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Translational neuroscience and magnetic-resonance microscopy
Review
Translational neuroscience and magnetic-resonance microscopy Helene Benveniste, Stephen J Blackband Lancet Neurol 2006; 5: 536–44 Medical Department, Brookhaven National Laboratory, Upton, NY, USA (H Benveniste MD); Department of Anesthesiology, Stony Brook University, Stony Brook, NY, USA (H Benveniste); Department of Neuroscience, University of Florida, Gainesville, FL, USA (S J Blackband PhD); and The National High Magnetic Field Laboratory, Tallahassee, FL, USA (S J Blackband) Correspondence to: Dr Helene Benveniste Medical Department, Brookhaven National Laboratory, Upton, NY 11973, USA [email protected]
Physical sciences
Magnetic-resonance microscopy is a rapidly growing and a widely applied imaging method in translational neuroscience studies. Emphasis has been placed on anatomical, functional, and metabolic studies of genetically engineered mouse models of human disease and the need for efficient phenotyping at all levels. Magnetic-resonance microscopy is now implemented in many laboratories worldwide due to the availability of commercial high-field magnetic-resonance instruments for use in small animals, the development of accessories (including miniature radio-frequency coils), magnetic-resonance compatible physiological monitoring equipment, and access to adjustable anaesthesia techniques. Two of the major magnetic-resonance microscopy applications in the neurosciences— structural and functional magnetic-resonance microscopy—will be reviewed.
Introduction MRI is probably one of the best examples of an imaging method that has developed into an indispensable technique in translational and preclinical studies of human neurological diseases. The term “translational” here refers to the use of small animal models for testing of the diagnostic and pathophysiological significance of magnetic-resonance parameters, such as T1, T2, T2*, and diffusion, perfusion, and magnetisation transfer. Over the past two decades anatomical, functional, and metabolic findings on MRI and spectra of the human brain have been translated into hypothesis-based and systematic MRI studies in animal models of neurological
Life sciences
MRI instrumention, magnetic resonance pulse sequences, radio frequency coils, bioinformatics, physiological monitoring, optical technology, Innovative animal models
New knowledge Hypothesis
Behaviour Disease Genetic profiling
Figure 1: Illustration of the translational research concept of MRI in the neurosciences Processes and features on MRI images in human beings have lead to hypothesis-driven research in small-animal models, which have been made possible by innovations and fusion of knowledge at the physical and life sciences interface. Knowledge and understanding of MRI contrast signatures of the normal and diseased brain from MRI research in small animals have generated new knowledge, which has been translated back into clinical practice as diagnostic and therapeutic tools for various diseases.
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diseases, and the findings applied as diagnostic and therapeutic monitoring methods in clinical practice (figure 1). For example, diagnosis of acute brain ischaemia1 and detection and assessment of stroke lesion maturation,2–4 post-ischaemic recirculation, and treatment efficacy were all derived largely from such translational MRI studies.5–8 From these and many other studies, it has become clear that magnetic-resonance parameters do not represent specific histopathology such as apotosis, degeneration, selective neuronal injury, fibrosis, and inflammation but are associated and correlated with these features. In other words, a given magneticresonance contrast signature for one disease can be the same as in another. With the completion of the human genome project in 2003, translational magnetic-resonance studies have developed even further with the availability of new genetic mouse models of human diseases and the need for diverse and time efficient morphological, physiological, and metabolic phenotyping. The MRI field now encompasses a wide range of disciplines and expertise; these include highfield magnetic-resonance instrumentation, multispectral imaging, miniature radio-frequency coil design operable in high-field systems, small-animal magnetic resonance that is compatible with physiological monitoring equipment, magnetic-resonance contrast agents for molecular imaging, mouse brain neuroinformatics, and magnetic-resonance phenotyping at all levels. There are already many strong reviews available on magneticresonance technology in general,9 magnetic-resonance microscopy and small animal imaging,10,11 and magneticresonance histology.12 In this review we focus on two major applications in the neurosciences; structural and functional magnetic-resonance microscopy with emphasis on translational applications.
Physical and technical concepts High resolution MRI and magnetic-resonance microscopy MRI was discovered more than 30 years ago and it continues to develop at a rapid pace. The primary drive behind this rapid development is the desire to improve http://neurology.thelancet.com Vol 5 June 2006
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the signal to noise ratio and image contrast, which together define the sensitivity of the magnetic-resonance technique and detect differences between normal and diseased tissues, changes in those differences (by natural progression or after intervention), or changes caused by exogenous drugs. Foremost among the technological developments has been the development of ever-stronger magnetic fields to improve signal-to-noise ratio and spectral resolution. Furthermore, the scientific drive to rapidly translate magnetic-resonance-microscopy data to use in the clinic has led to the development of instruments with user interfaces and pulse-sequence encoding that match clinical systems (eg, the Bruker ClinScan 7-T small animal MRI system that combines the clinical user interface of Siemens’ MRI technology). These developments enabled pulse sequence development and testing in a clinical setting. The physical size of the sample being tested determines the available magnetic-field strength because the field achievable is proportional to the diameter of the main solenoidal coil in superconducting magnets. Wide-bore magnets (89 mm vertical bore) at 600–800 MHz are becoming widely available; the National High Magnetic Field Laboratory recently commissioned the world’s first 900 MHz (11 cm bore) magnet.13 Small-animal systems (mainly 20–40 cm horizontal bore magnets) at 7–11·7 T are becoming more commonly available. 3 T human systems (80–100 cm bore) are now commonplace (and the clinical standard), 7–8 T systems are being assessed, and a recent study used a 9·4 T human system;14 an 11·7 T human system has also been proposed.15 However, after insertion of ancillary technology that is needed for imaging (gradients, shims, radio-frequency coils) only 40–60% of the magnet bore is available for the sample. Thus the higher field systems accommodate mice and isolated tissues, the animal systems small mammals up to small primates, and the human systems can obviously be used for human beings but also for larger animals. The most important determinant of the signal-to-noise ratio is the size of the radio-frequency coil (the ratio is linearly dependent on the radio-frequency coil diameter), which is used to detect the signal. Therefore, the size of the coil should be matched to the size of the object (or region on a larger object) for optimum signal to noise ratio. The smallest coils are only tens or hundreds of micrometres in diameter and have been used to image large, isolated, single cells.16 Array coils have been developed that can further improve signal-to-noise ratio, imaging speed, and image coverage.17 These array coils work by recognising that smaller coils give better signalto-noise ratio but a small field of view, and so compensate for the small field of view by using multiple small coils simultaneously. As a consequence, an array of independent receivers is needed. These coils are now standard on clinical machines and are being developed for high-field animal and vertical bore systems. http://neurology.thelancet.com Vol 5 June 2006
Given this basic technology, the image quality achievable in MRI depends on a complex interplay of the imaging technique (pulse sequence) used to achieve the desired type of contrast (T1, T2, diffusion, or magnetisation transfer) and the imaging speed needed (especially for functional imaging). Once these factors are determined, the signal-to-noise ratio will dictate the spatial resolution that can be attained. Other factors also affect image quality and signal-to-noise ratio, including, but not limited to, susceptibility effects, motion, in vivo sample viability. and temporal stability. If resolution is less than 100 μm in at least one spatial dimension, then this is typically referred to as “magnetic-resonance microscopy”.16 If resolution is above 100 μm (but below standard clinical imaging), imaging is usually referred to as “high resolution MRI”. These terms are commonly used interchangeably, but we stress that it is important not to be fixated on resolution. Imaging speed and contrast are equally important and, together with resolution, dictate what can actually be seen, thus determining the sensitivity and specificity of the technique. Generally, a resolution of tens of micrometres can be achieved on small samples (a couple of centimetres or less) in minutes to hours in a high magnetic-field strength. On larger samples at lower magnetic fields or over shorter times, especially in-vivo or functional studies, the resolution relaxes to about 50–200 μm, although this can depend on what the sample is and exactly what needs to be detected. A better idea of what can be achieved is provided by the many examples in the rest of this review.18,19
Contrast generation via exogenous agents Classic contrast agents influence the relaxation times of water in their immediate vicinity and include gadolinium derivatives (causing signal hyperintensity) and super paramagnetic oxide particles (causing signal hypointensity). These agents have great clinical utility, because they can leak out into and highlight damaged tissue, however, they are restricted to extracellular spaces. “Smart” or targeted contrast agents have been developed that attach to specific tissues (eg, tumour cells) or become activated only in the presence of another agent. Some of these agents have the potential to work their way into cells for monitoring gene expression,20 and agents based on nanoparticles and associated technology are being rapidly developed.21 Although these new agents have promising potential, the difficulties involved in their application cannot be underestimated. For example, the concentration of these agents is so small that ways to amplify their potential signal must be developed if they are to be clinically useful.22 In addition, when signal changes are potentially small, the image spatial resolution and acquisition time might be compromised. There is also great concern about the potential toxicity of these agents and the 537
