Another important issue when evaluating ICP monitoring modalities is the possible role of pressure gradients on the ICP scores displayed to the health care personnel. In this regard, some questions arise: what is the value of measuring supra-tentorial ICP in subjects with infra-tentorial mass lesions, and how representative is an ICP measured in the right hemisphere in an individual with a lesion in the left?
A growing lesion will create pressure gradients needed to displace tissue and fluids, which in turn will affect the measured ICP. This concern is relevant both in TBI and in evolving hydrocephalus [ 96 ].
Consequently, rapid treatment should be considered in patients with growing localized lesions and accompanied with symptoms, despite any discrepancy between clinical symptoms and ICP.
Moreover, pressure gradients may exist between the cranio-spinal compartments and should be kept in mind when making important clinical decisions. Comparing pressure scores between the intracranial and lumbar compartments revealed differences in both static ICP and ICP wave amplitudes [ 97 ]. In the case of pressure gradients, there should be differentiation between static and pulsatile gradients in ICP.
This question has been studied in the context of hydrocephalus, and particularly whether an outward pressure gradient can explain growing ventricles. In chronic cases, there were no trans mantle gradients in static ICP in individuals with communicating or non-communicating hydrocephalus [ 98 ], and no trans-mantle gradients in ICP wave amplitudes in communicating hydrocephalus [ 99 ].
Others reported that gradients in static pressure between the ventricular and parenchymal compartments could be attributed to hydrostatic pressure gradients, while differences in ICP wave amplitudes were minor [ 58 ]. Accordingly, the results may depend on the ICP scores in question.
The term pulsatile ICP refers to the pressure changes occurring during the cardiac cycle. Each heartbeat results in intracranial pressure variations in accordance with the cardiac cycle measured as the ICP waveforms see Fig.
Typically, a continuous ICP signal varies over time, characterized by a diastolic minimum pressure value and a systolic maximum pressure value, causing the calculated ICP scores to vary. This is further illustrated in Additional file 2 : Movie 2. Established attributes from the single wave amplitudes are the amplitude pressure difference between diastolic and systolic pressures , the rise time time from diastolic to systolic pressures and rise time coefficient amplitude divided by rise time, providing a measure of the steepness of the ICP waveform [ 32 ].
How the ICP waveforms are presented in the clinic varies according to the different monitors and devices. Some monitors show a few seconds of data providing poor time resolution, while others present long-time series with poor spatial resolution, which makes it a challenge to access the various morphological features. Some ICP equipment allows for storage of the ICP waveform data for post-processing, which is beneficial for research but has limited value in daily clinical practice.
This has been defined and studied in different ways throughout the past decades, but the amplitude metrics AMP and MWA are the main references in the clinical literature [ ]. The single wave amplitude AMP is determined in the frequency domain [ 63 , 73 ].
The first harmonic of the frequency Fourier spectrum corresponds to the heartbeat, and the value of the first harmonic can be used to estimate the single wave amplitude. The average of identified single wave amplitudes over a defined period 6 s is the MWA. These methods are not equivalent and provide different measures of the ICP wave amplitude [ ], which constitutes an obstacle when comparing results in the literature. To date, the transition to clinical decision making has been made only in a few locations.
Differences between the time and frequency domain methods are further illustrated in Fig. Differences between the time- and frequency-domain methods for estimating ICP scores. The pressure waveforms are usually presented in the time domain upper panel. The lower plots present the pressure data analyzed in the frequency domain, represented as a function of frequency.
The signal and the defined cardiac components separated from low frequency components such as respiration can be analyzed independently. Additional information available with frequency domain analysis is the phase, the frequency domain analog of timing in the time domain not shown.
The phase plot allows analysis of timing differences between the ICP and the reference waveform for each identified frequency component. ICP wave amplitude, slope, between peak time and has been used in an attempt to forecast increased ICP [ ]. The normal values of mean ICP wave amplitude MWA have not been determined since measurements cannot be performed in healthy individuals.
The thresholds determined from our experience refer to MWA scores obtained in individuals who have undergone ICP monitoring, but where no evidence of abnormality was found. One advantage of the ICP waveform derived indices, as compared to the ICP-derived indices, is the independence of reference pressure variability. The limitations related to the assessment of single ICP wave metrics involve the physiological processes creating single ICP waves and technological issues related to the proper identification of single ICP waves.
