thesecond term is negative and the measured signal decay shouldresemble an (partial) inversion recovery with initial non-zero datapoints. The exclusion of contributions from �908< aj< 908canbe achieved bymeans of the spoiled spin echo readout in rf pulsesequence b. According to Eq. (6) the monoexponential fitfunction for the rf pulse sequence a and b is

S TIð Þ ¼ y0 þ c1 1� exp � TI

T1

� �� þ c2 exp � TI

T1

� ��: (7)

The constant is c1 is positive in both rf pulse sequences. For rfpulse sequence a the constant c2 can be positive, negative orclose to zero whereas for rf pulse sequence b the constant c2 isnegative. Extension of Eq. (7) to the biexponential case gives

S TIð Þ ¼ y0 þ c1 1� A exp � TI

T1;a

� �� 1� Að Þ exp � TI

T1;b

� �� þ c2 A exp � TI

T1;a

� �þ 1� Að Þ exp � TI

T1;b

� �� (8)

The measurement parameters for the rf pulse sequence a andb were TR¼ 200ms, number of averages¼ 400, number ofdummy scans¼ 8 and TE¼ 0.65ms (sequence b). 33 incrementsof TI in the range of 1.025ms to 150ms were measured on thephantom and 45 increments of TI in the range of 0.325ms to100ms were measured in vivo. Calibration of the a/2 rf pulse andpost processing of the measured data was done in the same wayas in the T2 measurements.For all relaxation time measurements the Levenberg-

Marquardt fitting procedures were done using Origin1 (Origi-nLab, Northampton, USA).

5

RESULTS

The performance of the MR imaging sequence was tested on aphantom containing two flasks filled with a solution of NaCl andKCl. Figure 2 shows the 1H, 23Na and 35Cl images measured on the


phantom. The signal intensity in the 35Cl image shows left–rightsymmetry and decreases as the distance to the surface coilincreases. This behaviour corresponds to the usual sensitivityprofile of a surface coil.In the next step the multinuclear MRI of 1H, 23Na and 35Cl was

applied on the head of a healthy rat and on a rat displaying a focalcerebral infarction in the right hemisphere of the brain. Figure 3shows the results of the in vivomeasurements. In the T2 weighted1H images the area of infarction can be identified by thehyperintense areas in the right hemisphere of the brain. Thesignal enhancement is caused by liquid accumulation due toischemic swelling. Similar behaviour is observed in the 23Na and35Cl images. Note, that the 35Cl images were measured with asurface coil therefore mainly the brain of the rat is visible in thecorresponding images. For each slice, exemplary values for theSNR were determined for a ROI in the area of infarction and a ROIcontralateral to this area. The average SNRs are given in Table 1.The ratio of the two SNR values gives an estimate for the signalenhancement in the area of infarction. An average signalenhancement of a factor of 2.9� 0.6 (23Na) and of 2.2� 0.2 (35Cl)is observed in the area of infarction.Figure 4 shows the decay of the 35Cl signal owing to transversal

relaxation in the phantom and in the head of a living healthy rat.Note, that for reasons of clarity the T2 measurement with TEvalues in the range of 2ms to 151ms is not presented in Figure 4.The measured data were fitted using Equation (3) and (4) fromthe Methods section of this paper and the results are given inTable 2. For the phantom the monoexponential fit according toEq. (3) yields a T2 of (29.4� 0.4) ms. The fit shows excellentagreement with the measured data which is reflected in thecorrelation coefficient R2¼ 0.99936.The results of the curve fitting procedures on the in vivo data

are also given in Table 2. Compared to the phantommeasurement the in vivo data suggest a multi-exponentialsignal decay and the application of the biexponential fitfunction (4) reveals a short and a long T2 component. Theobtained T2 values differ for the two measurements coveringdifferent ranges of TE values (T2¼ 2.3ms and 20.7ms vs.T2¼ 1.3ms and 11.8ms).Figure 5 shows the results of the T1 measurements on the

phantom and in vivo. The results of the curve fitting proceduresare given in Table 3. Again, the phantom data suggest amonoexponential and the in vivo data a multi-exponential signalprogression. The monoexponential fit according to Equation (7)on the phantom data yields a T1 of (32.3� 0.5) ms for rf pulsesequence a and (31.7� 0.4) ms for rf pulse sequence b. For thein vivomeasurement the biexponential fit function (8) revealed ashort and a long T1 component with T1,a¼ (1.7� 0.4) ms andT1,b¼ (25.1� 1.4) ms.