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body’s ability to remove them: it may be necessary to develop additional agents to actively remove the contrast agent after they have been given. Because these methods are under development, they will not feature extensively in this review, which instead will concentrate on the better-established morphological and functional magnetic-resonance techniques. They remain however, an exciting area with great potential, particularly for studying the brain.23
In-vivo magnetic-resonance microscopy Awake animal studies Most animal studies with magnetic-resonance microscopy require that the animal is anaesthetised for ethical reasons and to prevent motion artifacts, stress, discomfort, and pain if surgical procedures are involved. A few magnetic-resonance microscopy studies that examined evoked cortical responses and absence seizures using blood oxygen level dependent (BOLD) contrast have been done with awake rats.24,25 However, functional MRI activation patterns in the brains of awake animals must be interpreted with careful parallel assessment of potential concomitant stress in the animal (eg, by noting stress markers such as increased heart and respiratory rates as well as neuroendocrine responses). In addition, in studies in awake animals in which functional MRI responses to drugs like cocaine and narcotics are investigated, extra attention must be paid to physiological parameters during data acquisitions because of the effects these drugs have on respiration and heart rate, which can confound the BOLD signal interpretation.26–28
Anaesthesia An ideal anaesthetic would keep the animal asleep, free from pain and stress, immobilised, haemodynamically stable, and not interfere with breathing, heart rate, and the physiological process to be studied. However, on the basis of these criteria there is no perfect anaesthetic and so when choosing one for a given magnetic-resonance microimaging study it is important to consider several factors. These factors include length of study, recovery time, intravenous access, spontaneous or mechanical ventilation, disease status of the animal (important for certain transgenic mouse strains), type of experiment (eg, anatomical, spectroscopic, or functional data acquisitions), and interference with other drugs during the study. For functional MRI studies, chloral hydrate, alpha chloralose, or propofol are commonly used because they interfere less with sensory evoked potentials, including brainstem auditory potentials.29,30 Alpha chloralose has been successfully used in functional imaging studies in rats characterising brain activation patterns during forepaw and hindpaw stimulation,31–33 visual stimulation,34 whisker stimulation,35 noxious stimulation,36 with cocaine administration,28,37 and with other drugs.27 538
Physiological monitoring equipment In the earliest MRI studies in rodents, no physiological parameters were monitored and during imaging the animal was wrapped loosely in plastic sheeting for immobilisation (and probably to control temperature) and placed supine in the imaging field.38 Over the past two decades, increasingly sophisticated anaesthesia delivery systems, physiological monitoring, and supportive equipment for small animals have been developed.39–44 Physiological monitors suitable for rats and smaller rodents are available to monitor respiratory rate, tidal volume, ventilation pressure (if mechanically ventilated), body temperature, heart rate, and electrocardiography. Typically the monitors are operated by fibre optics and therefore are magnetic-resonance compatible (eg Module 224002, Small Animal Instrument Inc Stony Brook, NY, USA). Importantly, the fibre-optic monitors function in magnetic-resonance systems, which operate at field strengths in excess of 11 T and can track heart rates up to 700 beats per minute. These new monitoring devices are very useful in magnetic-resonance microscopy studies on transgenic mouse models of neurological diseases such as multiple sclerosis, amyotrophic lateral sclerosis, or spinal-cord defects or injuries where the animals have increased sensitivity to anaesthetics and are therefore at risk of respiratory or cardiovascular compromise.
Morphology of mouse models of neurological diseases The ability to resolve rodent brain anatomy on MRI is essential not only for the assessment of macroscopic abnormalities, such as atrophy, hydrocephalus, demyelination, plaque deposition, or periventricular leucomalacia, but also to accurately quantify more subtle morphological changes that might occur in animals treated with drugs, disease models, or animals exposed to environmental and behavioural stressors. For genetically engineered mice, it is important to understand how the gene in question affects the morphological phenotype and whether the engineered genotype produces the desired phenotype—ie, does it actually mimic clinical disease? For example, the desired pathology for a mouse model of multiple sclerosis would be demyelination, demyelinated plaques, and reactive gliosis whereas for attention hyperactivity disorder it might be cerebellar volume asymmetry.45 Quantitative analysis of rat and mouse brain structures can be obtained from in vivo as well as in-vitro magnetic-resonance microscopy (table).46–56 A superoxide dismutase (SOD)1G93AG1H transgenic mouse—an animal model of familial human amyotrophic lateral sclerosis—was recently characterised by magneticresonance microscopy.57 These mice develop tremor, paresis, and paralysis of the hind limbs due to loss of motor neurons and degeneration of the CNS. On T2weighted magnetic-resonance microscopy, a 3 month SOD transgenic mouse had areas of hyperintensity in the http://neurology.thelancet.com Vol 5 June 2006
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Familial human amyotrophic lateral sclerosis
Strain/mouse model
MRI findings
Superoxide dismutase SOD1G93A G1H
T2-weighted MRI: areas of hyperintensity (neurodegeneration) in nucleus ambiguous, facial nucleus, and motor trigeminal nucleus46
Multiple sclerosis
Cuprizone model
T2-weighted MRI: enlargement of ventricles; hyperintesnsity (demyelination)47
Alzheimer’s disease
APP23 mice
Quantitative diffusion imaging (ADC): local ADC decreases in APP23 mice48
Perinatal brain injury
Human S100B transgenic (Tg) and S100B knock-out (KO) mice
T2-weighted MRI: hemispheric volume loss after ischaemia49,50
Neural stem cell injected into mouse brain
C3H mice
T2*-weighted magnetic resonance images: hypointensity associated with stem cells (previously incubated with SPIO)51
Duchenne muscular dystrophy
Mdx mouse
Quantitative T2: no change in frontal cortex52
Knobloch syndrome
Col18a1(-/-) mice
T2-weighted MRI: hydrocephalus53
Leber hereditary optic neuropathy, Leigh syndrome
Intraocular injection of recombinant rAAV in DBA/1J mice
T1-weighted MRI+contrast: thinning of optic nerve, not rescued with adenoassociated virus to deliver the human gene for SOD254
Cerebral ischaemia
C57BL/6J mice ApoE-deficient
Diffusion-weighted MRI: brain atrophy; enlargement of ventricles55
Head injury
C57BL/6J mice ApoE-deficient
Diffusion-weighted MRI: unilateral hemispheric swelling, obliteration of lateral ventricle56
Table: MRI findings in mouse models of human disoders
brainstem (nucleus ambiguous, rostroventrolateral reticular nucleus, and the lateral paragigantocellular nucleus) and trigeminal motor nuclei, which correlated with areas of vacuolar degeneration on histological sections.57 As expected, degeneration of the nucleus ambiguous and trigeminal nucleus was associated with deficits in the movement of larynx, pharynx, and reduced mastication similar to that seen in human beings with bulbar onset amyotrophic lateral sclerosis.57 These mice develop respiratory deficiencies and it is methodologically demanding to follow the long-term sequelae by magneticresonance microscopy; only one timepoint (3 months) has been investigated so far. The various transgenic mouse models of Alzheimer’s disease are indicative of the search for defining specific genes that affect the pathology of the disease and allow the assesment of potential therapies.57 An ideal mouse model for Alzheimer’s disease would display behavioural equivalents of cognitive decline (memory deficits) and hallmark pathological features such as deposition of amyloid plaques, neurofibrillary tangles, reactive gliosis, dystrophic neuritis, neurodegeneration, and brain atrophy. Amyloid-plaque detection has been investigated with radiolabelled positron emitters, PET, and single-photonemission CT.58–63 However, MRI detection of plaques in vivo was recently achieved with a fluorine-19-labelled compound that specifically labels amyloid-β plaques in the brain.64 FSB—(E,E)-1-fluoro-2,5-bis(3-hydroxycarbonyo4-hydroxy)styrylbenzene—an amyloidophilic Congo-red type compound that crosses the blood–brain barrier, was intravenously injected into APP transgenic mice and subsequently imaged by T1-weighted magnetic-resonance contrast;64 the amyloid deposits were clearly visualised.64 Brain atrophy can also be detected by magneticresonance microscopy in vivo55 and in vitro;65 because atrophy in the temporal poles and entorhinal cortex are one of the earliest clinical signs of early cognitive decline during Alzheimer’s disease in human beings, magnetichttp://neurology.thelancet.com Vol 5 June 2006
resonance microscopy can be used to assess the clinical utility of transgenic mouse models of Alzheimer’s disease.66 Magnetic-resonance microscopy has been used to assess the development of brain atrophy in formalinfixed brains of transgenic mouse models that overexpresses human mutant amyloid precursor protein (V717F) under control of the platelet derived growth factor promoter (PDAPP mice) at 40, 100, 365, and 630 days.65 In agreement with clinical data, the investigators reported a 12% reduction of hippocampal volume in the transgenic mice at 100 and 630 days compared with the wild-type control group.65 In addition, the corpus callosum was substantially longer in the control mice compared with the transgenic mice.65 In-vivo preliminary magneticresonance microscopy studies of PDAPP mice have not
Figure 2: MRI of a homozygous PDAPP mouse before and after formalin-fixation Left: diffusion-weighted MRI microscopy of a homozygous PDAPP mouse acquired with 7·1 T intrument at a spatial resolution of 0·0012 mm³ at optimum contrast-to-noise for visualisation of mouse anatomy in vivo. Right: same mouse brain after formalin-fixation. The formalin fixed brain was excised from the skull and imaged with a T2*-weighted contrast sequence at a spatial resolution of 0·00012 mm³ with a 9·4 T solenoid coil (CIVM, Duke University). Images generated by H Benveniste on magnetic-resonance systems at the Center for In Vivo Microscopy, Duke University.