We have limited knowledge of the mechanisms affecting single ICP wave morphology. Since the early exploration of single ICP waves in the s, experimental evidence suggests that the single ICP waves are affected by cerebral blood volume changes [ ], and that cerebral blood volume therefore may affect the ICP wave amplitudes [ , ]. On the other hand, the ICP wave amplitudes in individuals with iNPH were not related to cardiovascular parameters such as arterial blood pressure, cardiac output, stroke volume, oxygen consumption and systemic vascular resistance [ ].
Research indicates that the pulsatile ICP waveforms may also be affected by body position [ 72 ] and day-night cycles. As arterial BP waveform is the input signal to the ICP waveform, changes in vascular wall properties vascular compliance could also impact the ICP waveform. Given the influence of all physiological variables on the ICP waveform, it is not surprising that plotting MWA over time also reveals time-related variation see Fig. Therefore, the question of the degree to which these aspects impact the utility of pulsatile ICP monitoring remains unanswered.
One limitation for monitoring of single ICP wave-derived parameters in the clinical context is a methodology for identification of the single ICP waves. Regardless of the method used to measure ICP invasive, less invasive or non-invasive , the limited control of the ICP source signal means that the ways by which to control the information provided to the health care personnel are limited.
The ICP source signal may be corrupted for a number of reasons. Clearly, erroneous ICP scores provided to the health care personnel may result in inappropriate patient management. One main avenue for improvement of ICP monitoring practice is through incorporation of automatic algorithms for single wave identification along with ICP waveform analysis becoming part of standard practice, which would enhance control of the ICP source signal. The present review is largely based on the clinical experience done in this department throughout the early s.
Automatic identification and characterization of single ICP waves represents a challenge that requires a dedicated methodology [ ]. Such errors will affect the measurement of all wave attributes, including amplitudes, rise time and rise time coefficient.
Since measurement of pulsatile ICP metrics is highly technology-driven, differences in methodologies represent a limitation in terms of the build-up of knowledge.
However, depending on how ICP wave amplitudes are determined, these scores do not appear to be influenced by zero pressure levels [ 45 , 90 ]. When single wave amplitudes are determined as relative values within the pressure signal itself, the scores are not influenced by the absolute reference pressure. While pulsatile ICP refers to the pressure fluctuations during the cardiac beat contradiction, respiratory waves are low frequency fluctuations in static mean ICP related to respiration.
Respiratory waves occur at a slower frequency than the cardiac part of the ICP signal with a frequency of about 0. In the context of ICP monitoring, these respiratory waves are referred to as slow waves [ 63 ]. Assessment of these waves has not been implemented in current monitoring systems because they require dedicated software.
Therefore, the assessment of respiratory slow waves has come into clinical practice only to a very limited extent and is primarily done for research purposes. C waves are more frequent about 4—8 waves per minute elevations of mean ICP up to about 30 mmHg. While attention has been given to the role of B waves [ ], it has been difficult to incorporate assessment of B waves clinically because of difficulties in defining and quantifying them.
A recent review addressed the diversity in definitions of B waves and the variability of their presentation [ ].
B waves may have different patterns Fig. The clinical implications of the B waves frequency and magnitude content is a research question that remains to be determined. From Martinez-Tejada et al. B waves are elevations of mean ICP that may have different patterns. The timescale is in the order of minutes. It should be noted that amplitudes of B waves and single ICP waves are fundamentally different. There are currently limited options for measuring ICC directly in the clinical context, although this has been an objective since the introduction of clinical ICP monitoring in the s.
It has proven to be difficult to measure, however, without causing too much damage. The different approaches to measuring ICC are briefly mentioned in the following sections. The experiments required to investigate the intracranial pressure—volume relationship were highly invasive and was first explored in animal experiments [ , , , , ]. The animals were exposed to an intracranial volume increase with simultaneous measurement of ICP, which demonstrated a non-linear pressure—volume relationship see Fig.
Based on the experimental studies on intracranial volume—pressure relationships, the volume—pressure test VPT described the increase in ICP caused by administration of a minor volume of fluid to the ventricular CSF.
