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1 Introduction

Although the adult human brain assumes only about 2 % of the total body weight, it receives nearly 15 % of the resting cardiac output and 20 % of the total body oxygen consumption. This high metabolic rate reveals the need to ensure a constant delivery of oxygen and energy-providing substrates at the capillary level and to remove the waste products of metabolism. Owing to a complex autoregulatory mechanism, the cerebral perfusion is maintained relatively constant over a wider range of mean arterial pressures. A complete interruption of brain–blood supply, however, leads immediately to neuronal impairments because of the limited availability of nutrition and energy reserves in the brain. Hence, perfusion parameters are important key factors involved in major cerebrovascular and other brain diseases. Important information about cerebral characteristics is given by measurement of the cerebral blood flow (CBF) which describes the rate of blood delivery to the brain parenchyma. Traditionally, the respective literature reports CBF units in ml blood/100 g of tissue per minute. Other authors use ml instead of gram tissue to describe the volumetric character especially with imaging techniques. Due to the fact that 1 g of brain tissue nearly corresponds to 1 ml, both values can be more or less used interchangeably.

Global average CBF values in middle-aged healthy human subjects are around 55 ml/100 g/min (Kety and Schmidt 1945). However, CBF values vary regionally: In cortical gray matter, the CBF is 60–100 ml/100 g/min (Slosman et al. 2001; Sokoloff et al. 1957) and around 20 ml/100 g/min for white matter (Law et al. 2000). It has been demonstrated that neuronal activity as well as CBF is closely coupled to brain metabolism (Pantano et al. 1984; Roy and Sherrington 1890). This may explain the generally higher blood flow in younger subjects which is typically exceeding values that are 50–85 % higher than those of adults (age: 6–7 years, Chiron et al. 1992). This age dependency of CBF is probably caused, at least in parts, by an age effect on the gray-to-white matter ratio. Additionally, a decline of CBF with age is described (Brody 1955).

Kety and Schmidt (1945) were the first who introduced a method for quantitative CBF measurement in the unanesthetized human. This method is based on the Fick principle. During the inhalation of the freely diffusible nitrous oxide, the brain perfusion was calculated by measuring the amount of gas removed from the blood by the brain per minute and dividing this by the arteriovenous difference of N2O. This method was an important step to study brain function in humans and has contributed fundamentally to our understanding of physiological and pathological cerebral processes. However, this first technique lacked the possibility to measure CBF in different brain subregions. This limitation has led to further developments in perfusion imaging techniques. As important examples, radioactive agents such as 133Xe and 85Kr were later used to measure their washout with external radiation detectors (Lassen and Ingvar 1961; Veall and Mallett 1967; Bruce et al. 1973; Obrist et al. 1975). Today, a wide variety of brain perfusion imaging techniques is available for preclinical and clinical research settings as well as clinical routine applications. CBF measurement using positron emission tomography (PET) became available in the 1970s, followed by other in vivo imaging techniques, like single-photon emission computed tomography (SPECT), magnetic resonance imaging (MRI), and transmission computed tomography (CT), to obtain CBF readouts (Hoeffner 2005). This chapter will focus on CBF PET with [15O]H2O as this represents the gold standard technique for in vivo CBF measurements (Carroll et al. 2002; Hoeffner 2005). One possible application for its usage is studying CBF in brain activation studies to localize brain segments involved in processing certain tasks. More important today, however, is the diagnostic potential of CBF PET imaging mainly in cerebrovascular disorders and oncology. Hence alterations of cerebral perfusion can theoretically be detected by quantitative CBF measurement leading to optimal therapy and providing estimates of the potentially salvable part of the affected brain.

2 Radiochemistry of [15O]H2O

Oxygen-15 can be produced by different nuclear reactions using both high- and low-energy protons as well as deuterons as projectiles. The 14N(d, n)15O nuclear reaction is the most efficient and economic production pathway and is therefore applied most frequently. For cyclotrons that do not have the possibility of accelerating deuterons, either the 15N(p, n)15O (Powell and O’Neil 2006) for low-energy protons (>3.7 MeV) or the 16O(p, pn)15O (Beaver et al. 1976; Krohn et al. 1986) for high-energy protons (>16.6 MeV) nuclear reactions can be used as an alternative.

Two methods are available for conversion of the 15O to [15O]H2O, the in-target production method and the out-of-target external conversion method. In case of the in-target production method, [15O]H2O is produced by either addition of small amounts of hydrogen to the target gas or direct irradiation of H2 16O. For the first approach, the 15N(p, n)15O nuclear reaction is used (Powell and O’Neil 2006), and the trace amounts of H2 in the target lead to the formation of [15O]H2O by radiolytic reactions which can be trapped in a cooled stainless steel loop. The possible radioactive impurities such as 11CH4 and 13NH3 (resulting from trace amounts of N-14 and O-16 in the target gas) are directed to the exhaust. By heating the loop, [15O]H2O vapor is released and can be trapped again in a sterile water or saline solution. The major drawbacks are the high costs for enriched 15N2 which is used as target material. An alternative approach for [15O]H2O in-target production is the direct irradiation of H2 16O in a liquid target (Mulholland et al. 1990; van Naemen et al. 1996). As mentioned above this method is feasible only with high-energy protons, a restriction that excludes this approach for commonly used cyclotrons in a clinical setting.