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Figure 3: Diffusion-weighted images of a homozygous PDAPP and wild-type control Top: diffusion-weighted MRI of a homozygous PDAPP mouse at the level of the dorsal hippocampus (left) and ventral hippocampus (right). Bottom: diffusion-weighted MRI of a wild-type control mouse at the level of dorsal hippocampus (left) and ventral hippocampus (right); arrows show the lateral ventricles, which are clearly larger in the wildtype control compared with the PDAPP mouse.
confirmed brain atrophy at age 22 months67 probably because image contrast differs between in-vivo and invitro specimens (figure 2). A striking abnormality in PDAPP mice is a smaller ventricle volume compared with their control litter mates, the reasons for which are unknown (figure 3). The discrepancy between in vitro and in vivo findings is clearly associated with the difficulty in defining anatomical areas of interest on MRI in vivo because of scan time restriction and motion, which decreases overall contrast-to-noise ratio (borders between anatomical areas are obscured), signal-to-noise ratio, and therefore spatial resolution of the image. Additionally, the formalin fixation process in vitro chemically changes tissue characteristics so that the image contrast-to-noise ratio is improved compared with that obtained in vivo. Figure 4 shows that anatomical delineation is better in the T2*weighted MRI acquired from a formalin-fixed specimen 540
on a 17·6 T instrument at a spatial resolution of 0·001 mm³ (scan time 5·7 h) compared to a T2-weighted and diffusion-weighted MRI acquired in vivo on a 9·4 T instrument at a spatial resolution of 0·001 mm³ and 0·0047 mm³, respectively. The difference in segmentation accuracy on in-vitro magnetic-resonance microscopy images versus in-vivo is indicated in the coefficient of error (measure of how much spread there is in the set of values relative to the mean) of the volume measurements. For example, hippocampal volume assessments (in vitro) in C57BL6/J mice had a mean of 25·7 mm³ (SD 1·1),68 18·61 mm³ (1·0),68 and 21 mm³ (0·5),69 equal to a coefficient of error of 0·04 (1·1/25·7), 0·05, and 0·02, respectively compared with in-vivo hippocampus volumes of 12·1 mm³ (1·4) and (dorsal hippocampus only) 8·9 mm³ (0·9)55 equal to coefficient of error of 0·12 and 0·10, respectively. The large coefficients of error in vivo are likely related to difficulties with segmentation and affect statistical power if the sample size is not adjusted. For example, for a given set of experiments (independent samples) with a coefficient of error of 0·04, at the significance level of 0·05 (two-sided), five in each group would yield a power of 93% to detect a 10% change with the independent samples t test; however, for the set of experiments (independent samples) with a coefficient of error of 0·1, 20 in each group would give a power of 87% with the independent samples t test. However, the greater segmentation accuracy in vitro is offset by greater potential for errors from tissue shrinkage and other artifacts of fixation. These problems in quantification of volumetric in vivo changes can be overcome by overlaying, aligning, and warping better segmented brain atlas templates onto the in-vivo images.70 This approach is routinely used in human studies and is increasingly being applied in rodent studies using MRI and other imaging modalities as more digital rodent atlases become available. A recent study showed the advantages of using a contrast agent for in vivo imaging of mice. In this study mice were injected intraperitoneally 48 h before imaging with 20 mM/kg of MnCl2 which accumulates in brain cells and highlights contrast to noise ratio in cell-rich areas.71 Bock and colleagues71 were able to acquire in vivo MRIs with an isotropic resolution of 156 μm³ in less than 3 h. However, the actual anatomical segmentation advantage of this method still needs to be assessed.
Functional MRI in rodents Measurement of haemodynamic changes associated with brain activity using BOLD functional MRI is used extensively for cognitive and behavioural studies in normal and diseased human beings. Functional MRI studies in rodents have progressed slowly and are especially challenging because they are usually done in anaesthetised animals in which functional MRI signals are dampened or changed. Further, as the animal is not awake it cannot be imaged when performing tasks. http://neurology.thelancet.com Vol 5 June 2006
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Figure 4: Anatomical resolution in C57BL6/J mouse brain images acquired in vivo and in vitro with different magnetic resonance contrast techniques Left: in vivo T2 weighted RARE (rapid acquisition with relaxation enhancement) sequence, 100 μm×100 μm×100 μm, 9·4 T. Middle: in vivo diffusion-weighted spin-echo image, 0·009 mm3, 9·4 T. Right: in vitro T2*-weighted image, TR (repetition time)=150 ms, TE (echo time)=7·5 ms, 0·001 mm3, 17·6 T.
Therefore, functional MRI experiments of rodents need implementation of an externally applied stimulus, such as electrical stimulation, odour, visual stimulation, air puffs, chemical, drugs, or noxious stimulation. One of the earliest studies by Hyder and colleagues33 reported functional MRI activation patterns in the brain of an alpha-chloralose anaesthetised rat during forepaw stimulation. The forepaw stimulation paradigm has been used by several other investigators and is favoured over hind-paw stimulation because of susceptibility artifacts associated with skull structures (tympanic bulla and zygomatic arch) in the somatory sensory hind-paw area. Gyngell and colleagues31 examined variations of the functional MRI signal in response to the frequency of somatosensory stimulation and found that the activation intensity decreased for forepaw stimulation frequencies above 3 Hz, and was negligible at 9 Hz. High-resolution BOLD-based functional MRI images of this same paradigm revealed that the functional MRI response was maximum at the cortical surface and in cortical layer 4.32 Improvement of the temporal resolution of the functional MRI experiments from 18 s33 to 50 ms32 showed that the forepaw activated functional MRI signal starts deep within the layers of the cortex and then expands to outer parts of the somatosensory cortex.23,32 The usefulness of functional MRI studies in anaesthetised rodents may seem questionable, but important information can still be obtained. For example, assessment of integrity and alteration (plasticity) of neuronal pathways after a stroke can be studied and assessed using functional MRI in the anaesthetised rodent. Dijkhuizen and colleagues72 studied the correlation between infarct size and brain function recovery with functional MRI. Rats middle cerebral arteries were occluded for 2 h before anatomical and functional MRI after forelimb stimulation were done. http://neurology.thelancet.com Vol 5 June 2006
The total area of significant forelimb-stimulation-induced activation was calculated in each hemisphere and the laterality index was calculated, this index shows activity in the contralateral hemisphere relative to that in the ipsilateral hemisphere.72 The study showed that acute stimulation of the impaired forelimb (days 1 and 3) was accompanied by acute bilateral hemispheric activation (days 1 and 3); these findings are similar to those in patients with stroke and indicate ongoing neuroplasticity in the recovering brain. Furthermore, the laterality index was inversely related to lesion volume.72 Most functional MRI studies use BOLD contrast. However, an alternativee approach is to preload the brain with paramagnetic manganese (Mn-enhanced MRI) for 12–24 h, then expose the animal to a given stimulus when it is awake and doing tasks, and subsequently scan the anaesthetised animal using T1-contrast.73 Manganese enhances contrast in the rodent brain in several brain regions (olfactory bulb, ventricles, corpus callosum, CA3, dentate granule cells, amydala, and cerebellum) in a time-dependent manner.74 Manganese ions enter synaptically activated neurons through voltage-gated calcium channels23,74,75 and if electrically or pharmacologically stimulated, a signal enhancement will be visible on T1-weighted images. This approach is very similar to brain activation studies done using radiolabelled fluoro-deoxy-glucose and PET in which the stimulus is given at the time of injection of the radioactive tracer when the animal is awake and cognitively challenged; 30–40 min later when the tracer has been trapped intracellularly the PET scan is done during which the animal can either be awake or anaesthetised. Fluorodeoxyglucose PET or Mn-MRI approaches necessitate that the stimulus is constant, prolonged, or repeated during the uptake periods, and the brain activation pattern reported is therefore an integration of the activity seen in this interval. 541
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Yu and colleagues73 gave MnCl2 as an intraperitoneal injection to mice at a dose of 0·4 mmol/kg and after 24 h the mice were exposed to a sound-evoked auditory stimulus delivered at different frequencies when the animals were awake and freely moving in a specially-designed cage.73 A careful analysis of signal enhancement in the inferior colliculus revealed frequency-dependent tonotropic organisation. This is a unique new approach, which might be superior to BOLD-based functional MRI experiments in mouse brains where physiological stability, motion control, and functional MRI signals can be difficult to obtain in the anaesthetised state because of the small size. Functional MRI studies in transgenic mouse models using external stimuli or drugs are not yet routine given the difficulties involved in doing such studies. However, with the advancement in radio-frequency coil technology, positioning devices, pulse sequences, and hardware (shim tools) this technique will become more widespread. There are many transgenic mouse models, in which functional MRI approaches would provide insight into pathophysiology and disease processes. For example, dopamine transporter deficient mice and mice transgenic for human dopamine receptors DA 2 and DA 4 have been used in the study of several diseases such as addiction, attention hyperactivity disorder, and Parkinson’s disease, and we have not yet phenotypically explored functional neurocircuitry in these models.