From this, less invasive clinical approaches to assess ICC evolved. Different scores for the pressure—volume reserve in the clinical situation were developed, including the pressure—volume index PVI [ ], and volume—pressure response VPR wherein 1 ml was added or subtracted from the ventricular CSF [ ]. Favorable results from the clinical use of this monitor were reported [ ].
However, a major drawback with these methods of assessing ICC was the need to add or subtract volume in the intracranial compartment, which is invasive and risk-related.
In some clinical situations, even a minor volume change in individuals with impaired ICC might cause a harmful increase in ICP. Another approach to obtain information about ICC while avoiding artificial intracranial volume changes is deciphering ICC from the single ICP waveform characteristics [ , ]. Over the years, different approaches were reported: Szewczykowski et al.
According to this concept, the intracranial volume dV change per heartbeat was assumed to be rather constant, and the peak-to-peak amplitude an indicator of the ICC. Later, Czosnyka et al. The calculation of RAP was determined from the frequency domain method and involves extracting the amplitude of the fundamental frequency. Another method based on the frequency domain method is determining the centroid of the ICP power spectrum between 4 and 15 Hz , denoting the high-frequency centroid and introducing it as an indicator of the ICC [ ].
Mortality after TBI increased with an increasing mean high-frequency centroid value [ ]. Cardoso et al. According to this concept, when ICC is impaired, the tidal P2 and dicrotic P3 peaks exceed the systolic peak P1 with the disappearance of the dicrotic notch, while the systolic peak P1 exceeds the tidal P2 and dicrotic P3 peaks under normal conditions.
Others have more recently performed automatic identification of the ICP waveform peaks using an artificial neural network [ ], and confirmed an association between peak separation and ICE, but no relationship with resistance to CSF outflow.
Given the many approaches to quantify ICC [ 63 , , , ], evaluation of the ICP wave amplitude is one approach that has made its way into clinical practice [ 60 , ]. According to this concept, the MWA represents the pressure response to a net intracranial blood volume change of about 1 ml during each cardiac beat.
The MWA was found to correlate with the ICC measured by the Spiegelberg compliance monitor [ ], and with intracranial compliance computed during ventricular infusion testing [ ]. Since all physiological parameters vary over time, including the net intracranial blood volume change, we have incorporated monitoring of MWA over many hours, usually overnight, when monitoring is done for diagnostic purposes. The MWA is computed every 6 seconds, multi-hour monitoring, which provides several thousands of observations that may reduce the impact of variation over individual cardiac cycles.
When comparing three metrics of ICC i. As similar ICP levels can correspond to different ICC states, ICC monitoring could provide significant improvement in patient care as it would allow for early intervention in progressively worsening patient states [ ]. However, the non-invasive estimation of ICP wave amplitudes by phase-contrast MRI was not found feasible by others [ ].
Further studies are needed to determine the clinical benefits of ICC monitoring. Moreover, a randomized controlled trial showed significantly improved outcome in individuals with SAH who were managed according to MWA, as compared to traditional mean ICP guided management [ ]. The need for surgical penetration of the skull and dura for placement of parenchymal or ventricular sensors for extended periods of time limit the application of ICP measurements, especially outside of the ICU, where the bar for neurosurgical intervention is high.
This raises the need for alternative approaches to accessing information about ICP. The most widespread use of measuring ICP is via lumbar puncture LP and involves advancing a needle into the lumbar intrathecal space, which is linked on the other end to an external pressure transducer. The common way is to measure fluid level as centimeters of water H 2 O and use this as an indication of ICP. A requirement for this measurement methodology to provide relevant results is obstruction-free CSF communication pathways.
Lumbar CSF pressure measurements during so-called infusion tests also have a long tradition [ ]. In such procedures, the CSF pressure is measured during infusion of a fluid to the lumbar compartment and the pressure change in response to the administered fluid is interpreted as resistance to outflow of CSF. This is performed on a routine basis in several centers [ 63 ]. The main indication is to assess shunt dependency or shunt failure in individuals with tentative CSF circulation failure.
