J. Biomedical Science and Engineering, 2009, 2, 210-215
doi: 10.4236/jbise.2009.24034 Published Online August 2009 (http://www.SciRP.org/journal/jbise/
JBiSE
).
Published Online August 2009 in SciRes. http:// www.scirp.org/journal/jbise
Changes in cerebral perfusion detected by dynamic
susceptibility contrast magnetic resonance imaging:
normal volunteers examined during normal breathing
and hyperventilation
Ronnie Wirestam1*, Christian Engvall2, Erik Ryding3, Stig Holtås4, Freddy Ståhlberg1,4, Peter Reinstrup2
1Department of Medical Radiation Physics, Lund University, Lund, Sweden; 2Department of Anaesthesiology & Intensive Care,
Lund University Hospital, Lund, Sweden; 3Department of Clinical Neurophysiology, Lund University Hospital, Lund, Sweden;
4Department of Diagnostic Radiology, Lund University, Lund, Sweden.
Received 3 March 2009; revised 2 April 2009; accepted 24 April 2009.
ABSTRACT
Global cerebral perfusion parameters were
measured using dynamic susceptibility contrast
magnetic resonance imaging (DSC-MRI) in eight
healthy volunteers examined during normal
breathing and spontaneous hyperventilation.
DSC-MRI-based cerebral blood flow (CBF) de-
creased during hyperventilation in all volun-
teers (average decrease 29%), and the corre-
sponding global CBF estimates were 73±19ml/
(min100g) during normal breathing and 52±
7.9ml/(min100g) during hyperventilation (mean
±SD, n=8). Furthermore, the hypocapnic condi-
tions induced by hyperventilation resulted in a
prolongation of the global mean transit time
(MTT) by on average 14%. The observed CBF
estimates appeared to be systematically over-
estimated, in accordance with previously pub-
lished DSC-MRI results, but reduced to more
reasonable levels when a previously retrieved
calibration factor was applied.
Keywords: Magnetic Resonance Imaging; Perfu-
sion; Cerebral Blood Flow; Mean Transit Time;
Hypocapnia
1. INTRODUCTION
The use of dynamic susceptibility contrast magnetic
resonance imaging (DSC-MRI) for assessment of perfu-
sion-related parameters is promising [1,2], but the con-
cept shows a number of methodological complications.
For example, accurate registration of the arterial input
function (AIF), i.e. the concentration-versus-time curve
in an appropriate tissue-feeding artery, is hampered by
arterial signal saturation [3] and local geometrical distor-
tion [4] at peak concentration, as well as by partial-
volume effects [5]. Furthermore, the T2* relaxivities of
the paramagnetic contrast agent are, most likely, differ-
ent in arterial and tissue environments, and the non-lin-
ear relationship in whole blood between transversal re-
laxation-rate change (R2*) and contrast-agent concen-
tration needs to be considered when gradient-echo pulse
sequences are used [6]. Attempts to achieve absolute
quantification of perfusion parameters by standard DSC-
MRI have typically been characterized by overestimated
absolute values of cerebral blood volume (CBV) and
cerebral blood flow (CBF) [1,3,7,8,9], and these obser-
vations are attributed, at least in part, to a correspond-
ingly underestimated arterial concentration time integral.
Hence, most existing implementations of DSC-MRI
provide perfusion parameters only in relative terms.
Reproducible absolute quantification of CBF is indeed
desirable, for example, in the follow-up of tumour or
stroke therapy, for determining tissue at risk in acute
ischaemic stroke and when a global change in CBF can
be expected. The potential of DSC-MRI for absolute or
semi-absolute CBF quantification is not yet fully estab-
lished, and additional information is warranted. Hence,
in order to further investigate the usefulness of DSC-
MRI for absolute quantification as well as for detection
of controlled changes in perfusion, estimates of global
CBF and mean transit time (MTT) were acquired in a
group of normal volunteers examined during normal
breathing and spontaneous hyperventilation. The pri-
mary aim of this study was to test the capability of
DSC-MRI to detect and quantify changes in global cere-
bral perfusion caused by spontaneous hyperventilation.
2. METHODS
2.1. Subjects and Experimental Procedure
Eight healthy volunteers (mean age 33 years) were in-
*Corresponding author.