Conclusion We hope that this review has highlighted the importance of magnetic-resonance microscopy in translational neuroimaging studies. Growing research is increasingly focused on improving specificity of magnetic-resonance contrast and attempts to extract multicompartmental information from magnetic-resonance parameters. This will, to a large extent, determine the future of the clinical application of magnetic-resonance microscopy studies. The latter relates to the longterm goal in magneticresonance studies to use quantitative measurements to improve the sensitivity and specificity of MRI, especially with regard to T1 and T2 measurements. These goals have proven difficult to achieve, primarily because of inter- and intra-sample biological heterogeneity and variability in imaging techniques. Additionally, the data acquisition times are such that there is considerable exchange between tissue compartments at the cellular level, making data interpretation difficult. More recently, the search for quantitative techniques has been ressurected by the observation that diffusion measures can reveal multicompartmental information and can be collected with short echo times minimising compartmental exchange. Further, these compartmental signals change with tissue perturbations in a way that is consistent with cell swelling. Model systems are under detailed experimental and mathematical modelling investigations and it remains to be seen if these measures will have a clinical significance. 542
Search stategy and selection crtieria References for this review were identified by searches of MEDLINE and MEDLINEplus between 1969 and January 2006 and from relevant articles; numerous articles were identified through searhces of the extensive files of the authors. The search terms “MR microscopy”, “MR imaging”, “rodent”, “functional MRI”, “fMRI”, “neuroinformatics”, “morphology”, “brain”, “brain atlas” were used. Abstracts and reports from meetings were also included. Only papers published in English were reviewed. The final reference list was generated based on originality and relevance to the topics covered in this review.
Contributors Both authors contributed equally to the article. Conflicts of interest We have no conflicts of interest. Acknowledgments The authors would like to acknowledge the financial support of the NIH through grants R01 EB 00233-04 and P41 RR16105 and the National High Magnetic Field Laboratory. MRI data were obtained at the Advanced Magnetic Resonance Imaging and Spectroscopy facility in the McKnight Brain Institute of the University of Florida. In addition, we would like to thank the SBU/BNL 9.4T magnetic-resonance microscopy facility at Brookhaven National Laboratory. References 1 Kucharczyk J, Mintorovitch J, Asgari HS, Moseley M. Diffusion/ perfusion MR imaging of acute cerebral ischemia. Magn Reson Med 1991; 19: 311–15. 2 Gill R, Sibson NR, Hatfield RH, et al. A comparison of the early development of ischaemic damage following permanent middle cerebral artery occlusion in rats as assessed using magnetic resonance imaging and histology. J Cereb Blood Flow Metab 1995; 15: 1–11. 3 Minematsu K, Li L, Sotak CH, Davis MA, Fisher M. Reversible focal ischemic injury demonstrated by diffusion-weighted magnetic resonance imaging in rats. Stroke 1992; 23: 1304–10. 4 Verheul HB, Berkelbach van der Sprenkel JW, Tulleken CA, Tamminga KS, Nicolay K. Temporal evolution of focal cerebral ischemia in the rat assessed by T2-weighted and diffusion-weighted magnetic resonance imaging. Brain Topogr 1992; 5: 171–76. 5 Lo EH, Matsumoto K, Pierce AR, Garrido L, Luttinger D. Pharmacologic reversal of acute changes in diffusion-weighted magnetic resonance imaging in focal cerebral ischemia. J Cereb Blood Flow Meatb 1994; 14: 597–603. 6 Reith W, Hasegawa Y, Latour LL, et al. Multislice diffusion mapping for 3-D evolution of cerebral ischemia in a rat stroke model. Neurology 1995; 45: 172–77. 7 Roussel SA, van Bruggen N, King MD, et al. Monitoring the initial expansion of focal ischaemic changes by diffusion-weighted MRI using a remote controlled method of occlusion. NMR Biomed 1994; 7: 21–28. 8 Steinberg GK, Kunis D, DeLaPaz R, Poljak A. Neuroprotection following focal cerebral ischaemia with the NMDA antagonist dextromethorphan, has a favourable dose response profile. Neurol Res 1993; 15: 174–80. 9 Tyszka JM, Fraser SE, Jacobs RE. Magnetic resonance microscopy: recent advances and applications. Curr Opin Biotechnol 2005; 16: 93–99. 10 Pautler RG. Mouse MRI: concepts and applications in physiology. Physiology (Bethesda) 2004; 19: 168–75. 11 Pirko I, Fricke ST, Johnson AJ, Rodriguez M, Macura SI. Magnetic resonance imaging, microscopy, and spectroscopy of the central nervous system in experimental animals. NeuroRx 2005; 2: 250–64. 12 Henkelman RM, Chen XJ, Sled JG. Disease phenotyping: structural and functional readouts. Prog Drug Res 2005; 62: 151–84.
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Fu R, Brey WW, Shetty K, et al. Ultra-wide bore 900 MHz highresolution NMR at the national high magnetic field laboratory. J Magn Reson; 2005; 177: 1–8. Thulborn K. 9.4T. in International Society of Magnetic Resoance in Medicine 2005. Feder T. Issues and events: lab weds brain research and physics. Phys Today April 2005; 26. Aiken NR, Hsu EW, Blackband SJ. A review of NMR microimaging studies of single cells. J Magn Reson Anal 1995; 1: 41–48. Poemer PB, Edelstein WA, Hayes CE, Souza SP, Mueller OM. The NMR phased array. Magn Reson Med 1990; 16: 192–225. Benveniste H, Blackband SJ. MR microscopy and high resolution small animal MRI: applications in neuroscience research. Prog Neurobiol 2002; 67: 393–420. Benveniste H, Mareci T, Blackband SJ, Anatomical studies in the rodent brain and spinal cord: Applications of magnetic resonance microscopy. In van Bruggen N, Roberts T eds. Techniques of biomedical imaging in neuroscience. CRC Press, 2003. Bell JD, Taylor-Robinson SD. Assesing gene expression in vivo: magnetic resonance imaging and spectroscopy. Gene Ther 2000; 7: 1259–64. Lanza GM, Winther PM, Caruthers SD, et al. Magnetic resonance molecular imaging with nanoparticles. J Nucl Cardiol 2004; 11: 733–43. Delikatny EJ, Poptani H. MR techniques for in vivo molecular and cellular imaging. Radiol Clin North Am 2005; 43: 205–20. Koretsky AP. New developments in magnetic resonance imaging of the brain. NeuroRx 2004; 1: 155–64. Lahti KM, Ferris CF, Li F, Sotak CH, King JA. Comparison of evoked cortical activity in conscious and propofol-anesthetized rats using functional MRI. Magn Reson Med 1999; 41: 412–16. Sicard K, Shen Q, Brevard ME, et al. Regional cerebral blood flow and BOLD responses in conscious and anesthetized rats under basal and hypercapnic conditions: implications for functional MRI studies. J Cereb Blood Flow Metab 2003; 23: 472–81. Baslow MH, Dyakin VV, Nowak KL, Hungund BL, Guilfoyle DN. 2-PMPA, a NAAG peptidase inhibitor, attenuates magnetic resonance BOLD signals in brain of anesthetized mice: evidence of a link between neuron NAAG release and hyperemia. J Mol Neurosci 2005; 26: 1–16. Gozzi A, Schwarz AJ, Reese T, et al. Functional magnetic resonance mapping of intracerebroventricular infusion of a neuroactive peptide in the anaesthetised rat. J Neurosci Methods 2005; 142: 115–24. Marota JJ, Mandeville JB, Weisskoff RM, et al. Cocaine activation discriminates dopaminergic projections by temporal response: an fMRI study in rat. Neuroimage 2000; 11: 13–23. Shaw NA. The effects of chloralose on the brainstem auditory evoked potential of the rat. Neuropharmacology 1990; 29: 561–65. Helfaer MA, Kirsch JR, Traystman RJ. Anesthetic modulation of cerebral hemodynamic and evoked responses to transient middle cerebral artery occlusion in cats. Stroke 1990; 21: 795–800. Gyngell ML, Bock C, Schimtz B, Hoehn-Berlage M, Hossmann KA. Variation of functional MRI signal in response to frequency of somatosensory stimulation in alpha chloralose anesthetized ratas. Magn Reson Med 1996; 36: 13–15. Silva AC, Koretsky AP. Laminar specificity of fMRI onset times during somatosensory stimulation in the rat. Proc Natl Acad Sci USA 2002; 99: 15182–87. Hyder F, Bhehar KL, Martin MA, Blamire AM, Shulman RG. Dynamic magnetic resonance imaging of the rat brain during forepaw stimulation. J Cereb Blood Flow Metab 1994; 14: 649–55. Huang W, Plyka I, Li H, Eisenstein EM, et al. Magnetic resonance imaging (MRI) detection of the murine brain response to light: temporal differentiation and negative functional MRI changes. Proc Natl Acad Sci USA 1996; 93: 6037–42. Devor A, Ulbert I, Dunn AK, et al. Coupling of the cortical hemodynamic response to cortical and thalamic neuronal activity. Proc Natl Acad Sci USA, 2005; 102: 3822–27. Tuor UI, Malisza K, Foniok T, et al. Functional magnetic resonance imaging in rats subjected to intense electrical and noxious chemical stimulation of the forepaw. Pain 2000; 87: 315–24. Febo M, Segarra AC, Nair G, et al. The neural consequences of repeated cocaine exposure revealed by functional MRI in awake rats. Neuropsychopharmacology 2005; 30: 936–43.