The pressure measured by LP depends on the position of the lumbar region relative to the head since hydrostatic pressure differences will determine how lumbar CSF pressure compares with ICP. It has been reported that CSF pressure measured by LP compares very well with ICP [ 27 ], while a similar perfect match was not reported by others [ 97 ]. Serious complications are rare, but LP is contraindicated in cases when very high ICP is suspected due to the possibility of brain herniation [ ].
Further, this is not strictly an approach for ICP measurement as it is performed in the spinal region. LP measurements with current devices, however, only reflect an instantaneous ICP value and are therefore not useful for continuous mean ICP monitoring in its current state. Placement of an ICP sensor outside the dura mater is less invasive than placing a sensor within the brain parenchyma or ventricular CSF.
As sensor placement in the epidural space would reduce the risks of subdural or parenchymal hemorrhage, epidural placement was explored in the earlier years of ICP monitoring. Patients with increased risks of internal bleeds such as those with acute liver failure [ ] or hemorrhagic diseases could theoretically benefit even more from such a sensor placement.
However, since the introduction of epidural ICP sensors, numerous studies have reported errors in mean ICP monitoring [ , ], typically reporting ICP as too high.
Epidural ICP monitoring was therefore discontinued in most centers, although some centers have demonstrated that epidural ICP measurements provide very accurate readings with regard to pulsatile ICP, both with parenchymal probes placed epidurally [ ] and with commercially available epidural probes [ ]. This could be beneficial for certain patients and should therefore not be disregarded completely. Epidural ICP measurements also require a trepanation.
While the risk of bleeds is reduced, it is not eliminated. There is also a corresponding risk of infection. Epidural ICP monitoring highlights the importance of ICP source signal control because the handling of the source signal determines the safety of the information provided. The major problem with epidural ICP sensors relates to the ability to offer reliable mean ICP measurements [ , ].
This is because of issues related to zero reference pressure. However, measurements of epidural ICP wave amplitudes are feasible and accurate but require dedicated software because an algorithm for single ICP wave identification is needed [ , ]. Accordingly, even though it is considered that epidural ICP measurements do not provide valid mean ICP scores, experience from measuring ICP wave amplitudes has illustrated the value of the technique.
The results are better when using dedicated epidural ICP sensors rather than ICP sensors designed for parenchymal use [ ]. Preclinical long-term measurements of ICP in animals is required to understand the normal regulation of ICP, as well as mechanisms behind abnormal ICP in brain disease or injury.
Animal models have utilized fluid-filled ICP catheters placed in the ventricles, cisterna magna, thecal sac, epidural and subdural spaces, and fiberoptic ICP measurement systems in brain parenchyma [ ]. One study compared simultaneous measurements from ventricular, cisterna magna and parenchymal pressure measurement devices, and found the ventricular ICP monitoring preferable in terms of accuracy and least brain damage [ ]. Another study reporting a novel method for epidural measurements of ICP in rats reported a strong correlation between ventricular and epidural ICP score [ ].
More recently, telemetric devices utilizing subdural ICP catheters have been introduced for long-term monitoring of ICP in rats [ , , ]. To obtain reliable data from preclinical studies of ICP, adherence to the methodological limitations are of the utmost importance. Similar methodological issues as seen in humans such as shifts and drifts in static ICP occur in animals.
Preferably, preclinical studies should sample the continuous raw ICP signals and include assessment of both the static and pulsatile ICP after single wave assessment so as to ensure correct pressure assessments. However, in freely moving animals, high signal-to-noise ratio is expected. Because of the risks related to invasive ICP measurements, numerous non-invasive ICP source signals have been explored.
The potential benefits of non-invasive ICP monitoring seem clear. This section, therefore, discusses state of the art, the limitations and weaknesses of the nICP source signals, the nICP source signal control, and the issue of measuring absolute ICP non-invasively from a clinical perspective.
As non-invasive approaches to ICP monitoring require other input data than direct ICP measurements, they all rely on a form of model that permits the non-invasively obtained source signals to be altered into signals or information that can be utilized by the clinicians.
The first model of relevance was the three-compartment Monro-Kellie doctrine, which was a concise description of a highly intricate system. In the decades that followed, numerous models of varying complexity have been proposed. Some describe the cerebrovascular system using mechanistic models [ , , , , ], while others rely on statistics and ideally huge amounts of data in order to identify top-level statistical relations the clinicians can utilize [ , , ].