As an alternative to the in-target production of [15O]H2O, external out-of-target tracer production by catalytic conversion of a mixture of H2 and [15O]O2 has been demonstrated for all abovementioned nuclear reactions, i.e., 16O(p, pn)15O (Krohn et al. 1986), 15N(p, n)15O (Explora®H2O module, Siemens Healthcare, Erlangen, Germany), and 14N(d, n)15O (Clark et al. 1987; Clark and Tochon-Danguy 1991; Sajjad et al. 2000). For the 14N(d, n)15O nuclear reaction, N2 with up to 4 % O2 is irradiated with deuterons (≈7 MeV). In most cases, Pd is used as the catalyst; however, if the same target is used as well for production of C15O, it might be necessary to reduce the amount of O2 in order to minimize the production of toxic carrier CO. Under these conditions it has been shown that Pt as catalyst is superior (Berridge et al. 1990). The most simple production and application approach consists of an H2 supply that is connected to the target line via a T connector. The target gas/H2 mixture (2–10 %, Berridge et al. 1990; Sajjad et al. 2000) is then directed over the heated catalyst (≈170 °C, temperatures up to 450 °C have been reported as well (Berridge et al. 1990)), and the resulting [15O]H2O vapor is bubbled into a sterile reservoir containing water, saline, or preferably buffer since the catalyst might as well generate trace amounts of ammonia which un-buffered may lead to elevated pH values. The [15O]H2O-containing solution is then drawn up into a syringe and manually applied to the investigated subject. However, due to the relatively high radiation exposure of the medical personnel, it might be worth to automate the injection procedure. In a relatively simple setup, the injection is performed by means of two infusion pumps and a 4-port valve (Sajjad et al. 2000). A more sophisticated method uses a dialysis membrane to enable exchange of [15O]H2O with sterile water (Clark and Tochon-Danguy 1991) together with an infusion pump and several valves to enable automatic injection. This system is commercially available (Veenstra Instruments, Joure, Netherlands).

3 [15O]H2O Brain PET Data Generation

After a transmission scan for attenuation correction, the emission scan is acquired preferentially in 3D mode and initiated immediately before tracer administration. The tracer is injected as a fast bolus followed by a flush of inert saline solution. The amount of injected tracer in humans typically ranges between 550 and 1,000 MBq for an adult subject, but studies using up to 2.2 GBq were also found in the literature (Heiss et al. 2000).

For the bolus injection method, Kanno et al. (1991) investigated an optimal scan time for [15O]H2O to improve image quality and signal-to-noise ratios. A minimum scan duration of 90 s was recommended. For most applications, however, scan durations of 2–5 min are chosen. A typical protocol for a 5-min dynamic data acquisition is shown in Table 4.1.

Table 4.1 Acquisition protocol for a 5-min PET scan with [15O]H2O
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The advantageous short half-life of the tracer (122 s) enables the performance of multiple image acquisition scans in rapid sequence. Inter-scan intervals should, however, not be shorter than 15 min to allow isotope decay. Due to the fast radioactive decay, an on-site cyclotron or linear accelerator for isotope production becomes necessary. Further data analyses generally include the absolute quantification of CBF. For this purpose, the experimental setting implies arterial blood sampling in parallel to the PET acquisition. This is preferentially performed using an automated sampling system (e.g., ALLOGG AB blood sampler; Allogg Mariefred, Sweden) with a peripheral artery, e.g., the radial artery. Using such a device, arterial blood samples are continuously drawn at a constant speed with activity measurements for every 0.5 or 1 s. The blood sampler needs to be cross-calibrated to the PET scanner, allowing the decay-corrected blood data to be used as input function for kinetic modeling.

Obtaining arterial blood samples via the placement of an arterial catheter is accepted as the gold standard method for CBF quantification. However, there are some limitations, like invasiveness, complications for the patient, and sensitivity to errors (Hall 1971; Machleder et al. 1972). Aiming to replace the arterial canalization and corresponding input function, alternatives were suggested, like image-derived input functions (Zanotti-Fregonara et al. 2011) and methods without the need of any input function (Lammertsma 1994; Watabe et al. 1996). Another possibility to avoid arterial canalization is the use of arterialized venous blood, which is an often used method also with other PET tracers. With this technique, the hand of the patient or volunteer is heated well above 37 °C to achieve a shunting of arterial blood to the venous system (Wakita et al. 2000). Nevertheless, all alternative techniques have their drawbacks, and the CBF values obtained by them need to be handled with caution.

