Clinical Measurement and Monitoring
The ability to measure and monitor the physiology of patients is fundamental to modern anaesthesia. The anaesthetist is responsible for the correct use of sophisticated instruments which extend clinical observations beyond the human senses and enhance patient care. This requires vigilance and awareness of the limitations of the processes of measurement and the many causes of error. Uncritical acceptance of the recordings of monitoring equipment in the face of contradictory evidence is a common mistake. Unreliable measurements that are taken at face value and used to change patient management compromise the safety and effectiveness of care. It is essential that those who use monitors understand their limitations and are able to justify their risks.
Feasibility of measurement. The sensitivity and inherent variability of a clinical measurement depend on complex interactions and technical difficulties at the biological interface between the patient and the instrument.
Reliability of measurements is determined by the properties of the measurement system. This is influenced by the calibration and correct use of the instrument. Simple examples include the correct placement of ECG electrodes, or the appropriate size of cuff for non-invasive measurement of arterial pressure. Delicate equipment, e.g. a blood gas analyser, requires regular maintenance and calibration.
Interpretation depends on the critical faculties of the anaesthetist who interprets the significance of measurements in the context of complex physiological systems. Arterial pressure may be within the normal range despite severe hypovolaemia or derangement of cardiovascular function within the limits of physiological compensation. Global measurements of end-tidal carbon dioxide tension or oxygen saturation are influenced by many factors in a highly complex system. More information is required to deduce the cause of a change in the measurement.
Value of clinical measurements in patient care is defined by the role of a measurement in improving patient care. This includes the ease, convenience, continuity and usefulness of a clinical measurement, and evidence of improvement in patient safety and clinical outcome.
Monitoring is the process by which clinical measurements are assessed and used to direct therapy. In general, monitors consist of four components (Table 16.1): (1) a device which connects to the patient – this may either be a direct attachment or via a tube or lead; (2) a measuring device, often a transducer which converts the properties of the patient into an electrical signal; (3) a computer which may amplify the signal, filter it and integrate it with other variables to produce a variety of derived variables; (4) a display which may show the results as a wave, a number or a combination. It is important to appreciate that most monitors do not directly measure the displayed variable, and that the displayed variable may not reflect physiological function. For example, an electrocardiograph (ECG) does not measure cardiac function and therefore a normal ECG trace does not guarantee that the heart is pumping effectively. When interpreting measurements, the following questions should be asked:
Connection to patient
What is being measured? In the case of arterial pressure, there is an obvious answer. However, in some cases, for example ‘depth of anaesthesia’, it may not be clear what the monitor is measuring. In addition, many monitors use data from a variety of sources. For example, heart rate is usually derived from the ECG. However, if the ECG fails to provide the data required, the monitor often switches automatically to a rate from either a pulse oximeter or an arterial pressure waveform. Thus, the displayed value may change rapidly despite the patient remaining stable.
How is it measured? Arterial pressure is often measured by either a transducer attached to an arterial cannula or an automated oscillometer. Although a transducer is often regarded as the more accurate, the readings must be compared with the preoperative values recorded on the ward, usually with an oscillometer. Therefore, where accurate control of arterial pressure is essential, it is advisable to start invasive pressure monitoring before anaesthesia to avoid any confusion with non-invasive measures.
Is the environment appropriate? Many monitors have been designed for use in operating theatres and do not function correctly if exposed to the cold and vibration, for example in an ambulance or helicopter. Another example is the strong magnetic field produced by magnetic resonance imaging (MRI) scanners. The electrical currents induced may damage not only incompatible monitors but even produce burns to a patient’s skin.
Is the patient appropriate? Monitors designed for adult use often fail to produce reliable readings when used on small children. Particularly obese adults may require a large blood pressure cuff, and poor-quality ECG readings may be obtained.
Has the monitor been applied to the correct part of the patient? For example, in aortic coarctation, arterial pressure may be markedly different in each arm. Pulse oximeters also fail to work reliably if placed on a limb distal to a blood pressure cuff.