95

Figure 3. In vivo 1H, 23Na and 35Cl images from the head of a rat. Columns a-e: 3 coronal slices from a healthy rat. Columns f-j: 3 coronal slices from a rat

with focal cerebral infarction in the right hemisphere of the brain. Columns d, e and i, j show a cropped overlay of the inverted 1H images and the

color-encoded 23Na and 35Cl images.

S. KIRSCH ET AL.

596

DISCUSSION

Measurement of the relaxation times T1 and T2

The measurements of the relaxation times T1 and T2 on thephantom revealed monoexponential behaviour. This observation


is consistent with the expectation that in the used KCl solutionthe chloride ions are in the extreme narrowing limit and hence noresidual quadrupolar interaction is active.The T1 measurements on the phantom show that the modified

inversion recovery pulse sequence b provides initial data pointswith higher magnitude compared to the conventional inversion


Figure 4. The decay of the 35Cl signal owing to transversal relaxation measured with a spin echo sequence on the phantom (left) and in vivo (right). For

the results of the curve fitting procedure see Tab. 2.

Table 1. SNR determined from in vivo measurements

SNR 23Na SNR 35Cl

In vivo (ROI in area of infarction) 13.6 21.1In vivo (ROI contralateral toarea of infarction)

4.7 9.7

Note, a direct comparison of the 35Cl and 23Na SNR is notpossible due to the different coil types and imagingparameters.


recovery sequence a. This ‘increase’ of the initial signal isachieved at the cost of discarding signal contributions from areaswhere the local flip angle aj satisfies the condition�908<aj< 908, i.e. areas where only a saturation of themagnetization is achieved.To the best of our knowledge in this study we present the first

in vivo T1 and T2 measurement of 35Cl. Despite the unlocalizednature of the rf pulse sequences used for the determination ofthe relaxation times, comparison of the anatomical 1H and the35Cl images suggest that almost the entire in vivo signalmeasured with the 35Cl surface coil originates from the brain ofthe rat. The results of the biexponential curve fitting procedure

Table 2. Results of the curve fitting procedure for the determination of the T2 relaxation time of the 35Cl signal. R2 is the correlationcoefficient. n denotes the identification number of the rat

PhantomIn vivo (n¼ 1)min(TE)¼ 2ms

In vivo (n¼ 2)min(TE)¼ 0.65ms

Monoexponential fit y0 [a.u.] 0.045� 0.004 0.29� 0.02 0.27� 0.01S0 [a.u.] 1.03� 0.005 1.32� 0.05 0.77� 0.02T2 [ms] 29.36� 0.4 10.63� 0.81 3.92� 0.27

R2 0.99936 0.96949 0.97868Biexponential fit y0 [a.u.] — 0.23� 0.01 0.22� 0.01

S0 [a.u.] — 2.19� 0.16 1.04� 0.03A — 0.62� 0.02 0.64� 0.02

T2,a [ms] — 2.33� 0.26 1.33� 0.11T2,b [ms] — 20.71� 1.6 11.81� 1.1

R2 — 0.99691 0.99785

5


with Equation (4) on the in vivo data show that the valuesobtained for T2,a and T2,b depend on the range and the minimumof TE used in the spin echo sequence. For min(TE)¼ 2ms thesignal decay is dominated by the slow relaxing component andtherefore the fit may not be accurate with respect to the fastcomponent. The observed change from T2,a¼ 2.3ms formin(TE)¼ 2ms to T2,a¼ 1.3ms for min(TE)¼ 0.65ms leads tothe expectation that further improvements in accuracy of theshort T2 component may be achieved when the minimum TE isfurther decreased. However, the hardware limited minimum TEwas 0.65ms.The determined values of the amplitudes A¼ 0.64 and

(1–A)¼ 0.36 in the T2 measurements show good agreement withthe theoretical values of 0.6 and 0.4 predicted by Equation (2). Adetailed analysis of Equation (2) as it is presented in Ref. (30)shows that a fast relaxing component is expected in context withthe amplitude of 0.6, which is consistent with our results. Theamplitudes of A¼ 0.26 and (1–A)¼ 0.74 determined fromthe T1 measurement are also in agreement with the expectedamplitudes of 0.8 and 0.2 from Equation (1). This observationsupports the expectation that for the most chlorine ionsin the rat brain the extreme narrowing condition is notsatisfied, and therefore these ions experience a non-vanishingquadrupolar interaction as a consequence of restricted molecularmobility.