R. Wirestam et al. / J. Biomedical Science and Engineering 2 (2009) 210-215 211
SciRes Copyright © 2009 JBiSE
cluded in the study (Table 1). Each subject was exam-
ined by DSC-MRI during normal breathing and hyper-
ventilation on different occasions. The time interval be-
tween the two DSC-MRI experiments ranged from 3 to
35 days (mean time interval 16.4 days), and every sec-
ond subject started with the normocapnic conditions.
Hypocapnia was induced by spontaneous hyperventila-
tion under external guidance. During the DSC-MRI ex-
periments, end-tidal pCO2 (ETCO2) was monitored and
the subjects were at rest in the supine position, breathing
normal air with addition of extra O2 to a total level of
50%, with their eyes open and supplied with earplugs.
The study was approved by the local ethics committee,
and written informed consent was obtained from each
volunteer.
2.2. DSC-MRI Experiment
DSC-MRI was performed using a 1.5 T MRI whole-
body unit (Siemens Magnetom Vision, Siemens Medical
Systems, Erlangen, Germany). At each DSC-MRI ex-
periment, the subject received 0.2 mmol/kg bodyweight
of a gadobutrol MRI contrast agent (Gadovist 1.0,
Schering AG, Germany), administered into a peripheral
arm vein at an injection rate of 3 ml/s and followed by a
saline flush.
The first passage of the contrast-agent bolus through
the brain was tracked using dynamic gradient-echo echo-
planar imaging (GRE-EPI) during approximately 75 s at
a temporal resolution of 1.65 s. Ten slices with a slice
thickness of 8 mm were recorded and the imaging pa-
rameters were as follows: Echo time 54ms, matrix size
128×128 and field of view 250×250 mm2.
2.3. Post-Processing and Data Analysis
Estimates of CBF in ml/(min100g) were calculated ac-
cording to (1):








00
0
)()(
)(max)(
dttRdttC
tRdttC
kCBF
artery
H (1)
The tracer concentrations in tissue (C) and in artery
(Cartery) were calculated (in arbitrary units) using the re-
lationship C(t)  –ln[S(t)/S0], where S(t) is the signal at
time t and S0 is the baseline signal observed before arri-
val of the contrast-agent bolus [1,2]. The constant kH =
(1-Hlarge)/[(1-Hsmall)] was set to 0.705 ml/g in the pre-
sent study [1]. Hlarge and Hsmall are the haematocrit values
in large and small vessels, respectively, and  is the
brain density. R(t) is the tissue residue function, obtained
by deconvolution of the measured tissue concentration
time curve with the AIF, and max [R(t)] is the peak value
of this function. Deconvolution was performed using a
singular value decomposition algorithm.
The area under the AIF curve, i.e. the time integral of
the arterial concentration Cartery(t), was determined from
the same arterial locations in both the normal and the
hyperventilation case. A correction for the combined
consequences of partial-volume effects, arterial signal
saturation and signal displacement (due to local geomet-
rical distortion) at peak concentration was applied. The
employed correction resembled the approach described
by Knutsson et al. [9], although in the present study a
large brain-feeding artery (the internal carotid artery),
rather than the superior sagittal sinus, was used for the
AIF time-integral rescaling.
The rescaling procedure was based on the combined
concentration-versus-time information from the large
brain-feeding artery (showing distorted curve shape at
peak concentration due to partial-volume effects, signal
saturation and/or local geometric distortions) and a
smaller artery used as the AIF in the deconvolution pro-
cedure (assumed to show a reasonable curve shape but
suffering from an underestimated area under curve). The
concentration curve from the smaller artery (in practice
obtained from pixels very close to the middle cerebral
artery or its branches) was rescaled, with retained shape,
to fit the flanks and baseline of the distorted curve from
the large artery. The time integral of the rescaled
small-artery curve was used in (1) as an approximation
to the true concentration time integral of the large
Table 1. Volunteer data (sex, age, ETCO2 levels) and observed whole-brain average DSC-MRI CBF estimates.
ETCO2 [kPa] CBF [ml/(min 100g)]
Volunteer no. Sex Age
[years] Normocapnia Hypocapnia Normocapnia Hypocapnia
1 M 40 6.0 4.6 106 49.0
2 M 39 5.6 4.1 79.7 61.8
3 M 35 5.0 3.7 70.1 53.6
4 M 31 5.7 2.7 84.1 61.9
5 M 31 5.8 3.1 63.0 48.1
6 M 30 5.8 4.4 52.4 46.6
7 M 30 5.6 4.0 84 58.1
8 M 29 5.3 3.6 47.5 39.9

212 R. Wirestam et al. / J. Biomedical Science and Engineering 2 (2009) 210-215
SciRes Copyright © 2009 JBiSE
brain-feeding artery.
The retrieval of R(t) by deconvolution also allows for
calculation of the mean transit time using Zierler’s area-
to-height relationship [1,10,11]:

 )(max
)(
0
tR
dttR
MTT



Parametric CBF and MTT maps were calculated, in
absolute terms, using (1) and (2), respectively. For CBF
as well as MTT, whole-brain average estimates were
calculated as the mean of all brain-tissue voxel values in
the 10 slices obtained by DSC-MRI. Large-vessel con-
tributions were eliminated by excluding all pixels with
values exceeding 2.5 times the average CBF value of the
entire volume [12]. Spurious MTT values were similarly
removed by excluding a small number of pixels with
values below 0.2 times the mean MTT value and above
2.5 times the mean MTT value of the entire volume.
2.4. Statistical Analysis
A Wilcoxon matched-pairs signed-ranks test was applied
to determine whether or not the ETCO2, CBF and MTT
values observed during normal breathing were signifi-
cantly different from those observed during hyperventila-
tion. The CBF-versus-ETCO2 and MTT-versus-ETCO2
relationships were evaluated by linear-regression analyses.
3. RESULTS
In all 8 volunteers, the DSC-MRI-based CBF estimates
decreased when the subject was hyperventilating (Table
1 & Figure 1a), and the mean CBF decrease during hy-
perventilation was 29%. Similarly, in all 8 subjects a
longer MTT was observed during hyperventilation (Fig-
ure 1b), with a mean MTT increase of 14%.
Average whole-brain CBF and MTT estimates at nor-
mocapnic and hypocapnic conditions, together with the
corresponding ETCO2 levels, are given in Table 2 (mean
±SD, n=8). The statistical analysis showed that ETCO2
levels as well as CBF and MTT estimates differed sig-
nificantly between normal-breathing conditions and hy-
perventilation (p<0.01).
The obtained relationship between CBF and ETCO2 is
displayed in Figure 2, indicating that CBF increased with
ETCO2. Figure 3 shows MTT versus ETCO2 and the
trendline suggests a slight decrease in MTT when ETCO2
increases. The significant prolongation of MTT observed
during hyperventilation implies, according to the central
volume theorem (MTT=CBV/CBF), that the relative de-
crease in CBF, induced by hyperventilation, was larger
than the corresponding relative decrease in CBV.
4. DISCUSSION
It is indeed encouraging that the expected decrease in
CBF during hyperventilation could be detected in all of
the volunteers. Analysis of the results from the whole
population showed that the CBF estimates obtained dur-
ing hyperventilation were significantly different from
those seen during normal breathing conditions (p<0.01),
and the observed average CBF decrease of 29% is in
quite reasonable agreement with previous studies [13,14].
The observed CO2 reactivity of CBF during hypocapnia,
corresponding to approximately 2.1% of reduction in
global CBF per mmHg change in ETCO2, is not at all
unreasonable, and the current estimate is in excellent
agreement with previous findings by Fortune et al. [13]
and Reinstrup et al. [15]. Other investigators have ob-
served a somewhat higher CO2 reactivity [e.g. 14,16],
corresponding to approximately 3% reduction in CBF
per mmHg change in ETCO2, but the characteristics of
previously investigated populations may have differed
with regard to, for example, sex, age and state of health,
and some previous studies were limited to grey matter.
110
100
Figure 1. Individual estimates of (a) cerebral blood flow (CBF) and (b) mean transit time (MTT),
measured by DSC-MRI in eight normal male subjects during normal breathing and hyperventilation.
CBF [ml/(min 100g)]
90
80
70
60
50
40
30
a b
MTT [s]
9
8
7
6
5
4
Normal
breathing
Normal Hyper-
ventilation
Hyper-
ventilation breathing