Is the variable within the range of the monitor? Most monitors are validated on healthy patients in laboratories. Whether such monitors continue to provide accurate results during the extreme physiological changes of, for example, anaphylaxis is uncertain. This does imply that the usefulness of monitors declines with the health of the patient: that is, they are least reliable when needed most. In most cases of acute perioperative patient deterioration, additional monitoring is needed.
Has the monitor been checked, serviced and calibrated at the correct intervals? To reduce costs, departments may re-use single-use equipment and fail to ensure that service checks are carried out. All equipment should be tagged with a service sticker. This should identify the date serviced, when the next service is due and who to contact in case of malfunction. Equipment which has not been serviced or is past its service date should not be used.
What is being measured?
What method is being used?
Has the monitor been serviced and calibrated?
Is the environment appropriate?
Is the patient appropriate?
Is it attached to the appropriate part of the patient?
Is the range appropriate?
Can the display be read?
Are the alarms on and have the limits been set?
Mechanical instruments use the signal energy to drive a display, with minimal intermediate processing. The height of a fluid-column manometer provides a visible index of pressure. The expansion of mercury within the confines of a thin glass column is a measure of temperature. Mechanical springs and gearing translate the rotation of a vane into the recording of expired volume on a dial. However, the overwhelming trend is for nonelectrical signals to be converted by a transducer to an electrical signal suitable for electronic processing by digital computers.
The development of digital microprocessors over the last 25 years has revolutionized anaesthetic practice. Beautifully engineered mechanical instruments, e.g. the von Recklinghausen oscillotonometer, are now obsolete in developed countries.
All clinical measurement systems detect a biological signal and reproduce this input signal in the form of a display or record which is presented to the operator. The degree to which a discrete measurement is a true reflection of the underlying signal is defined by its accuracy and precision.
Accuracy is the difference between the measurements and the real biological signal, or in practice, a different and superior ‘gold standard’ measurement. Calibration against predetermined signals is used to test and optimally adjust measuring instruments. For absolute measurements, e.g. arterial pressure, one point must be a fixed reference or ‘zero’.
Precision describes the reproducibility of repeated measurements of the same biological signal. This dispersion is usually described by summary statistics, standard deviation for normally distributed measurements, or the range for non-normal distributions. A single recording is unreliable when the measurement is imprecise. This is especially true of tests which require patient cooperation, practised skill or effort, e.g. peak expiratory flow rate. Repeated measurements demonstrate the variability in response.
The anaesthetist must be satisfied with the accuracy and precision of any clinical measurement used in patient management. Repeated measurements which are consistent ensure that the measurement is representative, i.e. precise, but do not ensure accuracy. For example, repeated recordings of invasive arterial pressure may be extremely consistent, but erroneous if the transducer is not calibrated against the correct zero point. Defences against the uncritical acceptance of inaccurate measurements include meticulous care in calibrating instruments and recording of clinical measurements, and reflection on clinical measurements which do not fit the clinical state of the patient or other related measurements. A discrepant result should be rechecked, using a different measurement technique if possible, before it is used to change patient management. This is especially true of complex, operator-dependent techniques such as measurement of cardiac output.
Continuous signals, which include the majority of modern clinical measurements such as biological electrical signals and the electrical output of signal transducers, introduce the complication of the response of the measuring instrument to a changing signal over time. The reliability with which a continuous signal is reproduced is defined by the relationship between input and output of the measurement system over the clinical range of signal magnitude and frequency. The input–output function of an accurate clinical measurement system would demonstrate good zero and gain stability, minimal amplitude non-linearity and hysteresis, and an adequate frequency response. This cannot be taken for granted, particularly with older equipment or with variations in environmental temperature or humidity.
The ability of a measurement to maintain a zero reading on the display or record when the input signal is zero defines the zero stability. The importance of zero instability depends on the magnitude relative to the signal, e.g. a zero drift of a few millimetres of mercury is much less important for the measurement of arterial pressure than it is for intracranial pressure.