97

Figure 5. The progression of the 35Cl signal owing to longitudinal relaxation measured with the conventional a and the modified inversion recovery

pulse sequence b on the phantom (left). In vivo signal progression measured with the inversion recovery pulse sequence b (right). For the results of thecurve fitting procedure see Tab. 3.

S. KIRSCH ET AL.

598

However, an alternative explanation for the multi-exponentialbehaviour of the signal could be a superposition of 35Cl signalsfrom different compartments or tissue types with varying T1 andT2 values. In this scenario the measured signal decay represents acomplicated superposition of mono- and/or biexponential signaldecays and the fit functions (3,4) and (7,8) are no longer valid. Thebiexponential fit was chosen because it represents the simplestmulti-exponential fit function and doing this corresponds to anad hoc assumption with respect to the signal decay. Higher ordermulti-exponential fits appear not useful since, depending on theamount of data points measured, the convergence of these fits isnot ensured. Therefore, care should be taken when interpretingthe results of the relaxation time measurements. The measuredvalues for the in vivo relaxation times should be considered asestimates which allow optimization of the MRI pulse sequenceparameters.Future 35Cl MRI/MRS studies employing multiple-quantum

filtered rf pulse sequences may allow separation of 35Cl nucleiwhich are exposed to a residual quadrupolar interaction. Sinceresidual quadrupolar interactions are the consequence ofrestricted molecular mobility, a discrimination of different35Cl pools may then be possible.Paramagnetic chemical shift reagents are known to allow

separation between extra- and intracellular sodium (32) and

Table 3. Results of the curve fitting procedure for the determination of the T1 relaxation time of the 35Cl signal. R2 is the correlationcoefficient. n denotes the identification number of the rat

Phantom rf pulsesequence a

Phantom rf pulsesequence b

In vivo (n¼ 2) rf pulsesequence b

Monoexponential fit y0 [a.u.] 0.01� 0.004 ! 0 0.25� 0.02c1 [a.u.] 1.0� 0.006 1.01� 0.01 0.67� 0.02c2 [a.u.] �0.37� 0.007 �0.67� 0.01 �0.32� 0.03T1 [ms] 32.33� 0.49 31.69� 0.39 15.34� 1.04

R2 0.99934 0.99836 0.96394Biexponential fit y0 [a.u.] — — 0.26� 0.01

c1 [a.u.] — — 0.72� 0.02c2 [a.u.] — — �0.41� 0.02

A — — 0.26� 0.02T1,a [ms] — — 1.73� 0.39T1,b [ms] — — 25.07� 1.4

R2 — — 0.99359


chloride (33). In future studies, the use of chemical shift reagentsand combined 23Na and 35Cl MRI/MRS may allow non-invasiveinvestigation of the extra- and intracellular sodium and chloridesignal under different physiological and pathological conditionsin vivo.The results of our measurements of the in vivo relaxation times

of 35Cl indicate the presence of signal components with veryshort T2 in the rat brain. Physiological processes associated withthe infraction may result in exchange or redistribution of chlorideions between compartments with very short and long relaxationtimes, thus giving a contribution to the 35Cl signal intensity in theMR images which does not depend on the concentration ofchloride ions. On the other hand, Boada et al. (6) showed that23Na MRI pulse sequences with radial k-space acquisition andvery short echo time of TE< 0.5ms provide accurate estimates forthe tissue sodium concentration, even if short T2 componentswith T2 � 2ms are present. Assuming equal mobility of sodiumand chloride ions, theory predicts shorter transverse relaxationtimes for 35Cl (34). Therefore, for 35Cl MRI it is not obvious that theargumentation of Boada et al. holds and that the 35Cl signalintensity exclusively represents the concentration of chlorideions.Another source of concentration independent signal change in

the presented images may result from different saturation of the



longitudinal magnetization of 23Na and 35Cl. Depending on thevalue of T1 and the TR used both nuclei may be imaged with adifferent steady-state magnetization. An increase of long T1components (T1 of 23Na �34ms (31), T1 of 35Cl �25ms) as theresult of liquid accumulation in the area of infarction may yielddifferent signal intensities which are not related to changes inconcentration.