R. Wirestam et al. / J. Biomedical Science and Engineering 2 (2009) 210-215 213
SciRes Copyright © 2009 JBiSE
Table 2. Whole-brain average DSC-MRI CBF and MTT esti-
mates and the corresponding ETCO2 levels during normal
breathing and hyperventilation (mean±SD, n=8).
ETCO2
[kPa]
CBF
[ml/(min 100g)] MTT[s]
Normal breathing 5.6 ± 0.3273 ± 19 6.5 ± 0.65
Hyperventilation 3.8 ± 0.6452 ± 7.9 7.5 ± 0.49
Relative change[%] -33 -29 +14
Wilcoxon test p-value 0.008 0.008 0.008
Figure 2. Estimates of cerebral blood flow
(CBF) obtained during normal breathing and
hyperventilation as a function of end-tidal pCO2
(ETCO2).
Figure 3. Estimates of mean transit time (MTT)
obtained during normal breathing and hyperven-
tilation as a function of end-tidal pCO2 (ETCO2).
The observed prolongation of the MTT during hyper-
ventilation is related to a decreased vascular blood ve-
locity during hypocapnia [14], and this effect manifests
itself as a well-documented smaller relative decrease in
CBV than in CBF during the hypocapnic conditions in-
duced by hyperventilation [13,14]. Ito et al. [17] ad-
dressed this topic by investigating the relative impor-
tance of the arterial, capillary and venous blood-volume
fractions in hypocapnia. The authors concluded that
changes in human CBV during hypocapnia are caused
by changes in the arterial blood-volume component
without changes in the venous and capillary blood vol-
ume [17].
The absolute global CBF values obtained from the
present experiments were somewhat high, in accordance
with previous experimental DSC-MRI investigations
[1,3,7,8,9] and theoretical predictions [6]. Partial-volume
effects and local geometric distortions at peak concen-
tration [3,4] can be problematic, but these effects were
most likely reduced by the applied correction of the arte-
rial concentration time integral. Hence, the remaining
CBF overestimation is probably related to the funda-
mental problem with DSC-MRI in that the response to a
given contrast-agent concentration differs between large
vessels and the capillary environment [6]. In spite of the
applied AIF-area correction, accurate determination of
Cartery(t) and the associated time integral was difficult in
some cases, and the relatively large standard deviations
(SDs) seen in the DSC-MRI results might be a reflection
of this difficulty. Identification of an appropriate AIF
location is another crucial issue in CBF and MTT quan-
tification. It has recently been established that the de-
sired linear relationship between R2* and con-
trast-agent concentration in arteries can indeed be ob-
tained by careful selection of the AIF from pixels not
completely located inside the vessel [18]. In the present
study, we were aware of the AIF-selection guidelines
provided by Bleeker et al. for single-shot EPI [18], and
tried, as far as possible, to consider them in the AIF
identification procedure.
CBF [ml/(min 100g)]
ETCO2 [kPa]
Even if DSC-MRI turns out to provide inherently
overestimated CBF estimates they may still be useful,
provided that DSC-MRI results consistently can be
shown to exhibit a high degree of linear correlation with
a reference CBF technique, for example, Xe-133 SPECT
[9] or positron emission tomography (PET) [7]. The
possibility to rescale DSC-MRI-based perfusion esti-
mates by application of an appropriate calibration factor,
based on such comparative studies, has been suggested,
although it has been pointed out that the retrieval of a
universal conversion factor, applicable to a variety of
DSC-MRI implementations, may be challenging [7].
MTT [s]
ETCO2 [kPa]
A CBV calibration factor, applicable to the current
DSC-MRI setup, has previously been obtained, in the
same group of volunteers as examined in the present
study, using SPECT imaging of Tc-99m-labelled eryth-
rocytes as a reference CBV method [19]. This calibration
factor can theoretically be used also to appropriately
correct corresponding DSC-MRI-based CBF values,
since CBF=CBV/MTT (according to the central volume
theorem), provided that MTT values can be correctly
estimated. The global MTT estimates observed in the
present study (mean value 6.5 s at normoventilation)

214 R. Wirestam et al. / J. Biomedical Science and Engineering 2 (2009) 210-215
SciRes Copyright © 2009 JBiSE
were quite reasonable, and in accordance with previ-
ously published PET results from normal subjects. For
example, Kaneko et al. [20] observed MTT values of 6.1
s in grey matter and 8.1 s in white matter, and the large
study by Leenders et al. [21] showed CBV-to-CBF ratios
of 5.7 s in insular grey matter and 7.3 s in white matter.
Application of the calibration factor to the present data
resulted in a corrected whole-brain average CBF of ap-
proximately 42 ml/(min100g). Literature values of nor-
mal global CBF in humans at rest vary over a consider-
able range [22,23], but are typically between 40 and
50ml/(min100g) for the adult population. For example,
Knutsson et al. [9] obtained a whole-brain average CBF
of 40ml/(min100g) (in elderly normal subjects) by
Xe-133 SPECT, Slosman et al. [24] observed a global
CBF of 43ml/(min100g) in male volunteers (age interval
29-38 years), also by use of Xe-133 SPECT, Dörfler et
al. [22] reported a global CBF estimate of 48ml/
(min100g) based on extracranial sonography and Mat-
thew et al. [25] observed 40 ml/(min100g) using H2
15O
PET. Finally, Yonas et al. [26] employed stable xenon
computed tomography (Xe-CT) and extracted regional
CBF values of 92ml/(min100g) in the highest-flow
compartments, 54ml/(min 100g) in mixed-cortical re-
gions (calculated from linear-regression equations and
corresponding to the age of 33 years) and an age-inde-
pendent white-matter regional CBF of 20ml/(min100g).
In conclusion, DSC-MRI showed promising results in
the detection of controlled perfusion changes, induced
by spontaneous hyperventilation, in individual subjects.
In accordance with previously reported DSC-MRI ex-
periments, uncorrected absolute CBF values appeared to
be overestimated.
5. ACKNOWLEDGEMENTS
This study was supported by the Swedish Research Council (project no.
13514), the Swedish Cancer Society and the Crafoord Foundation,
Lund.
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