As for all brain PET tracers, the resulting PET data require a correlation with structural information from MRI (preferably 3D-T1 data). In this regard, the use of combined PET/MRI scanners might improve this situation in the future. Combined PET/MR imaging gives the opportunity for accurate registration and exact correlation of PET functional aspects with anatomical information from MRI. This will result in better image quality because of the comparably lower spatial resolution of PET in contrast to MRI and also shortens imaging times for the patients/anesthetized research animals.

4 Kinetic Modeling of CBF

The first method to measure CBF in humans was proposed by Kety and Schmidt in 1945 who utilized nitrous oxide for CBF detection. Later, other tracers like 133Xenon (Veall and Mallett 1967) were applied within this concept which was based on the Fick principle and a single-tissue compartment model. It states that the amount of a metabolically inert and freely diffusible gas that is taken up by a tissue per unit of time is equal to the product of the blood flow through that tissue and the difference between the amount of gas entering it via the arterial blood and the gas leaving in the venous blood.

The single-tissue compartment model describes the behavior of a freely diffusible tracer like [15O]H2O in tissue as shown in Fig. 4.1 and can be used to determine the local arterial blood flow in the brain (on a region of interest (ROI) or voxel basis). The model consists of two parameters that have to be estimated from the data of a dynamic PET scan (q(t)) and from the measured arterial input function c a(t).

Fig. 4.1
figure 1

Schematic description of a single-tissue compartment. V tissue is the anatomical tissue volume (1 cm3), and V d is the volume of the tissue compartment accessible to the tracer. The products of flow and concentration describe the amount of tracer which enters or leaves the tissue compartment per time unit. Please refer to the main text for further explanation of the other symbols

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Assuming that the transport of tracer from the vessel into the tissue compartment is fast (high permeability surface area product) compared to the delivery by the arterial blood flow, the tracer dynamics can be described by a one-tissue compartment model with one input function. The mass balance for the tracer (Fick principle) yields the differential equation

$$ \frac{\mathrm{d}q(t)}{\mathrm{d}t}={f}_{\mathrm{a}}{c}_{\mathrm{a}}(t)-{f}_{\mathrm{a}}{c}_{\mathrm{v}}(t) $$
(4.1)

where q(t) is the quantity of tracer per unit volume of tissue (kBq cm−3), f a is the local arterial blood flow per unit volume (ml min−1 cm−3), and c a(t) and c v(t) are the tracer concentrations in arterial and venous blood (kBq ml−3).

The tracer concentration in the venous blood (c v(t)) is related to the tracer concentration in the tissue space (q(t)) through the relative volume of distribution (V d (ml cm−3)) by the Kety–Schmidt assumption:

$$ q(t)={V}_{\mathrm{d}}{c}_{\mathrm{v}}(t) $$
(4.2)

reflecting the assumption that the concentrations in the water spaces of venous blood and tissue are always equilibrated (this assumption is not generally valid, because a diffusion limitation exists for [15O]H2O at low CBF rates).

From Eqs. (4.1) and (4.2), the differential equation

$$ \frac{\mathrm{d}q(t)}{\mathrm{d}t}={f}_{\mathrm{a}}{c}_{\mathrm{a}}(t)- kq(t) $$
(4.3)

is obtained with the washout constant k (min−1) defined to be

$$ k=\frac{f_{\mathrm{a}}}{V_{\mathrm{d}}}. $$
(4.4)

By estimating values of f a and k, the partition coefficient of the tracer can be calculated for every tissue voxel. An example of parameter estimation for a cortical region of interest based on the arterial input function is shown in Fig. 4.2.

Fig. 4.2
figure 2

Model parameter estimation. The arterial input function (a) and corresponding tissue response in a cortical ROI (b) after bolus injection of [15O]H2O are shown. Blood activity data were determined by taking arterial blood samples with a dedicated sampling device. The input function and tissue response are corrected for tracer arrival times and bolus dispersion. Further kinetic modeling is based on the resulting corrected curves (c)

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Equation (4.3) has the solution

$$ q(t)={f}_{\mathrm{a}}{e}^{- kt}\otimes {c}_{\mathrm{a}}(t) $$
(4.5)

where ⊗ describes the convolution of an exponential function with the arterial input function c a(t):

$$ {e}^{- kt}\otimes {c}_{\mathrm{a}}(t)={\displaystyle {\int}_0^tt{e}^{- kt}{c}_{\mathrm{a}}\left(t-\hat{o}\right)\mathrm{d}\hat{o}.} $$
(4.6)

The experimental design is further specified in a way that the tissue response, q(t), as well as the arterial, c a(t), input is measured and thus known for the duration of the experiment. Additionally, q(0) = 0.