The degree of amplification of the signal should be constant over the whole range of signal amplitudes. Manufacturers usually specify the degree of linearity of electronic components over a certain amplitude range. The amplitude linearity of a complete clinical system may be confirmed easily in an electronics laboratory by comparing the output to known, standardized test signals.
Some instruments, such as thermistors and humidity sensors, may display hysteresis. This is a special case of non-linearity, in which the output differs depending on whether the input signal is increasing or decreasing.
Many biological signals vary in a complicated and rapidly changing pattern. Accurate reproduction of a complex waveform requires that all of the component frequencies which make up the waveform are processed in an identical manner. This requires more than simply equal amplification irrespective of frequency, i.e. no amplitude distortion. It also implies that the relative positions of the various frequency components of the waveform are not shifted, i.e. no phase distortion. In practice, accurate reproduction up to the 10th harmonic of the fundamental frequency is sufficient for clinical purposes, e.g. 30 Hz for an arterial pressure waveform associated with heart rates up to 180 beats min–1 (3 Hz).
Biological signals are obscured to a variable degree by unwanted or extraneous signals which have similar physical characteristics and are described as noise, e.g. heart sounds become difficult to detect in the presence of continuous, noisy breath sounds. The efficiency of isolation of the signal from unwanted biological signals and electronic noise sources in the equipment is defined by the signal-to-noise ratio. The variability of the amplitudes of signal and noise is enormous and the signal-to-noise ratio is described using a logarithmic scale of decibels. Microvolt EEG measurements are particularly susceptible to noise from many sources. Biological noise includes contaminating ECG and EMG potentials, particularly from the scalp muscles, and interference from electrochemical activity at the skin–electrode interface. Electrostatic and electromagnetic linkage between the recording wires and nearby sources of mains electricity generates noise which is predominantly 50 Hz frequency and harmonics. Radiofrequency noise from diathermy or transmitters may also be picked up at this stage. Physical disturbance of the recording wires causes tiny changes in capacitive potentials and may add low-frequency noise, called microphony. Thermal noise is added during amplification, particularly at the input stage when the signal is in the microvolt range. Good amplifier design, electronic filtering of unwanted frequencies and modern techniques of digital signal processing may extract small signals from considerable background noise, but this inevitably introduces some distortion of the signal. Prevention of contamination of the signal by minimizing sources of noise before the signal is amplified is always preferable. The operator is responsible for correctly using measuring instruments to optimize the signal and for applying knowledge of the physical principles of the measurement to minimize contamination by noise.
Measuring instruments based on mechanical principles lack the flexibility and automated control of computerized devices, but use ingenious methods for processing and displaying analogue measurements. For example, mechanical spirometers use precision-engineered gears to translate the movement of a piston or vane into the rotation of a calibrated dial.
Analogue computers use hardware comprising electronic circuits and operational amplifiers. Signals are processed in the form of continuously variable electric potentials. Analogue hardware components continuously perform a wide variety of mathematical functions on a rapidly changing input waveform. Integration and differentiation are formidable mathematical tasks for a digital computer, which can be solved simply and cheaply using analogue circuits comprising capacitors and resistors. Integration of the flow signal from a pneumotachograph produces a volume waveform.
Digital signal processing offers a powerful alternative to mechanical processing and analogue computation. A fundamental step in this process is the conversion of a continuous analogue electrical signal into a discrete digital form. This analogue-to-digital conversion is achieved by measuring or ‘sampling’ the continuous input signal at regular intervals, to produce a series of discrete measurements over time which are in a suitable format for digital computation. The overwhelming advantage of digital processing is that the manipulation of the digitized signal is performed by a flexible and unlimited series of software calculations which range from mathematical functions to the analysis of statistical properties and trends.