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Zhang W, Oya S, Kung MP, et al. F-18 stilbenes as PET imaging agents for detecting beta-amyloid plaques in the brain. J Med Chem 2005; 48: 5980–88. Higuchi M, Iwata N, Matsuba Y, et al. 19F and 1H MRI detection of amyloid beta plaques in vivo. Nat Neurosci 2005; 8: 527–33. Redwine JM, Kosofsky B, Jacobs RE, et al. Dentate gyrus volume is reduced before onset of plaque formation in PDAPP mice: a magnetic resonance microscopy and stereologic analysis. Proc Natl Acad Sci USA 2003; 100: 1381–86. Albert MS. Cognitive and neurobiologic markers of early Alzheimer disease. Proc Natl Acad Sci USA 1996; 93: 13547–51. Zhang L, Kim K-R, Bales K, Paul S-M, Benveniste H. Tracking of pathology in PDAPP mice by MRI: a two-year longitudinal study. Abstr Soc Neurosci Program No. 572.14, 2001; 27. Ma Y, Hof PR, Grant SC, et al. A Three-dimensional digital atlas database of the adult C57B/6J Mouse brain by magnetic resonance microscopy. Neuroscience 2005; 135: 1203–15. Kovacevic N, Henderson JT, Chan E, et al. A three-dimensional MRI atlas of the mouse brain with estimates of the average and variability. Cereb Cortex 2005; 15: 639–45.
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Translational neuroscience and magnetic-resonance microscopy Helene Benveniste, Stephen J Blackband Lancet Neurol 2006; 5: 536–44 Medical Department, Brookhaven National Laboratory, Upton, NY, USA (H Benveniste MD); Department of Anesthesiology, Stony Brook University, Stony Brook, NY, USA (H Benveniste); Department of Neuroscience, University of Florida, Gainesville, FL, USA (S J Blackband PhD); and The National High Magnetic Field Laboratory, Tallahassee, FL, USA (S J Blackband) Correspondence to: Dr Helene Benveniste Medical Department, Brookhaven National Laboratory, Upton, NY 11973, USA [email protected]
Physical sciences
Magnetic-resonance microscopy is a rapidly growing and a widely applied imaging method in translational neuroscience studies. Emphasis has been placed on anatomical, functional, and metabolic studies of genetically engineered mouse models of human disease and the need for efficient phenotyping at all levels. Magnetic-resonance microscopy is now implemented in many laboratories worldwide due to the availability of commercial high-field magnetic-resonance instruments for use in small animals, the development of accessories (including miniature radio-frequency coils), magnetic-resonance compatible physiological monitoring equipment, and access to adjustable anaesthesia techniques. Two of the major magnetic-resonance microscopy applications in the neurosciences— structural and functional magnetic-resonance microscopy—will be reviewed.
Introduction MRI is probably one of the best examples of an imaging method that has developed into an indispensable technique in translational and preclinical studies of human neurological diseases. The term “translational” here refers to the use of small animal models for testing of the diagnostic and pathophysiological significance of magnetic-resonance parameters, such as T1, T2, T2*, and diffusion, perfusion, and magnetisation transfer. Over the past two decades anatomical, functional, and metabolic findings on MRI and spectra of the human brain have been translated into hypothesis-based and systematic MRI studies in animal models of neurological
Life sciences
MRI instrumention, magnetic resonance pulse sequences, radio frequency coils, bioinformatics, physiological monitoring, optical technology, Innovative animal models
New knowledge Hypothesis
Behaviour Disease Genetic profiling
Figure 1: Illustration of the translational research concept of MRI in the neurosciences Processes and features on MRI images in human beings have lead to hypothesis-driven research in small-animal models, which have been made possible by innovations and fusion of knowledge at the physical and life sciences interface. Knowledge and understanding of MRI contrast signatures of the normal and diseased brain from MRI research in small animals have generated new knowledge, which has been translated back into clinical practice as diagnostic and therapeutic tools for various diseases.
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diseases, and the findings applied as diagnostic and therapeutic monitoring methods in clinical practice (figure 1). For example, diagnosis of acute brain ischaemia1 and detection and assessment of stroke lesion maturation,2–4 post-ischaemic recirculation, and treatment efficacy were all derived largely from such translational MRI studies.5–8 From these and many other studies, it has become clear that magnetic-resonance parameters do not represent specific histopathology such as apotosis, degeneration, selective neuronal injury, fibrosis, and inflammation but are associated and correlated with these features. In other words, a given magneticresonance contrast signature for one disease can be the same as in another. With the completion of the human genome project in 2003, translational magnetic-resonance studies have developed even further with the availability of new genetic mouse models of human diseases and the need for diverse and time efficient morphological, physiological, and metabolic phenotyping. The MRI field now encompasses a wide range of disciplines and expertise; these include highfield magnetic-resonance instrumentation, multispectral imaging, miniature radio-frequency coil design operable in high-field systems, small-animal magnetic resonance that is compatible with physiological monitoring equipment, magnetic-resonance contrast agents for molecular imaging, mouse brain neuroinformatics, and magnetic-resonance phenotyping at all levels. There are already many strong reviews available on magneticresonance technology in general,9 magnetic-resonance microscopy and small animal imaging,10,11 and magneticresonance histology.12 In this review we focus on two major applications in the neurosciences; structural and functional magnetic-resonance microscopy with emphasis on translational applications.