[15O]H2O is the most commonly used tracer for brain perfusion imaging and CBF quantification with PET. However, there are some notable restrictions concerning the diffusion limitation of [15O]H2O. A study of Eichling et al. investigated the cerebral behavior of [15O]H2O after administration to rhesus monkeys. They found that only about 90 % of the injected tracer freely exchanges with the brain tissue, with even lower rates at higher flow rates. This incomplete first-pass extraction (80–90 % in gray matter structures) resulted in a slightly underestimation of CBF especially in high-flow regions (Eichling et al. 1974; Bolwig and Lassen 1975; Raichle et al. 1983).

In this respect, lipophilic gaseous tracers (like nitrous oxide and xenon) behave superior for CBF measurements as compared with [15O]H2O.

Further consideration should be done concerning partial volume effect that among others derive from relatively low spatial resolution of typically 6–10 mm in PET. The resulting spread out of signal is a phenomenon that introduces distortion effects in the targeted region and adjacent tissue (Links et al. 1996). The partial volume effect primarily leads to an increased bias for small anatomical structures (like vessels) in the brain (Rousset et al. 1998).

Since the early 1980s, several methods have been described to calculate CBF from PET measurements with [15O]H2O, including simplified techniques that include parameter fixation, e.g., fixation of the partition coefficient (Watabe et al. 1996). The partition coefficient of [15O]H2O is a parameter that has been investigated in many studies and varies from 0.77 to 1.05 ml/ml (Herscovitch and Raichle 1985; Iida et al. 1993; Kanno et al. 1991). A good approximation for the whole brain was considered by Herscovitch and Raichle to be 0.9 ml/ml (Herscovitch and Raichle 1985). However, CBF calculation methods with fixed values for V d are only applicable to identify global blood flow changes. In presence of regional blood flow deficits, CBF calculation results in incorrect values.

5 Role of PET for CBF Measurements

5.1 General Principles for CBF Measurements

A general distinction must be made between the behaviors of different CBF tracers. One class of techniques utilizes agents which are restricted to the intravascular space and do not interact with the nonvascular space. These can be, in cases of an intact blood–brain barrier, referred to as nondiffusible CBF tracers. Many brain imaging techniques, such as contrast-enhanced MR and CT, use these intravascular tracers to calculate CBF on the basis of the indicator dilution theory (Meier and Zierler 1954; Zierler 1962). In contrast, a direct tracer exchange from the arterial vascularity to the parenchyma occurs in case of freely diffusible CBF tracers and enables to give a direct measurement of parenchymal blood flow, as originally described by Kety (1951). This concept is utilized, for example, with [15O]H2O PET, Xenon–CT, and Xenon SPECT. Some basic characteristics appear to be necessary for a useful CBF measurement method: Ideally, the incorporated indicator should be early and completely mixed with blood and must stay identifiable for position and concentration in time of image acquisition. Further, the indicator should be metabolically inert and rapidly eliminated.

5.2 Advantages and Disadvantages of Perfusion Imaging Methods

Various modalities have been developed to obtain hemodynamic parameters in research and clinical settings. These include the older 133Xe inhalation method, PET, SPECT, X-ray computed tomography methods, and several MRI techniques. However, each technique has its own advantages and drawbacks. It depends on the study subject and the targeted question which method to choose appropriately. A review from Wintermark et al. (2005) gives a comparative overview about current brain perfusion measurement techniques and their clinical relevance. At this point, a brief overview is given on the role of [15O]H2O PET in this context.

5.2.1 Nuclear Medicine Methods

CBF can be measured after incorporation of radioactive agents which are detected outside the investigated subject with dedicated scintillation detectors. Based on this principle, tracers like [99mTc]HMPAO or [99mTc]ECD and 133Xe were commonly used for CBF measurements with SPECT (Barthel et al. 2001; Lass et al. 1998; Sakai et al. 1987). In comparison to the coincidence method in PET imaging, the detection of single photons is less sensitive. SPECT imaging with [99mTc]HMPAO or [99mTc]ECD only allows for semiquantitative CBF estimation (Markus 2004). In contrast, the 133Xe SPECT method relies on the Kety–Schmidt model (Kety and Schmidt 1945) and is considered to give quantitative measures of CBF (Wintermark et al. 2005). In several studies, however, a systematic CBF overestimation was reported in low-flow areas, as well as an underestimation of cortical CBF (Matsuda et al. 1996; Skyhøj Olsen et al. 1981).