Both categories are included when various approaches to non-invasive ICP estimation are described in the following sections. While model-based estimation clearly has the advantage of resonating with clinicians and aiding everyone in understanding the mechanics of the intracranial compartment, the black box models are an exciting approach as digitalization of health data and immense computational power is becoming a reality.
This approach seems logical, given that the arterial BP waveform serves as an input signal to the ICP waveform.
Mathematical approaches have been explored, including the use of central aortic BP waveforms [ ]. However, we did not find that central aortic BP waveforms could be used as an nICP source signal [ ] and the use of arterial BP measures alone seems less useful. We also stress the role of assessing and comparing the pressure waveforms and their time alignment when conducting such and similar research [ ].
The use of arterial BP signals has been combined with other non-invasive signals, in particular, transcranial Doppler TCD. The TCD technology was initially developed as a non-invasive tool for vasospasm detection after SAH and for evaluating cerebral circulation [ ]. It utilizes the principle of the Doppler effect.
As ICP can affect the blood flow and the cross-section of vessels, CBFV can provide added information about the state of the intracranial space. This is related to the fact that the physiological parameters vessel compliance, autoregulation, and arterial BP vary over time [ ]. A recent study, however, reported that a combination is superior to using only one of the metrics [ ].
An approach combining radial artery BP and CBFV waveforms with a mechanistic model has also provided promising results [ , ] with regard to mean ICP estimation. Another approach for estimating nICP using Doppler technology is the two-depth ophthalmic artery Doppler ultrasonography developed by Ragauskas [ ]. A more recent study validating this technique reported a good correlation between invasive and nICP scores [ ].
However, pulsatile ICP measurements are not possible with this approach, and continuous monitoring of CBFV waveforms must be improved before the technique can be clinically useful. The use of TCD is also not possible in certain sub-populations [ ] because a cranial window is required. It is well known that TCD is highly user-dependent as results depend on the direction by which vessels are approached.
Even minor changes in the direction of the probe may significantly affect the measured Doppler signal. Moreover, CBFV is significantly affected by other changes in physiology such as medications, autoregulation, and hyperemia. Therefore, it seems infeasible to monitor nICP over several hours. The cochlear aqueduct represents a connection between the CSF in the intracranial cavity and the inner ear, which allows for CSF exchange when patent. This technique has since been used as a tool for mean ICP assessment in various studies with variable, but not convincing, clinical results [ , , ].
Some approaches to non-invasive estimation of ICC through the otic connection have also been presented in the literature. Davids et al. Direct assessment and comparison of ICP waveforms and tympanic membrane pressure waveforms have also been conducted [ , ] and found to have limited clinical potential due to the patient dependent cochlear aqueduct. In our study Fig. Adapted from Evensen et al.
Tympanic membrane pressure TMP as a surrogate marker of intracranial pressure. The non-invasive TMP waveforms were measured in the outer ear and used as input for the estimation of non-invasive ICP. The resulting output from the combination of b1 and the inverse of b2 is presented in b3. A similar technique to the stapedial reflex base proposed by Marchbanks aims to use the change in otoacoustic emission that occurs when ICP changes.
The principle behind this technique is that evoked CSF pressure in the inner ear will alter the load on the stapes and, as a result, the sounds generated in the inner ear in response to sound excitation will change when ICP changes [ , ]. However, this approach to nICP estimation is also, to a large degree, subject to significant interpatient variability. Using otic source signals for nICP is dependent on a patent cochlear aqueduct, which serves as a mechanical filter for the transmission of ICP-derived signals.
In most patients, this filter provides a substantial distortion of ICP signals. At present, the use of otic source signals for nICP estimation does not seem useful for clinical application. Various radiological approaches have been proposed to identify increased ICP. Traditional measures such as ventricular size, however, are not useful. For example, in a study of patients, no significant correlation was found between invasively measured ICP and the size of the cerebral ventricles measured by computed tomography CT [ ], which is often highlighted as a symptom of elevated ICP.