The core processing units of digital computers assume one of two stable states, i.e. a binary, rather than decimal, code. This imposes a limit on the resolving power of the digital processor, although with increasing processing power this limit has become negligible. Earlier 8-bit computing comprised a binary number of eight digits representing 28 integer decimal numbers, from binary 00000000 = decimal 0 to binary 11111111 = decimal 255. In short, an 8-bit converter can resolve an analogue signal with an accuracy of one part in 255, i.e. with an amplitude resolution of 0.4% of full scale. A 12-bit converter is more accurate, with a resolving power of one part in 4095 or 0.02% of full scale. More modern processors are capable of 32-bit computing corresponding to a range of 232 integer decimal numbers (a resolving power of 0.00000002%), with 64-bit processors now commonly found in domestic computer equipment. Whilst highly accurate, the cost of this improvement in resolution is more expensive hardware to digitize, process and store considerably more digital information.
Amplitude resolution is not the only determinant of the accuracy of analogue-to-digital conversion. Resolution over time, determined by the sampling frequency, is also important. A relatively low sampling frequency may provide a representative sample of values for a slowly changing waveform but it may inadequately represent high-frequency components and introduce an aliasing error, in which different signals become indistinguishable. The Nyquist theorem suggests that the minimum sampling frequency to maintain the integrity of the waveform is at least twice the highest frequency component with significant amplitude in the input signal waveform, e.g. a sampling frequency of 100 Hz would adequately capture the fastest rate of change in a physiological pressure signal.
The immensely powerful and complicated hardware and software programming instructions responsible for performing the tasks of digitizing, processing, storing and displaying the input signal are hidden from view in the commercial ‘black box’.
A continuously variable signal, such as pressure or temperature, is represented by an analogue display in terms of the amplitude of a physical quantity on a calibrated scale, dial, electrical meter or printed record. The glass thermometer incorporates a wedge-shaped lens which magnifies the appearance of the mercury column against the calibrated background scale. The height of a water column manometer is a linear, visual scale of pressure. Simple mechanical displays are accurate and easily understood, but are inconvenient to read and most suitable for intermittent discrete measurements.
Mechanical spirometers and flowmeters record flow on a dial driven by gears. Electrical moving coil meters use a coil of wire suspended in a magnetic field which rotates in proportion to the applied current and moves a pointer on a calibrated dial. Alternatively, the amplified and filtered electrical signal could drive a chart recorder which produces a continuous printed record of the amplitude of measurements against time. Limitations common to these mechanical devices include fragile moving parts, and inertia which impairs the frequency response to rapidly changing signals.
The cathode ray oscilloscope is an effective screen-based display for continuous analogue electrical signals. A heated cathode generates a stream of electrons which are focused and accelerated onto a phosphorescent coating which lines the flat surface of the tube to generate a bright spot. The position of the electron beam in both x- and y-axes is controlled by electrostatic plates. The continuously varying input signal is applied to the y-plates so that deflection in the vertical y-axis is proportional to the amplitude of the signal. The absence of mechanical parts results in a high-frequency response. An electronic time-base circuit delivers a saw-tooth voltage to the x-plates which drives the electron beam across the x-axis at a constant rate and returns the beam to the left-hand side at the start of each sweep. This produces a dynamic image of signal amplitude against time. Alternatively, a second input signal may be applied to the x-plates to produce an x-y graphical plot, e.g. pressure–volume loop. Cathode ray oscilloscopes are widely used in electronic engineering and signal processing, but have been replaced in clinical practice by microprocessor-controlled displays.
Digital signal processing has revolutionized clinical measurement. Modern monitors comprise a single system integrating various measurements of physiology, and display information as discrete measurements as well as continuous analogue waveforms (Fig. 16.1). This paradox, the conversion of analogue information into digital and then back to analogue, illustrates the real power of digital signal processing to manipulate and present information in a relevant and user-friendly manner. Data patterns (waveforms, trends, graphs) can be recreated or processed in other ways from the original digital signal to assist the anaesthetist, with the original digital signal being stored in a computer record without degradation of the quality of the signal.