However, the coincidence PET technique is commonly accepted to be the reference standard for CBF imaging. With PET, the tissue perfusion can be directly measured by using the diffusible radiotracer [15O]H2O. This method is well validated and combines several favorable properties. The tracer is easy to produce, and the fast acquisition time permits repetitive measurements with whole brain coverage. Additionally, the major advantage of CBF measurement with PET is the high accuracy for assessing quantitative parameter maps as well as a high reproducibility (Carroll et al. 2002; Matthew et al. 1993).

Additionally, PET imaging with 15O-labeled compounds is of special interest for studying cerebrovascular diseases and if a comprehensive view on brain hemodynamic is demanded. In addition to the determination of CBF with [15O]H2O, a successive 15O inhalation allows for quantitative determination of essential parameters of hemodynamics and energy metabolism like oxygen consumption (cerebral metabolic rate of oxygen; CMRO2) and oxygen extraction fraction (OEF) (Frackowiak et al. 1980; Ibaraki et al. 2004). Further, because of its binding to hemoglobin in red blood cells (Martin et al. 1987), radiolabeled CO is used as an intravascular tracer to measure the cerebral blood volume (CBV). Compared to other modalities, PET is the only technique which is able to gain all of these different functional parameters noninvasively and in 3D for the entire brain, a fact which allowed PET imaging to become the gold standard method in the field of brain circulation physiology and pathophysiology imaging (Hoeffner 2005).

In comparison to SPECT, which is widely accessible and a routine perfusion imaging tool, PET imaging is technically more demanding and its availability is limited by complex logistics. Not only a PET scanner, but also the constant access to a cyclotron producing the radiopharmaceutical online is required, a fact limiting the application of the method especially in emergency settings. A further restriction occurs in patients who will receive a thrombolytic therapy. Because quantitative PET preferentially requires invasive arterial blood sampling in order to obtain an input function for kinetic CBF modeling, this procedure is not applicable in these patients. As a consequence, [15O]H2O PET did not manage to become a clinical routine imaging tool in the acute stroke situation. Clinical applications instead mainly refer to chronic cerebrovascular disorders, brain tumors, and brain activation studies. Moreover, due to the wide acceptance of PET as standard for CBF visualization and quantification, the method is used as the reference to validate other brain perfusion imaging techniques, like perfusion-weighted or arterial spin labeling MRI (Zaro-Weber et al. 2010a, b; Chen et al. 2008)

Apart from [15O]H2O PET, CBF measurements with [11C]butanol have been suggested to be an alternative tracer for detection with PET. In comparison to radiolabeled water, butanol has the advantage of being permeable through the blood–brain barrier to 100 %. However, this tracer is not used in routine practice because of its complex and radiochemical synthesis (Herscovitch et al. 1987).

[14C]iodoantipyrine (IAP) autoradiography is another nuclear medicine perfusion imaging method used in preclinical research (Hatakeyama et al. 1992; Jay et al. 1988). As with butanol and H2O, IAP is also able to freely cross the blood–brain barrier. It is not metabolized and as such accumulates in the brain tissue depending on the regional CBF. After tracer application, the animals need to be sacrificed to prevent tracer diffusion and to autoradiographically determine the CBF at the time-point of tracer injection. Sequential arterial blood samples can be used for absolute CBF quantification. This autoradiographic ex vivo method provides accurate and high-resolution quantitative CBF values for a specific time-point and is therefore mainly used in small animal studies.

5.2.2 Computed Tomography Methods

The physical principle that underlies the CT technique is based on tissue-specific attenuation of X-rays that are directed to the body. The image contrast then resulted from variations in attenuation depending on tissue density. Due to similar densities in white and gray matter structures, this technique is not the ideal tool to image anatomical brain structures (Griffiths et al. 2001). However, with a bolus injection of a contrast agent, such as iodine, most prerequisites of the abovementioned indicator dilution theory are satisfied to measure blood flow in the brain. However, due to different acquisition hardware, acquisition protocols, varying post-processing protocols, and differences in the interpretation of perfusion CT data, a reliable CBF quantification remains challenging and varies widely between centers (Kudo et al. 2010). A further approach for CBF measurement uses inhaled Xenon to detect concentration changes of the substance (Pindzola and Yonas 1998). The lipophilic gas is soluble in water, and its X-ray attenuation is similar to that of iodine. As in 133Xe SPECT, the Xe–CT technique also utilizes the Kety–Schmidt method to calculate quantitative CBF maps with sufficient accuracy (Wintermark et al. 2005). Although newer CT scanners are able to achieve whole brain coverage, a main limitation of commonly used CT scanners derives from the limited anatomical coverage, which is restricted to few brain slices.