Physical and technical concepts High resolution MRI and magnetic-resonance microscopy MRI was discovered more than 30 years ago and it continues to develop at a rapid pace. The primary drive behind this rapid development is the desire to improve http://neurology.thelancet.com Vol 5 June 2006
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the signal to noise ratio and image contrast, which together define the sensitivity of the magnetic-resonance technique and detect differences between normal and diseased tissues, changes in those differences (by natural progression or after intervention), or changes caused by exogenous drugs. Foremost among the technological developments has been the development of ever-stronger magnetic fields to improve signal-to-noise ratio and spectral resolution. Furthermore, the scientific drive to rapidly translate magnetic-resonance-microscopy data to use in the clinic has led to the development of instruments with user interfaces and pulse-sequence encoding that match clinical systems (eg, the Bruker ClinScan 7-T small animal MRI system that combines the clinical user interface of Siemens’ MRI technology). These developments enabled pulse sequence development and testing in a clinical setting. The physical size of the sample being tested determines the available magnetic-field strength because the field achievable is proportional to the diameter of the main solenoidal coil in superconducting magnets. Wide-bore magnets (89 mm vertical bore) at 600–800 MHz are becoming widely available; the National High Magnetic Field Laboratory recently commissioned the world’s first 900 MHz (11 cm bore) magnet.13 Small-animal systems (mainly 20–40 cm horizontal bore magnets) at 7–11·7 T are becoming more commonly available. 3 T human systems (80–100 cm bore) are now commonplace (and the clinical standard), 7–8 T systems are being assessed, and a recent study used a 9·4 T human system;14 an 11·7 T human system has also been proposed.15 However, after insertion of ancillary technology that is needed for imaging (gradients, shims, radio-frequency coils) only 40–60% of the magnet bore is available for the sample. Thus the higher field systems accommodate mice and isolated tissues, the animal systems small mammals up to small primates, and the human systems can obviously be used for human beings but also for larger animals. The most important determinant of the signal-to-noise ratio is the size of the radio-frequency coil (the ratio is linearly dependent on the radio-frequency coil diameter), which is used to detect the signal. Therefore, the size of the coil should be matched to the size of the object (or region on a larger object) for optimum signal to noise ratio. The smallest coils are only tens or hundreds of micrometres in diameter and have been used to image large, isolated, single cells.16 Array coils have been developed that can further improve signal-to-noise ratio, imaging speed, and image coverage.17 These array coils work by recognising that smaller coils give better signalto-noise ratio but a small field of view, and so compensate for the small field of view by using multiple small coils simultaneously. As a consequence, an array of independent receivers is needed. These coils are now standard on clinical machines and are being developed for high-field animal and vertical bore systems. http://neurology.thelancet.com Vol 5 June 2006
Given this basic technology, the image quality achievable in MRI depends on a complex interplay of the imaging technique (pulse sequence) used to achieve the desired type of contrast (T1, T2, diffusion, or magnetisation transfer) and the imaging speed needed (especially for functional imaging). Once these factors are determined, the signal-to-noise ratio will dictate the spatial resolution that can be attained. Other factors also affect image quality and signal-to-noise ratio, including, but not limited to, susceptibility effects, motion, in vivo sample viability. and temporal stability. If resolution is less than 100 μm in at least one spatial dimension, then this is typically referred to as “magnetic-resonance microscopy”.16 If resolution is above 100 μm (but below standard clinical imaging), imaging is usually referred to as “high resolution MRI”. These terms are commonly used interchangeably, but we stress that it is important not to be fixated on resolution. Imaging speed and contrast are equally important and, together with resolution, dictate what can actually be seen, thus determining the sensitivity and specificity of the technique. Generally, a resolution of tens of micrometres can be achieved on small samples (a couple of centimetres or less) in minutes to hours in a high magnetic-field strength. On larger samples at lower magnetic fields or over shorter times, especially in-vivo or functional studies, the resolution relaxes to about 50–200 μm, although this can depend on what the sample is and exactly what needs to be detected. A better idea of what can be achieved is provided by the many examples in the rest of this review.18,19
Contrast generation via exogenous agents Classic contrast agents influence the relaxation times of water in their immediate vicinity and include gadolinium derivatives (causing signal hyperintensity) and super paramagnetic oxide particles (causing signal hypointensity). These agents have great clinical utility, because they can leak out into and highlight damaged tissue, however, they are restricted to extracellular spaces. “Smart” or targeted contrast agents have been developed that attach to specific tissues (eg, tumour cells) or become activated only in the presence of another agent. Some of these agents have the potential to work their way into cells for monitoring gene expression,20 and agents based on nanoparticles and associated technology are being rapidly developed.21 Although these new agents have promising potential, the difficulties involved in their application cannot be underestimated. For example, the concentration of these agents is so small that ways to amplify their potential signal must be developed if they are to be clinically useful.22 In addition, when signal changes are potentially small, the image spatial resolution and acquisition time might be compromised. There is also great concern about the potential toxicity of these agents and the 537
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body’s ability to remove them: it may be necessary to develop additional agents to actively remove the contrast agent after they have been given. Because these methods are under development, they will not feature extensively in this review, which instead will concentrate on the better-established morphological and functional magnetic-resonance techniques. They remain however, an exciting area with great potential, particularly for studying the brain.23
In-vivo magnetic-resonance microscopy Awake animal studies Most animal studies with magnetic-resonance microscopy require that the animal is anaesthetised for ethical reasons and to prevent motion artifacts, stress, discomfort, and pain if surgical procedures are involved. A few magnetic-resonance microscopy studies that examined evoked cortical responses and absence seizures using blood oxygen level dependent (BOLD) contrast have been done with awake rats.24,25 However, functional MRI activation patterns in the brains of awake animals must be interpreted with careful parallel assessment of potential concomitant stress in the animal (eg, by noting stress markers such as increased heart and respiratory rates as well as neuroendocrine responses). In addition, in studies in awake animals in which functional MRI responses to drugs like cocaine and narcotics are investigated, extra attention must be paid to physiological parameters during data acquisitions because of the effects these drugs have on respiration and heart rate, which can confound the BOLD signal interpretation.26–28
Anaesthesia An ideal anaesthetic would keep the animal asleep, free from pain and stress, immobilised, haemodynamically stable, and not interfere with breathing, heart rate, and the physiological process to be studied. However, on the basis of these criteria there is no perfect anaesthetic and so when choosing one for a given magnetic-resonance microimaging study it is important to consider several factors. These factors include length of study, recovery time, intravenous access, spontaneous or mechanical ventilation, disease status of the animal (important for certain transgenic mouse strains), type of experiment (eg, anatomical, spectroscopic, or functional data acquisitions), and interference with other drugs during the study. For functional MRI studies, chloral hydrate, alpha chloralose, or propofol are commonly used because they interfere less with sensory evoked potentials, including brainstem auditory potentials.29,30 Alpha chloralose has been successfully used in functional imaging studies in rats characterising brain activation patterns during forepaw and hindpaw stimulation,31–33 visual stimulation,34 whisker stimulation,35 noxious stimulation,36 with cocaine administration,28,37 and with other drugs.27 538
Physiological monitoring equipment In the earliest MRI studies in rodents, no physiological parameters were monitored and during imaging the animal was wrapped loosely in plastic sheeting for immobilisation (and probably to control temperature) and placed supine in the imaging field.38 Over the past two decades, increasingly sophisticated anaesthesia delivery systems, physiological monitoring, and supportive equipment for small animals have been developed.39–44 Physiological monitors suitable for rats and smaller rodents are available to monitor respiratory rate, tidal volume, ventilation pressure (if mechanically ventilated), body temperature, heart rate, and electrocardiography. Typically the monitors are operated by fibre optics and therefore are magnetic-resonance compatible (eg Module 224002, Small Animal Instrument Inc Stony Brook, NY, USA). Importantly, the fibre-optic monitors function in magnetic-resonance systems, which operate at field strengths in excess of 11 T and can track heart rates up to 700 beats per minute. These new monitoring devices are very useful in magnetic-resonance microscopy studies on transgenic mouse models of neurological diseases such as multiple sclerosis, amyotrophic lateral sclerosis, or spinal-cord defects or injuries where the animals have increased sensitivity to anaesthetics and are therefore at risk of respiratory or cardiovascular compromise.