A study of 20 TBI patients by Pappu et al. A review of various radiological measures used to assess raised ICP in children was published recently [ ].
One method that has attracted interest is combining MRI images with fluid mechanics to use the measures of blood and CSF volumes that enter and leave the brain during the cardiac cycle to compute a brain elastance metric that allowed for differentiating between elevated and normal ICP [ , ].
However, contradictory results have been reported in other studies examining pulsatile ICP information from MRI images [ , ]. Although potentially useful as a screening tool for very high ICP, brain imaging techniques alone are currently not reliable for clinical management [ , ].
They are also not appropriate for continuous assessments. Currently, ICP cannot be inferred from brain imaging modalities. If this became feasible, a significant flaw would be that it would only provide short-term assessments as repetitive measurements would not be possible. In addition, imaging modalities are costly, and thus not available in many settings. Acoustic-based approaches aim to estimate ICP from the acoustical properties of the skull or the constituents of the intracranial compartment.
One approach, proposed by Levinsky et al. This method is based on a source signal of Hz being sent from an earplug in one ear and received by an earplug in the other ear. The receiving earplug also recorded head generated sounds , which is important in the estimation of ICP.
A training set, where both invasive ICP measurements and TCA were accessible, was used to establish a mathematical model. The non-invasive estimates of ICP were found after splitting the recorded signal into different frequency bands corresponding to different processes in the body blood flow, breathing and test signal.
The study found a mean difference of 0. However, further validation of the approach has not been presented. Another well-documented approach is methods based on ultrasonic time of flight TOF measurements. The rationale behind these techniques is the assumption that the acoustic properties of the intracranial structures, as well as the size of the cranial vault, can change in the case of elevated ICP.
Such changes should then affect the propagation speed and frequency attenuation of emitted ultrasound pulses [ , , , ]. This has produced good results in small scale studies [ ].
One drawback of this technique involves variable data quality as well as its ability to only measure ICP relative to a known reference ICP value [ 59 ]. Currently, there are no validated acoustic methods for nICP measurement.
The extent to which acoustic signals will become useful as source signals is a matter for future studies. The CSF of the intracranial cavity also surrounds the optic nerve, which in turn is surrounded by a sheath of meningeal layers. It has been demonstrated that when ICP increases, the radial pressure increases in the CSF surrounding the optic nerve, causing the diameter of the sheet to expand.
There have been several approaches to quantify this diameter using different imaging techniques such as MRI, CT, ultrasound imaging and optical coherence tomography.
This approach has proven quite successful in separating the low and high levels of ICP by comparing population-averaged values to patient-specific measurements of the sheath diameter [ ]. In particular, the ultrasound-based variant of this technique can be seen to have promising clinical value due to its applicability and accessibility.
One advantage of this technique, especially the ultrasound-based solution, is its applicability and accessibility. These measurements can be completed within minutes of a TBI and are possible without much medical training. A recent systematic review and meta-analysis [ ] concluded that this technology is promising for nICP measurements. The main drawback of this technique is that it is unfit for continuous measurements.
In addition, its ability to separate between high and low ICP primarily makes it a triage tool and less valuable at the bedside of patients with less imminent diseases of the CNS. It should be noted that when there is interocular ONSD asymmetry in the same subjects [ ], it should be taken into account.
Ideally, both eyes should be assessed, but this may be difficult in traumatic cases because facial and eye damage is often involved. The nICP approaches referred to in this paper cover a subset of the contributions with the aim of highlighting the most promising and most investigated techniques.
For example, a possible technique that shows some promise is near infra-red spectroscopy NIRS [ ], which is a non-invasive technique for monitoring cerebral oxygenation, and potentially a source signal for nICP. A more comprehensive list can be found in other review papers [ 59 , ].
However, despite considerable effort over many years, these and other review papers [ 40 , , , , , , ] came to the same conclusion reached in this paper: no nICP technique manifests as a complete universal solution applicable for all patients in all situations.
However, there are a few techniques that seem promising for specific clinical situations. While trained physicians know how to recognize and manage patients with very high ICP, this might not be the case for individuals such as a soccer coach or for people operating in first-line health care. In cases of acute brain injury, the combination of portable ultrasound tools and measurements of ONSD seems promising [ ] as it may allow people who are not medical professionals to distinguish high and low ICP.