5.2.3 Magnetic Resonance Methods

Several methods for CBF estimation by means of MR had been developed. The most commonly used method for neuroimaging studies is the dynamic susceptibility contrast (DSC) MRI method. It relies on changes in relaxation time on T2*-weighted images. With the bolus of a paramagnetic contrast agent agents (e.g., gadolinium-DTPA) passing through the vascular system, a detectable signal loss occurs in T2*-weighted sequences. Mathematical conclusions were then drawn from the signal reductions to further calculate several perfusion or perfusion-related parameters including mean transit time (MTT), time to peak (TTP), relative cerebral blood volume (rCBV), and relative CBF (Ostergaard et al. 1996a, b). Contrast agents for MR imaging are not radioactive and relatively inexpensive as compared with PET and SPECT tracers. A further advantage of this method is the short acquisition time that enables to visualize perfusion-weighted measurements within a few minutes. However, the absolute quantification of CBF remains unsolved (Wintermark et al. 2005). The use of a local internal input function (Calamante et al. 2004) is necessary to receive parametric maps. As such, the detection of a plausible AIF which is influenced by numerous factors, such as partial volume effects, is important for reproducible and reliable perfusion values (van Osch et al. 2001). Another respective challenge is the localization of the intracranial region of interest for the AIF calculation (Zaro-Weber et al. 2012). Thus, most of the calculated parameter maps in perfusion MR are named as “relative” (Griffiths et al. 2001; Jezzard 1998). Nevertheless, perfusion MRI is employed for diagnostic purposes, for instance in acute stroke, and in clinical settings regarding the combination of the various readout parameters. In comparison to gadolinium-based MRI, the arterial spinlabeling approach is another promising MR perfusion technique. Here, magnetically labeled water protons are used as endogenous tracer. However, problems with image interpretation may occur due to a limited signal-to-noise ratio and in the presence of prolonged blood transit times, like in patients with stroke or atherosclerosis (Petersen et al. 2006). In latter cases (e.g., stroke patients), the labeled water spins did not reach the target brain tissue within a given time, with the consequence of underestimating the real blood flow values (Jezzard 1998; Kimura et al. 2005).

6 Applications for CBF PET

6.1 Cerebral Ischemia

The increasing incidence for vascular diseases, like atherosclerosis, is associated with a worldwide increasing number of ischemic attacks (Feigin et al. 2009). Interruptions in brain–blood supply rapidly leads to ischemic cell damage that results in necrotic tissue if no sufficient therapy or spontaneous reperfusion becomes available. Current therapies aim to restore perfusion in the ischemic, salvageable brain tissue. In order to determine this hypoperfused yet viable tissue – the so-called ischemic penumbra (Astrup et al. 1981) – in acute stroke patients, perfusion imaging techniques were developed to identify this “tissue at risk” and to separate it from the already necrotic infarction core. The transition from reversible to irreversible damage is a function of ischemia duration as well as of CBF. A hemodynamic determination of the ischemic penumbra is used for many research studies in laboratory animals and humans, and CBF thresholds to characterize different tissue states were proposed: While normal human CBF is in the range of 50–80 ml/100 g/min, reversible ischemia (“ischemic penumbra”) is evident when CBF drops below values around 22 ml/100 g/min, and neuronal cell death occurs below a CBF of 8 ml/100 g/m in (Baron 2001). However, penumbra detection, based on CBF thresholds, is highly dependent on a reliable and accurate quantitative imaging method in an acute stroke diagnostic setting.

The opportunity to investigate different parameters of brain function, like CBF, CBV, OEF, and CMRO2, within one PET imaging session is a further advantage in the investigation of cerebrovascular diseases such as ischemic stroke. Here, it is essential to reliably separate primary perfusion deficits from events of decreased metabolic demand. This is as CBF decreases may not only appear in the surrounding tissue of the stenotic vessel, but sometimes also distant from the obviously damaged part of the brain. This well-recognized phenomenon is called “diaschisis.” One example is the so-called crossed cerebellar diaschisis in which a CBF reduction in the cerebellum contralateral to the stroke-affected brain hemisphere occurs as a result of crossed functional deafferentiation (Baron et al. 1981; Feeney and Baron 1986).

6.1.1 PET Perfusion Imaging in Preclinical Stroke Research

This paragraph will handle the employment of [15O]H2O PET in translational research studies of experimental stroke.

In 2008, our group proposed a new large animal model that is applicable for acute and chronic stroke induction (Boltze et al. 2008) and is highly suitable to reflect the human brain pathophysiology. Due to a similar cerebral anatomy and the favorable ovine brain size, brain imaging protocols, scanners, and data analysis techniques as used in clinical routine become feasible. This enables us to perform studies while meeting main conditions for translational research. Like in humans, in sheep, the middle cerebral artery (MCA) usually gives rise to three arterial branches. Different stroke sizes can be induced by permanent transcranial occlusion of one, two, or all three MCA branches, with the latter being referred to as permanent MCA occlusion (pMCAO). Fig. 4.3 shows examples of different occlusion types for experimental pMCAO in sheep. One-, two-, and three-branch (total) occlusions of the MCA could clearly be visualized by magnetic resonance angiography (MRA), together with the resulting CBF defects in [15O]H2O PET. In addition, slight CBF decreases, probably due to the transcranial surgery, were also detected in the sham-operated animals. Further, it was possible to demonstrate that the ischemic strokes induced by the transcranial pMCAO lead to reproducible CBF deficits, which remain stable over time and eventually leading to necrotic brain tissue.