Morphology of mouse models of neurological diseases The ability to resolve rodent brain anatomy on MRI is essential not only for the assessment of macroscopic abnormalities, such as atrophy, hydrocephalus, demyelination, plaque deposition, or periventricular leucomalacia, but also to accurately quantify more subtle morphological changes that might occur in animals treated with drugs, disease models, or animals exposed to environmental and behavioural stressors. For genetically engineered mice, it is important to understand how the gene in question affects the morphological phenotype and whether the engineered genotype produces the desired phenotype—ie, does it actually mimic clinical disease? For example, the desired pathology for a mouse model of multiple sclerosis would be demyelination, demyelinated plaques, and reactive gliosis whereas for attention hyperactivity disorder it might be cerebellar volume asymmetry.45 Quantitative analysis of rat and mouse brain structures can be obtained from in vivo as well as in-vitro magnetic-resonance microscopy (table).46–56 A superoxide dismutase (SOD)1G93AG1H transgenic mouse—an animal model of familial human amyotrophic lateral sclerosis—was recently characterised by magneticresonance microscopy.57 These mice develop tremor, paresis, and paralysis of the hind limbs due to loss of motor neurons and degeneration of the CNS. On T2weighted magnetic-resonance microscopy, a 3 month SOD transgenic mouse had areas of hyperintensity in the http://neurology.thelancet.com Vol 5 June 2006
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Familial human amyotrophic lateral sclerosis
Strain/mouse model
MRI findings
Superoxide dismutase SOD1G93A G1H
T2-weighted MRI: areas of hyperintensity (neurodegeneration) in nucleus ambiguous, facial nucleus, and motor trigeminal nucleus46
Multiple sclerosis
Cuprizone model
T2-weighted MRI: enlargement of ventricles; hyperintesnsity (demyelination)47
Alzheimer’s disease
APP23 mice
Quantitative diffusion imaging (ADC): local ADC decreases in APP23 mice48
Perinatal brain injury
Human S100B transgenic (Tg) and S100B knock-out (KO) mice
T2-weighted MRI: hemispheric volume loss after ischaemia49,50
Neural stem cell injected into mouse brain
C3H mice
T2*-weighted magnetic resonance images: hypointensity associated with stem cells (previously incubated with SPIO)51
Duchenne muscular dystrophy
Mdx mouse
Quantitative T2: no change in frontal cortex52
Knobloch syndrome
Col18a1(-/-) mice
T2-weighted MRI: hydrocephalus53
Leber hereditary optic neuropathy, Leigh syndrome
Intraocular injection of recombinant rAAV in DBA/1J mice
T1-weighted MRI+contrast: thinning of optic nerve, not rescued with adenoassociated virus to deliver the human gene for SOD254
Cerebral ischaemia
C57BL/6J mice ApoE-deficient
Diffusion-weighted MRI: brain atrophy; enlargement of ventricles55
Head injury
C57BL/6J mice ApoE-deficient
Diffusion-weighted MRI: unilateral hemispheric swelling, obliteration of lateral ventricle56
Table: MRI findings in mouse models of human disoders
brainstem (nucleus ambiguous, rostroventrolateral reticular nucleus, and the lateral paragigantocellular nucleus) and trigeminal motor nuclei, which correlated with areas of vacuolar degeneration on histological sections.57 As expected, degeneration of the nucleus ambiguous and trigeminal nucleus was associated with deficits in the movement of larynx, pharynx, and reduced mastication similar to that seen in human beings with bulbar onset amyotrophic lateral sclerosis.57 These mice develop respiratory deficiencies and it is methodologically demanding to follow the long-term sequelae by magneticresonance microscopy; only one timepoint (3 months) has been investigated so far. The various transgenic mouse models of Alzheimer’s disease are indicative of the search for defining specific genes that affect the pathology of the disease and allow the assesment of potential therapies.57 An ideal mouse model for Alzheimer’s disease would display behavioural equivalents of cognitive decline (memory deficits) and hallmark pathological features such as deposition of amyloid plaques, neurofibrillary tangles, reactive gliosis, dystrophic neuritis, neurodegeneration, and brain atrophy. Amyloid-plaque detection has been investigated with radiolabelled positron emitters, PET, and single-photonemission CT.58–63 However, MRI detection of plaques in vivo was recently achieved with a fluorine-19-labelled compound that specifically labels amyloid-β plaques in the brain.64 FSB—(E,E)-1-fluoro-2,5-bis(3-hydroxycarbonyo4-hydroxy)styrylbenzene—an amyloidophilic Congo-red type compound that crosses the blood–brain barrier, was intravenously injected into APP transgenic mice and subsequently imaged by T1-weighted magnetic-resonance contrast;64 the amyloid deposits were clearly visualised.64 Brain atrophy can also be detected by magneticresonance microscopy in vivo55 and in vitro;65 because atrophy in the temporal poles and entorhinal cortex are one of the earliest clinical signs of early cognitive decline during Alzheimer’s disease in human beings, magnetichttp://neurology.thelancet.com Vol 5 June 2006
resonance microscopy can be used to assess the clinical utility of transgenic mouse models of Alzheimer’s disease.66 Magnetic-resonance microscopy has been used to assess the development of brain atrophy in formalinfixed brains of transgenic mouse models that overexpresses human mutant amyloid precursor protein (V717F) under control of the platelet derived growth factor promoter (PDAPP mice) at 40, 100, 365, and 630 days.65 In agreement with clinical data, the investigators reported a 12% reduction of hippocampal volume in the transgenic mice at 100 and 630 days compared with the wild-type control group.65 In addition, the corpus callosum was substantially longer in the control mice compared with the transgenic mice.65 In-vivo preliminary magneticresonance microscopy studies of PDAPP mice have not
Figure 2: MRI of a homozygous PDAPP mouse before and after formalin-fixation Left: diffusion-weighted MRI microscopy of a homozygous PDAPP mouse acquired with 7·1 T intrument at a spatial resolution of 0·0012 mm³ at optimum contrast-to-noise for visualisation of mouse anatomy in vivo. Right: same mouse brain after formalin-fixation. The formalin fixed brain was excised from the skull and imaged with a T2*-weighted contrast sequence at a spatial resolution of 0·00012 mm³ with a 9·4 T solenoid coil (CIVM, Duke University). Images generated by H Benveniste on magnetic-resonance systems at the Center for In Vivo Microscopy, Duke University.
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Figure 3: Diffusion-weighted images of a homozygous PDAPP and wild-type control Top: diffusion-weighted MRI of a homozygous PDAPP mouse at the level of the dorsal hippocampus (left) and ventral hippocampus (right). Bottom: diffusion-weighted MRI of a wild-type control mouse at the level of dorsal hippocampus (left) and ventral hippocampus (right); arrows show the lateral ventricles, which are clearly larger in the wildtype control compared with the PDAPP mouse.
confirmed brain atrophy at age 22 months67 probably because image contrast differs between in-vivo and invitro specimens (figure 2). A striking abnormality in PDAPP mice is a smaller ventricle volume compared with their control litter mates, the reasons for which are unknown (figure 3). The discrepancy between in vitro and in vivo findings is clearly associated with the difficulty in defining anatomical areas of interest on MRI in vivo because of scan time restriction and motion, which decreases overall contrast-to-noise ratio (borders between anatomical areas are obscured), signal-to-noise ratio, and therefore spatial resolution of the image. Additionally, the formalin fixation process in vitro chemically changes tissue characteristics so that the image contrast-to-noise ratio is improved compared with that obtained in vivo. Figure 4 shows that anatomical delineation is better in the T2*weighted MRI acquired from a formalin-fixed specimen 540
on a 17·6 T instrument at a spatial resolution of 0·001 mm³ (scan time 5·7 h) compared to a T2-weighted and diffusion-weighted MRI acquired in vivo on a 9·4 T instrument at a spatial resolution of 0·001 mm³ and 0·0047 mm³, respectively. The difference in segmentation accuracy on in-vitro magnetic-resonance microscopy images versus in-vivo is indicated in the coefficient of error (measure of how much spread there is in the set of values relative to the mean) of the volume measurements. For example, hippocampal volume assessments (in vitro) in C57BL6/J mice had a mean of 25·7 mm³ (SD 1·1),68 18·61 mm³ (1·0),68 and 21 mm³ (0·5),69 equal to a coefficient of error of 0·04 (1·1/25·7), 0·05, and 0·02, respectively compared with in-vivo hippocampus volumes of 12·1 mm³ (1·4) and (dorsal hippocampus only) 8·9 mm³ (0·9)55 equal to coefficient of error of 0·12 and 0·10, respectively. The large coefficients of error in vivo are likely related to difficulties with segmentation and affect statistical power if the sample size is not adjusted. For example, for a given set of experiments (independent samples) with a coefficient of error of 0·04, at the significance level of 0·05 (two-sided), five in each group would yield a power of 93% to detect a 10% change with the independent samples t test; however, for the set of experiments (independent samples) with a coefficient of error of 0·1, 20 in each group would give a power of 87% with the independent samples t test. However, the greater segmentation accuracy in vitro is offset by greater potential for errors from tissue shrinkage and other artifacts of fixation. These problems in quantification of volumetric in vivo changes can be overcome by overlaying, aligning, and warping better segmented brain atlas templates onto the in-vivo images.70 This approach is routinely used in human studies and is increasingly being applied in rodent studies using MRI and other imaging modalities as more digital rodent atlases become available. A recent study showed the advantages of using a contrast agent for in vivo imaging of mice. In this study mice were injected intraperitoneally 48 h before imaging with 20 mM/kg of MnCl2 which accumulates in brain cells and highlights contrast to noise ratio in cell-rich areas.71 Bock and colleagues71 were able to acquire in vivo MRIs with an isotropic resolution of 156 μm³ in less than 3 h. However, the actual anatomical segmentation advantage of this method still needs to be assessed.
Functional MRI in rodents Measurement of haemodynamic changes associated with brain activity using BOLD functional MRI is used extensively for cognitive and behavioural studies in normal and diseased human beings. Functional MRI studies in rodents have progressed slowly and are especially challenging because they are usually done in anaesthetised animals in which functional MRI signals are dampened or changed. Further, as the animal is not awake it cannot be imaged when performing tasks. http://neurology.thelancet.com Vol 5 June 2006
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Figure 4: Anatomical resolution in C57BL6/J mouse brain images acquired in vivo and in vitro with different magnetic resonance contrast techniques Left: in vivo T2 weighted RARE (rapid acquisition with relaxation enhancement) sequence, 100 μm×100 μm×100 μm, 9·4 T. Middle: in vivo diffusion-weighted spin-echo image, 0·009 mm3, 9·4 T. Right: in vitro T2*-weighted image, TR (repetition time)=150 ms, TE (echo time)=7·5 ms, 0·001 mm3, 17·6 T.