However, this approach is not suitable for continuous monitoring and does not provide any ICC metric. One technique that allows for continuous ICP monitoring to a greater extent is the mechanistic model-based approach based on CBFV waveforms [ , ].
However, long-term, reliable continuous monitoring of these is challenging with the current measurement equipment. In addition, this approach does not allow for any computation of an ICC metric.
There are a limited number of windows into the cranial cavity, especially in adults, that could allow for non-invasive monitoring of ICP. It is important to keep in mind that there are several different clinical areas for nICP monitoring, which has implications for how we assess the nICP source signals. In a pre-hospital setting or a hospital admittance department, the desire might be to determine whether abnormal ICP is present.
This may be useful for triage and the need to determine further treatment. Evidence of very high ICP in wounded soldiers on the battlefield may be useful. However, the nICP equipment may not be very accurate in these situations. For example, the goal might be to determine whether ICP is above a certain level e. For this particular purpose, the technique must be easy to operate, user-independent, practically risk-free, and independent of the operative environment. For a new method to be adopted in neuro-intensive care, it must provide accuracy at a comparable level to the invasive gold standard and allow for continuous assessments of ICP, both in the ICU and at the bedside.
Finally, assessment of individuals with chronic complaints e. The source signals for non-invasive monitoring of ICP used to date have shown limited success. We conclude that the currently available source signals are of limited value. Given that the comprehensive research on non-invasive ICP monitoring has not yet provided any technique that can be readily adopted in clinical practice, it is possible that current research should be shifted to smaller and more achievable goals.
Minimally invasive techniques such as extradural measurements provide pulsatile ICP readings similar to parenchymal measurements Fig. Therefore, sensor development with this in mind appears to be a technological advancement that can be incorporated more easily into clinical practice than completely non-invasive mean ICP estimation.
Sensor placement in the lumbar region rather than in the parenchyma will likely cause less immunologic response and tissue damage. The issues related to ICP source signal control and reference pressure variability are even more important when it comes to implantable ICP sensors.
For several neurological and neurosurgical diseases, improved patient care would include long term monitoring of ICP and ICC on timescales of weeks to months.
The invasive catheter systems can only be used in the hospital setting. Measurements outside the hospital require implantable systems. Telemetric devices have therefore been proposed as an alternative to the traditional invasive cable bound techniques and are often mentioned as an intermediate step towards completely non-invasive long-term monitoring.
Many development attempts based on this proposal have been made throughout the past few decades [ , , , ]. Currently, there are two commercially available telemetric ICP sensors on the market [ 31 ]. They are both strain gauge micro transducers extended into the intracranial space through a burr hole.
This communicates with an external telemetric reader connected to a portable data logger that displays the mean ICP scores [ , ]. The device allows for storage of ICP recordings in a short- and a long-play mode.
The short-play mode stores five mean ICP values per second 5 Hz. The long-play mode stores one ICP value per second 1 Hz ; hence, long-term pulsatile monitoring was not an option.
Some challenges remain, however. The chosen sampling frequency of 5 Hz is too low for adequate ICP waveform analysis [ ], but sufficient for calculation of the PRx [ 30 , ]. In addition, the device only allows for 72 h of monitoring in the short-play mode, and the data processing software also requires improvement [ ]. The loss of signal and disruption of measurements due to telemetric reader misalignment has also been reported [ 30 ].
This system also consists of an implanted sensor and an external reader unit, but the sensor is integrated into an existing shunt system. This system has the added benefit of therapeutic shunt drainage and will reveal over or under drainage of CSF. However, long term monitoring of ICP and accurate pulsatile ICP readings are not feasible with the current device [ 31 , ]. Major challenges concerning implantable ICP micro sensors include biocompatible electronics, power sources and efficient telemetry, in addition to the issue of reference pressure drift [ ].
A major issue is how zero reference pressures are affected, and how to control whether this is the case or not. It seems clear that the mean ICP values provided by the telemetric devices must be interpreted with caution. Since telemetric devices are implanted, the risks of bleeds and infection are the same as for other invasive ICP measurement devices.