Fig. 4.3
figure 3

Differential CBF deficits dependent on extent of experimental permanent middle cerebral artery occlusion in sheep. The extent of the post-pMCAO CBF deficit is clearly visualized with [15O]H2O PET. Corresponding to angiographic MRI findings, the CBF deficit increases in the order sham >1-branch pMCAO > 2-branch pMCAO > total pMCAO. MCAO middle cerebral artery occlusion, MRA magnetic resonance angiography (Modified from Boltze et al. (2008))

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The major advantage of [15O]H2O PET is the option to perform serial scans in a short time period due to the short half-life of the 15O (122 s). As an example, CBF PET measurements were applied for a controlled preclinical study that aimed to test inhaled nitrous oxide (iNO) for its potential to protect the ischemic tissue in the penumbra in acute ischemic stroke. All animals were subjected to repeated PET scans at 110, 150, 175, and 210 min following pMCAO (a total of four PET scans within 100 min). In the treatment group, 50 ppm iNO were applied from 120 to 180 min after pMCAO. By using kinetic modeling, parametric CBF maps were created. Based on the abovementioned commonly accepted CBF thresholds (Baron 2001), operator-independent brain volumes of interest were defined for penumbra, infarction core, and remaining normal brain tissue (Fig. 4.4). Our experiments showed that iNO selectively restores CBF in the ischemic penumbra. While the volume of the necrotic core was not affected, the volume of the penumbra decreased by up to 50 % turning into normally perfused tissue (>22 ml/100 g/min) under iNO application, but remained unchanged in the untreated control animals (p < 0.05 vs. baseline and vs. control; Figs. 4.5 and 4.6; Terpolilli et al. 2012).

Fig. 4.4
figure 4

Parametric CBF map as obtained by [15O]H2O PET in the Leipzig permanent middle cerebral artery occlusion sheep stroke model. For volume of interest analysis, the stroke-related regions were defined as follows: infarction core (a): <8 ml/100 g/min, ischemic penumbra: (b) 8–22 ml/100 g/min, normal brain tissue (c): >22 ml/100 g/min

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Fig. 4.5
figure 5

Serial [15O]H2O PET scans show therapeutic effect of iNO in the Leipzig permanent middle cerebral artery occlusion sheep stroke model. Stroke-related tissue regions were defined on the basis of quantitative CBF maps. Volumetric analysis showed a significant decrease of the penumbra volume under iNO therapy in favor of the normal brain tissue compartment. This beneficial effect was not detected in control animals (Terpolilli et al. 2012)

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Fig. 4.6
figure 6

iNO improves penumbral blood flow after permanent middle cerebral artery occlusion in sheep. Quantification of normally perfused, ischemic, and penumbral tissue volumes revealed that penumbral volume decreased significantly during NO inhalation (n = 3 per group; *p < 0.05 vs. control and vs. baseline PET at t-10 min) in favor of the normally perfused brain tissue (Terpolilli et al. 2012)

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To give an outlook for further projects on CBF PET imaging in the Leipzig sheep pMCAO stroke model, Fig. 4.7 shows first images acquired by a simultaneous PET/3T-MRI system (Biograph mMR, Siemens). Comparative PET and MR imaging studies in acute stroke setting will greatly benefit from the new possibility to acquire data of both modalities simultaneously. This will significantly improve the investigation of the very fast pathophysiological processes in early ischemia. The first experience with this new simultaneous imaging approach, however, triggers great enthusiasm to employ this new technique for further preclinical and clinical research in the acute stroke situation.

Fig. 4.7
figure 7

Multimodal PET-MRI imaging in the Leipzig sheep stroke model. Four hours after permanent middle cerebral artery occlusion, (a) CBF was determined by [15O]H2O PET (shown here in overlay with anatomical T2-MRI) demonstrating the typical poststroke deficit. (b) T2-MRI acquired to exclude hemorrhage. (c) Diffusion-weighted MRI with deficit similar to the CBF abnormality. The image data were acquired simultaneously using a PET/3T-MR (Siemens mMR) system