Therefore, functional MRI experiments of rodents need implementation of an externally applied stimulus, such as electrical stimulation, odour, visual stimulation, air puffs, chemical, drugs, or noxious stimulation. One of the earliest studies by Hyder and colleagues33 reported functional MRI activation patterns in the brain of an alpha-chloralose anaesthetised rat during forepaw stimulation. The forepaw stimulation paradigm has been used by several other investigators and is favoured over hind-paw stimulation because of susceptibility artifacts associated with skull structures (tympanic bulla and zygomatic arch) in the somatory sensory hind-paw area. Gyngell and colleagues31 examined variations of the functional MRI signal in response to the frequency of somatosensory stimulation and found that the activation intensity decreased for forepaw stimulation frequencies above 3 Hz, and was negligible at 9 Hz. High-resolution BOLD-based functional MRI images of this same paradigm revealed that the functional MRI response was maximum at the cortical surface and in cortical layer 4.32 Improvement of the temporal resolution of the functional MRI experiments from 18 s33 to 50 ms32 showed that the forepaw activated functional MRI signal starts deep within the layers of the cortex and then expands to outer parts of the somatosensory cortex.23,32 The usefulness of functional MRI studies in anaesthetised rodents may seem questionable, but important information can still be obtained. For example, assessment of integrity and alteration (plasticity) of neuronal pathways after a stroke can be studied and assessed using functional MRI in the anaesthetised rodent. Dijkhuizen and colleagues72 studied the correlation between infarct size and brain function recovery with functional MRI. Rats middle cerebral arteries were occluded for 2 h before anatomical and functional MRI after forelimb stimulation were done. http://neurology.thelancet.com Vol 5 June 2006
The total area of significant forelimb-stimulation-induced activation was calculated in each hemisphere and the laterality index was calculated, this index shows activity in the contralateral hemisphere relative to that in the ipsilateral hemisphere.72 The study showed that acute stimulation of the impaired forelimb (days 1 and 3) was accompanied by acute bilateral hemispheric activation (days 1 and 3); these findings are similar to those in patients with stroke and indicate ongoing neuroplasticity in the recovering brain. Furthermore, the laterality index was inversely related to lesion volume.72 Most functional MRI studies use BOLD contrast. However, an alternativee approach is to preload the brain with paramagnetic manganese (Mn-enhanced MRI) for 12–24 h, then expose the animal to a given stimulus when it is awake and doing tasks, and subsequently scan the anaesthetised animal using T1-contrast.73 Manganese enhances contrast in the rodent brain in several brain regions (olfactory bulb, ventricles, corpus callosum, CA3, dentate granule cells, amydala, and cerebellum) in a time-dependent manner.74 Manganese ions enter synaptically activated neurons through voltage-gated calcium channels23,74,75 and if electrically or pharmacologically stimulated, a signal enhancement will be visible on T1-weighted images. This approach is very similar to brain activation studies done using radiolabelled fluoro-deoxy-glucose and PET in which the stimulus is given at the time of injection of the radioactive tracer when the animal is awake and cognitively challenged; 30–40 min later when the tracer has been trapped intracellularly the PET scan is done during which the animal can either be awake or anaesthetised. Fluorodeoxyglucose PET or Mn-MRI approaches necessitate that the stimulus is constant, prolonged, or repeated during the uptake periods, and the brain activation pattern reported is therefore an integration of the activity seen in this interval. 541
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Yu and colleagues73 gave MnCl2 as an intraperitoneal injection to mice at a dose of 0·4 mmol/kg and after 24 h the mice were exposed to a sound-evoked auditory stimulus delivered at different frequencies when the animals were awake and freely moving in a specially-designed cage.73 A careful analysis of signal enhancement in the inferior colliculus revealed frequency-dependent tonotropic organisation. This is a unique new approach, which might be superior to BOLD-based functional MRI experiments in mouse brains where physiological stability, motion control, and functional MRI signals can be difficult to obtain in the anaesthetised state because of the small size. Functional MRI studies in transgenic mouse models using external stimuli or drugs are not yet routine given the difficulties involved in doing such studies. However, with the advancement in radio-frequency coil technology, positioning devices, pulse sequences, and hardware (shim tools) this technique will become more widespread. There are many transgenic mouse models, in which functional MRI approaches would provide insight into pathophysiology and disease processes. For example, dopamine transporter deficient mice and mice transgenic for human dopamine receptors DA 2 and DA 4 have been used in the study of several diseases such as addiction, attention hyperactivity disorder, and Parkinson’s disease, and we have not yet phenotypically explored functional neurocircuitry in these models.
Conclusion We hope that this review has highlighted the importance of magnetic-resonance microscopy in translational neuroimaging studies. Growing research is increasingly focused on improving specificity of magnetic-resonance contrast and attempts to extract multicompartmental information from magnetic-resonance parameters. This will, to a large extent, determine the future of the clinical application of magnetic-resonance microscopy studies. The latter relates to the longterm goal in magneticresonance studies to use quantitative measurements to improve the sensitivity and specificity of MRI, especially with regard to T1 and T2 measurements. These goals have proven difficult to achieve, primarily because of inter- and intra-sample biological heterogeneity and variability in imaging techniques. Additionally, the data acquisition times are such that there is considerable exchange between tissue compartments at the cellular level, making data interpretation difficult. More recently, the search for quantitative techniques has been ressurected by the observation that diffusion measures can reveal multicompartmental information and can be collected with short echo times minimising compartmental exchange. Further, these compartmental signals change with tissue perturbations in a way that is consistent with cell swelling. Model systems are under detailed experimental and mathematical modelling investigations and it remains to be seen if these measures will have a clinical significance. 542
Search stategy and selection crtieria References for this review were identified by searches of MEDLINE and MEDLINEplus between 1969 and January 2006 and from relevant articles; numerous articles were identified through searhces of the extensive files of the authors. The search terms “MR microscopy”, “MR imaging”, “rodent”, “functional MRI”, “fMRI”, “neuroinformatics”, “morphology”, “brain”, “brain atlas” were used. Abstracts and reports from meetings were also included. Only papers published in English were reviewed. The final reference list was generated based on originality and relevance to the topics covered in this review.
Contributors Both authors contributed equally to the article. Conflicts of interest We have no conflicts of interest. Acknowledgments The authors would like to acknowledge the financial support of the NIH through grants R01 EB 00233-04 and P41 RR16105 and the National High Magnetic Field Laboratory. MRI data were obtained at the Advanced Magnetic Resonance Imaging and Spectroscopy facility in the McKnight Brain Institute of the University of Florida. In addition, we would like to thank the SBU/BNL 9.4T magnetic-resonance microscopy facility at Brookhaven National Laboratory. References 1 Kucharczyk J, Mintorovitch J, Asgari HS, Moseley M. Diffusion/ perfusion MR imaging of acute cerebral ischemia. Magn Reson Med 1991; 19: 311–15. 2 Gill R, Sibson NR, Hatfield RH, et al. A comparison of the early development of ischaemic damage following permanent middle cerebral artery occlusion in rats as assessed using magnetic resonance imaging and histology. J Cereb Blood Flow Metab 1995; 15: 1–11. 3 Minematsu K, Li L, Sotak CH, Davis MA, Fisher M. Reversible focal ischemic injury demonstrated by diffusion-weighted magnetic resonance imaging in rats. Stroke 1992; 23: 1304–10. 4 Verheul HB, Berkelbach van der Sprenkel JW, Tulleken CA, Tamminga KS, Nicolay K. Temporal evolution of focal cerebral ischemia in the rat assessed by T2-weighted and diffusion-weighted magnetic resonance imaging. Brain Topogr 1992; 5: 171–76. 5 Lo EH, Matsumoto K, Pierce AR, Garrido L, Luttinger D. Pharmacologic reversal of acute changes in diffusion-weighted magnetic resonance imaging in focal cerebral ischemia. J Cereb Blood Flow Meatb 1994; 14: 597–603. 6 Reith W, Hasegawa Y, Latour LL, et al. Multislice diffusion mapping for 3-D evolution of cerebral ischemia in a rat stroke model. Neurology 1995; 45: 172–77. 7 Roussel SA, van Bruggen N, King MD, et al. Monitoring the initial expansion of focal ischaemic changes by diffusion-weighted MRI using a remote controlled method of occlusion. NMR Biomed 1994; 7: 21–28. 8 Steinberg GK, Kunis D, DeLaPaz R, Poljak A. Neuroprotection following focal cerebral ischaemia with the NMDA antagonist dextromethorphan, has a favourable dose response profile. Neurol Res 1993; 15: 174–80. 9 Tyszka JM, Fraser SE, Jacobs RE. Magnetic resonance microscopy: recent advances and applications. Curr Opin Biotechnol 2005; 16: 93–99. 10 Pautler RG. Mouse MRI: concepts and applications in physiology. Physiology (Bethesda) 2004; 19: 168–75. 11 Pirko I, Fricke ST, Johnson AJ, Rodriguez M, Macura SI. Magnetic resonance imaging, microscopy, and spectroscopy of the central nervous system in experimental animals. NeuroRx 2005; 2: 250–64. 12 Henkelman RM, Chen XJ, Sled JG. Disease phenotyping: structural and functional readouts. Prog Drug Res 2005; 62: 151–84.
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