MR imaging

The signal intensity of 35Cl, although expected to be approxi-mately 9.6 times lower than that of 23Na, was sufficiently high toperform in vivoMRI in a reasonable time. Since we used a custommade surface coil, the sensitivity decreased with increasingdistance (an inherent drawback of surface coils in general).Nevertheless, tissue penetration was sufficiently strong togenerate complete coronal sections of the rat brain. The coilsetup and the parameters of the MRI pulse sequence weredesigned to ensure roughly equivalent image quality in the23Na and 35Cl images with respect to SNR and resolution.Therefore, a direct comparison of the SNR of the 23Na and35Cl images does not differ by a factor of 9.6 as one would predictfrom the theoretically estimated signal intensity ratio. Futurestudies which involve quantification may use B1 sensitivity mapsin order to correct the images of the 35Cl surface coil.We observed a clear regional increase of the 35Cl signal in the

occluded hemisphere affecting cortical and subcortical areas.Although we cannot exclude methodological reasons such aschanges in T2 (this possibility does not appear probable in view ofthe short TE of 0.448ms and the expected liquid accumulation inthe area of infarction) we consider the increase of the 35Cl signalbeing mainly due to changes of the ion concentration in thetissue because it fits with the extent of ischemic damage to beexpected in the filament occlusionmethod and it goes alongwithan increase of the 23Na signal intensity. Previously, an increase ofthe tissue Naþ concentration in the early phase of focal ischemiahas been described in a subhuman model of focal cerebralischemia, and similar changes were found at later time points instroke patients (9). This increase of the Naþ concentration may bedue to several reasons including the loss of parenchymal cells (9)and formation of vasogenic brain edema in still viable tissuesurrounding the ischemic necrosis. In either case there is anexpansion of the extracellular space and accumulation of fluidcontaining high concentrations of Naþ and Cl�. Unfortunately,the present tools do not provide absolute values of the Cl�

concentration in the tissue. We, therefore, related the signalintensity to that in the mirror region in the contralateralnon-ischemic hemisphere and obtained a 2.2 fold increase. Theincrease of the 23Na signal was considerably higher, amounting to2.9 fold resulting in a ratio of 0.75 (35Cl to 23Na). Actually, this ratiois in perfect agreement with the ratio present in the extracellularfluid (normal plasma concentrations of Cl� and Naþ are in therange of 110 and 140mMol l�1, respectively). This, again, arguesin favor of an increased Cl� ion concentration underlying theincrease of the 35Cl signal in ischemic areas.With Naþ being themajor cation in the extracellular space there is

usually a tight coupling between Naþ movement and flux of water,either over cell membranes leading to cell swelling or over thecerebroendothelial lining, the so-called blood–brain barrier, leadingto the development of vasogenic brain edema. However, aremarkable discrepancy in the changes of the Naþ signal and the


apparent diffusion coefficient (ADC) for water has been reported inearly focal brain ischemia (9). While the Naþ signal was found toincrease steadily during the first 10 hours after establishment ofischemia the ADC for water decreased almost immediately uponvessel occlusion and stayed constant thereafter. Similarly, in skeletalmuscle the ADC for the tissue Naþ concentration was found todecrease after 4h of ischemia while that of water did not. Thus, thereis again a clear discrepancy between the behavior of the Naþ andwater signals. The reason(s) for this discrepancy is not yet clear.However, monitoring of the Cl� signal may prove worthwhile inunderstanding thepathophysiological changes of ion concentrationsand water flux in focal ischemia, trauma, or tumors. In addition,further developments in Cl� signal recording may help to betterunderstand regulation of neuronal excitability, and it may also provehelpful in characterizing the emerging role of Cl� ions in diseases ofdifferent organs including the central nervous system (35).