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With regard to the abovementioned possibility of PET imaging to provide multiparameter readouts (CBF, OEF, CMRO2, CBV), in experimental ischemic stroke, so far mainly the brains of monkeys (Kuge et al. 2001; Pappata et al. 1993), pigs (Sakoh et al. 2000a) and felines (Heiss et al. 1994) were investigated by using [15O]H2O for CBF measurements. With examinations of different time points after stroke, the studies by Heiss et al. aimed at monitoring important parameters of the temporal transversion of penumbral tissue to the final infarct core. The ischemic penumbra is characterized by an initial increase of OEF and CBV and preserved values of CMRO2. In progress of the infarct and decreasing blood flow supply, CMRO2 declines to values of 25 % of baseline within the first hour after stroke, while the initially increase of OEF becomes less prominent. In the final stage, the lack of CBF is associated with a decrease in OEF and the occurrence of brain tissue necrosis. This progress from still viable to necrotic tissue was defined as a function of duration and severity of stroke (Heiss et al. 1994, 1997; Heiss and Rosner 1983). In further experiments from the Cologne group, postischemic hemodynamic and metabolic processes were investigated after temporary MCAO (from 30 to 120 min of duration) in anesthetized cats (Heiss et al. 1997). CBF PET was performed immediately after MCAO and was repeated at 30-min intervals. After reopening the vessel, a distinct hyperperfusion was found in all animals, depending on the duration of the MCAO. After 30 min of ischemia, the reactive CBF increase was found to be transient with a fast normalization to pre-occlusion CBF values and without major tissue necrosis. By comparison, the reperfusion period after 60 or 120 min of ischemia was associated with severe hyperperfusion (CBF increase up to 300 % compared to basal levels) and irreversible tissue damage, depending on the severity of prior ischemia. Approximately 50 % of animals died after the prolonged ischemic period and suffered from higher CBF than the surviving animals from the same group. One explanation might be the hyperperfusion processes with massive cerebral edema that resulted in malignant brain swelling. These observations let to assume a relatively high tolerance against moderate hyperemia, whereas extensive hyperperfusion leads to an increased mortality (Heiss et al. 1997).

Other large animal experiments utilized [15O]H2O (Sakoh et al. 2000a, b) in combination with [15O]CO PET to study regional correlations of CBF and CBV after acute stroke and in cases of reperfusion (Sakoh et al. 2000a). For that purpose, 13 pigs underwent PET, and the results were compared to MRI measurements of the same parameters. A good agreement was found between the values of both modalities in the ischemic tissue. After MCAO, a significant correlation between CBF reduction and CBV increase was observed. However, the decrease of CBF below 60 % of the contralateral side was found to induce a reduction in CBV. In contrast, both parameters were less correlated in cases of reperfusion (Sakoh et al. 2000a).

6.2 Activation Studies

Perfusion parameters are closely coupled with changes in neuronal activity (Roy and Sherrington 1890; Fox et al. 1988; Villringer and Dirnagl 1995). In the last decades, numerous neuronal activation studies have reported this effect, which, nevertheless, is not completely understood (Peterson et al. 2011). The possibility to perform sequential measurements within a certain time window made CBF PET an attractive tool to study perfusion during cognitive, motor, or sensomotor tasks in humans. During local brain activation, cerebral areas that are involved in task performance were identified by CBF changes (Feng et al. 2004; Worsley et al. 1992).

6.3 Other Applications for CBF PET

In general, PET is a useful tool to study perfusion and oxygen metabolism in any cerebral disease, such as dementia (Yao et al. 1990; Cohen et al. 1997), schizophrenia (Ragland et al. 2001), and migraine. Latter approaches addressed hemodynamic changes and oxygen metabolism during and between acute attacks. An increase in CBF was found in cortical areas and in the brainstem which appeared to be pain related (Andersson et al. 1997; Cutrer et al. 2000). Similar to these migraine studies, other preclinical and clinical investigations focus on epilepsy to determine perfusion changes during and between epileptic seizures (Szabo et al. 2007; Kahane 1999; Gaillard et al. 1995). A few studies also used CBF PET (mostly in combination with other PET readouts, like oxygen metabolism from 15O inhalation) for detailed characterization of brain tumors (Leenders 1994; Hino et al. 1990; Tomura et al. 1993). Due to the neo-formed vessels which often appear in the tumor parenchyma or its surrounding, a regional CBF increase could be detected and may provide additionally useful information about the progression of the brain tumors.

7 Summary and Conclusions

PET imaging with [15O]H2O represents the gold standard to visualize and quantify CBF in vivo. As such, this method plays a relevant role mainly for the investigation of cerebrovascular diseases. Despite advantageous characteristics of the tracer, like the possibility to monitor CBF changes over time in sequential scans, [15O]H2O PET imaging remains a technological and infrastructural challenge, preventing it from gaining a wider acceptance in a clinical routine setting. In contrast to clinical routine, [15O]H2O PET plays an important role in preclinical and clinical research, for instance, in (1) clarifying the pathophysiology of ischemic stroke and other cerebrovascular disorders, (2) cross-evaluation of alternative imaging methods to estimate CBF, and (3) the testing of new stroke treatment concepts.