One limitation of the implantable devices described in the previous section is the need for surgical retrieval procedures, which subjects patients to the distress and risks associated with re-operation. Biodegradable sensors have therefore been explored in some detail, as they could provide added benefits because the complications are avoided. There are currently no commercially available products in this category, but technological advances have been made over the past years.
Kang et al. The pressure sensor and cables are entirely biodegradable when immersed in aqueous solutions such as CSF. The data transmitter with the potentiostat is placed under the skin subdermally and is not biodegradable, but because the wires through the skull are biodegradable, this can easily be removed once the pressure sensor and cables have degraded.
Stable, continuous operation has been proven for up to 3 days in vivo in rats. The pressure sensor volume is 0. However, very little measurement data has been provided so far. Extending the operational lifetime of these devices is a daunting challenge in material science as the end goal is still for the sensor to be naturally eliminated and dissolved into biologically safe products [ , ].
The use of ultrathin films of silicon dioxide t-SiO 2 as encapsulation layers has been suggested as a potential solution and has been applied to biodegradable ICP sensors to increase the operational lifetime.
On day 25, a negative drift of 4 mmHg was reported, comparable to clinical ICP monitors. After this, the signals from the device disappear, possibly due to dissolution of the sensors [ ].
This field is still in its early stages, and in vivo biodegradable sensors are emerging as a powerful tool in biomedical research that has significant potential in diagnostic medicine. Clinical utility of biodegradable miniature sensors may have greater promise.
While this field is also in its early phase, the focus of clinical neuroscientists should shift towards this aspect, rather than focusing on nICP approaches. There are numerous challenges related to more permanent implantable pressure sensors that the field has yet to overcome. The implantable device must operate under hostile conditions.
The environment within the human body is humid, filled with proteins, enzymes and ions. In addition, in vivo ICP measurements are especially challenging, because the pressure-sensitive part of the sensor must be in physical contact with the medium in which the pressure detection is being performed. Combined efforts between the scientific fields, addressing materials, biology and medicine, are therefore the most important goal going forward [ ].
What are the big questions that need to be answered in this area in the coming ten or 20 years? Where should clinicians, physicists and engineers invest their time and effort to develop new technologies?
How can one bring neuromonitoring into the 21st century? From our engineering and clinical perspective, we have highlighted some areas that we regard as most important.
As in other fields of medicine, progress in this field is highly technology-driven. The areas for improvement presented here require technology development to lift ICP measurement practice to another level. Given that mean ICP is the most prevalent ICP score independent of measurement modality, more focus is needed on the fact that this score can be affected by the variability of reference pressure.
There is a very small risk of bleeding in the brain during placement of the probe but this is rare. Once the monitoring is complete, the surgeon removes the sensor at the bedside and the skin is sutured back into place. Your doctor will carefully analyze the information obtained from this procedure and consider it along with data from other tests, such as eye examination, shunt studies, MRI or CT scans. This process can take up to a week. Complex cases need to be discussed at the weekly case conference to provide the best course of treatment.
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The infection, meningitis , is serious, but treatable with a day course of intravenous antibiotics. Leg or back pain : Occasionally, as the spinal needle or catheter is entering the lumbar space, it may touch a nerve root, causing a brief pain shooting down the back or leg. Moving the body to a different position usually alleviates the pain, and there is rarely any lasting injury to the nerve root.
Headache : During the procedure, some people develop a headache that gets worse when sitting or standing. You may also experience nausea, dizziness or lightheadedness. Stopping the fluid drain for a few minutes will usually relieve these symptoms.
A similar headache can occur after the tube is removed. You should drink plenty of caffeine containing fluids and take ibuprofen or acetaminophen as needed. If this kind of headache persists, a blood patch may be applied to the area to seal the fluid leak.
Bleeding : Bleeding is minimal during the procedure. There is a potential for bleeding around the brain if excess fluid is drained. To decrease the chance of this complication, please notify the nurse when you change positions or have to use the bathroom.
Spinal fluid leak : Rarely, spinal fluid may leak around the catheter and its dressing. If it leaks onto a non-sterile area, the physician may decide to remove the catheter. Depending on how much you have drained, a new one may be put in, or multiple spinal taps may be done to drain more spinal fluid.
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