REFERENCES

1. Hilal SK, Maudsley AA, Simon HE, Perman WH, Bonn J, Mawad ME,Silver AJ, Ganti SR, Sane P, Chien IC. In vivo NMR imaging of tissuesodium in the intact cat before and after acute cerebral stroke. AJNRAm. J. Neuroradiol. 1983; 4(3): 245–249.

2. Moseley ME, Chew WM, Nishimura MC, Richards TL, Murphy-Boesch J,Young GB, Marschner TM, Pitts LH, James TL. In vivo sodium-23magnetic resonance surface coil imaging: observing experimentalcerebral ischemia in the rat. Magn. Reson. Imaging. 1985; 3(4): 383–387.

3. Hilal SK, Maudsley AA, Ra JB, Simon HE, Roschmann P, Wittekoek S,Cho ZH, Mun SK. In vivo NMR imaging of sodium-23 in the humanhead. J. Comput. Assist. Tomogr. 1985; 9(1): 1–7.

4. Perman WH, Thomasson DM, Bernstein MA, Turski PA. Multiple short-echo (2.5-ms) quantitation of in vivo sodium T2 relaxation. Magn.Reson. Med. 1989; 9(2): 153–160.

5. Ra JB, Hilal SK, Oh CH, Mun IK. In vivo magnetic resonance imaging ofsodium in the human body. Magn. Reson. Med. 1988; 7(1): 11–22.

6. Boada FE, Christensen JD, Huang-Hellinger FR, Reese TG, Thulborn KR.Quantitative in vivo tissue sodium concentration maps: the effects ofbiexponential relaxation. Magn. Reson. Med. 1994; 32(2): 219–223.

7. Constantinides CD, Kraitchman DL, O’Brien KO, Boada FE, Gillen J,Bottomley PA. Noninvasive quantification of total sodium concen-trations in acute reperfused myocardial infarction using 23Na MRI.Magn. Reson. Med. 2001; 46(6): 1144–1151.

8. Ouwerkerk R, Bleich KB, Gillen JS, Pomper MG, Bottomley PA. TissueSodium Concentration in Human Brain Tumors as Measured with23Na MR Imaging. Radiology. 2003; 227: 529–537.

9. Thulborn KR, Davis D, Snyder J, Yonas H, Kassam A. Sodium MRimaging of acute and subacute stroke for assessment of tissueviability. Neuroimaging Clin. N. Am. 2005; 15(3): 639–653.

10. Jerecic R, Bock M, Nielles-Vallespin S, Wacker C, Bauer W, Schad LR.ECG-gated 23Na-MRI of the human heart using 3D-radial projectiontechnique with ultra-short echo times. Magn. Reson. Mater Phy. 2004;16: 297–302.

11. Nielles-Vallespin S, Weber MA, Bock M, Bongers A, Speier P, Combs SE,Wohrle J, Lehmann-Horn F, Essig M, Schad LR. 3D radial projectiontechnique with ultrashort echo times for sodium MRI: Clinical appli-cations in human brain and skeletal muscle. Magn. Reson. Med. 2007;57: 74–81.

12. Nagel AM, Laun FB, Weber MA, Matthies C, Semmler W, Schad LR.Sodium MRI using a density adapted 3D radial acquisition technique.Magn. Reson. Med. 2009; 62(6): 1565–1573.

13. Augath M, Heiler P, Kirsch S, Schad LR. In vivo 39K, 23Na and 1H NMRimaging using a triple resonant RF coil setup. J. Magn. Reson. 2009;200: 134–136.

14. Jentsch TJ, Stein V, Weinreich F, Zdebik A. Molecular structure andphysiological function of chloride channels. Physiol. Rev. 2002; 82:503–568.

15. Cottam GL, Ward RL. 35Cl-NMR studies on zinc-pyruvate kinasecomplexes. Arch. Biochem. Biophys. 1969; 132(1): 308–314.


599

S. KIRSCH ET AL.

600

16. Norne JE, Hjalmarsson SG, Lindman B, Zeppezauer M. Anion bindingproperties of human serum albumin from halide ion quadrupolerelaxation. Biochemistry. 1975; 14(15): 3401–3408.

17. Jonsson B, Lindman B. 35Cl NMR study of the interaction of chloridewith arginine, histidine and lysine. FEBS Lett. 1977; 78(1): 67–71.

18. Price WS, Kuchel PW, Cornell BA. A 35Cl and 37Cl NMR study of chloridebinding to the erythrocyte anion transport protein. BiophysicalChemistry. 1991; 40(3): 329–337.

19. Andersson T, Thulin E, Forsen S. Ion binding to cytochrome c studiedby nuclear magnetic quadrupole relaxation. Biochemistry. 1979;18(12): 2487–2493.

20. Falke JJ, Pace RJ, Chan SI. Direct observation of the transmembranerecruitment of band 3 transport sites by competitive inhibitors. A35Cl NMR study. J. Biol. Chem. 1984; 259(10): 6481–6489.

21. Brauer M, Spread CY, Reithmeier RAF, Sykes BD. 31P and 35Cl nuclearmagnetic resonance measurements of anion transport in humanerythrocytes. J. Biol. Chem. 1985; 260(21): 11643–11650.

22. Gupta RK, Dowd TL, Spitzer A, Barac-Nieto M. 23Na, 19F, 35Cl and 31Pmultinuclear magnetic resonance studies on perfused rat kidney. Ren.Physiol. Biochem. 1989; 12(3): 144–160.

23. Jelicks LA, Gupta RK. Multinuclear NMR studies of the Langendorffperfused rat heart. J. Biol. Chem. 1989; 264(26): 15230–15235.

24. Harris RK, Becker ED, Cabral de Menezes SM, Goodfellow R, Granger P.NMR Nomenclature. Nuclear spin properties and conventions forchemical shifts. Pure Appl. Chem. 2001; 73(11): 1795–1818.

25. Burstein D, Fossel ET. Intracellular sodium and lithium relaxation timesin the perfused frog heart. Magn. Reson. Med. 1987; 4: 261–273.


26. Parrish TB, Fieno DS, Fitzgerald SW, Judd RM. Theoretical basis forsodium and potassium MRI of the human heart at 1.5T. Magn. Reson.Med. 1997; 38: 653–661.

27. Hubbard PS. Nonexponential nuclear magnetic relaxation byquadrupole interactions. J. Chem. Phys. 1970; 53(3): 985–987.

28. Woitzik J, Schneider UC, Thome C, Schroeck H, Schilling L.Comparison of different intravascular thread occlusion models forexperimental stroke in rats. J. Neurosci. Meth. 2006; 151(2): 224–231.

29. Robson MD, Gatehouse PD, Bydder M, Bydder GM. Magnetic reson-ance: an introduction to ultrashort TE (UTE) imaging. J. Comput. Assist.Tomogr. 2003; 27(6): 825–846.

30. Bull TE. Nuclear magnetic relaxation of spin-3/2 nuclei involved inchemical exchange. J. Magn. Reson. 1972; 8: 344–353.

31. Goodman JA, Kroenke CD, Bretthorst GL, Ackerman JJH, Neil JJ.Sodium ion apparent diffusion coefficient in living rat brain. Magn.Reson. Med. 2005; 53: 1040–1045.

32. Hutchinson RB, Shapiro JI. Measurement of intracellular sodium withNMR methods. Con. Magn. Reson. 1991; 3: 215–236.

33. Diven CF, Wang F, Abukhdeir AM, Salah W, Layden BT, Geraldes CFGC,Freitas DM. Evaluation of [Co(gly)3]

� as a 35Cl� NMR shift reagent forcellular studies. Inorg. Chem. 2003; 42: 2774–2782.

34. Deverell C. Nuclear magnetic shielding of hydrated alkali andhalide ions in aqueous solution. Molec. Physics. 1969; 16(5):491–500.

35. Puljak L, Kilic G. Emerging roles of chloride channels in humandiseases. Biochim. Biophys. Acta. 2006; 1762(4): 404–413.


in vivo chlorine-35, sodium-23 and proton magnetic resonance imaging of the rat brain

Documents