Most of the current monitoring systems follow good ergonomic principles, with different variables separated consistently by position on the screen and by colour. This allows the most important information to be placed centrally in large symbols or fonts and in bold colours, with less important data either relegated to small print, or placed in submenus. However, the flexibility of most monitors implies that it is still possible for individuals to change colours and priorities, often making the monitor much less effective. Whenever possible, departments should ensure that all monitors have identical default settings to reduce confusion (these are usually password protected). Unfortunately, the lack of international standards means that confusion may still occur if monitors from multiple sources are used in the same unit.
Despite many attempts to simplify patient data into geometric shapes or bar graphs, data continue to be displayed most often as simple numbers, supported by waveforms, e.g. invasive pressure, and a graphical display of trends over time. Trends are particularly useful when clinical problems may produce gradual change. For example, in neurosurgery, a gradual decrease in end-tidal carbon dioxide concentration is often associated with multiple air emboli.
Depolarization of the cell membrane of excitable cells is fundamental to the action of these cells and generates a transient potential difference between the active cell and surrounding tissues. The summation of synchronous extracellular potentials from a large number of excitable cells generates a widespread electric field which can be detected by electrodes on the body surface. The electrocardiogram (ECG) and electroencephalogram (EEG) are two well-established measures of biological electrical activity.
Biological electrical signals are detected using electrodes constructed of silver and electrolytically coated with silver chloride. Low, stable impedances minimize mains interference. Symmetrical electrode impedance and insignificant polarization control drift. However, care is still required to achieve optimum results. The silver chloride layer is very thin, prone to deterioration and only suitable for single use. Movement artefacts which alter the electrode potential and impedance are greatly reduced if the electrode surface is separated from the skin by a foam pad impregnated with electrolyte gel. It is no longer necessary to abrade the skin to achieve ultra-low impedance, but de-greasing with alcohol before applying the electrode helps to reduce skin impedance and ensures satisfactory adhesion.
The amplitude of tiny bioelectrical signals must be increased by amplification, and unwanted noise and interference minimized. Calibration voltages may be incorporated for correct adjustment of the gain of the amplifier.
Amplifiers for biological signals require high common mode rejection and high input impedance. The input and electrode impedances act as a potential divider: high electrode impedance and low amplifier input impedance attenuate the electrical signal across the amplifier. The input impedance of modern amplifiers exceeds 5 MΩ to avoid problems, and careful attention must be paid to minimizing electrode impedance, particularly for EEG recordings.
Differential amplification is a powerful method of reducing unwanted noise. The potential difference between two input signals is amplified, but electrical signals common to both are attenuated. This feature is termed ‘common mode rejection’ and very effectively reduces mains interference in all biological signals and electrocardiographic contamination of much smaller electroencephalographic signals. The common mode rejection ratio (CMRR) for a typical differential amplifier exceeds 10 000:1. In other words, a signal applied equally to both input terminals would need to be 10 000 times larger than a signal applied between them for the same change in output.
The bandwidth of the amplifier must cover the range of frequencies which are important in the signal. In practice, amplifiers require a flat frequency response for ECG from 0.14 to 50 Hz, for EEG from 0.5 to 100 Hz and for EMG from 20 Hz to at least 2 kHz.
Low-frequency interference, largely caused by slow fluctuating potentials generated in the electrodes, produces baseline instability and drift. This is removed by incorporating a network of resistors and capacitors which function as a simple high-pass filter allowing biological signals to pass, but attenuating low-frequency noise. This introduces a compromise in amplifier design between signal trace fidelity and stability of recording. For example, amplifiers designed for diagnostic electrocardiography have long time constants with optimal reproduction of the waveform at the expense of baseline instability, especially to movement. In comparison, continuity of recording is more important when the electrocardiogram is used for monitoring during anaesthesia; high-pass filtering produces a short time constant and good baseline stability at the expense of waveform reproduction. Low-frequency elements of the ECG, such as the T wave, may become differentiated by phase shift in the high-pass filter and appear distorted or biphasic.
Other filters can attenuate particular frequencies. Highly selective band reject filters attenuate 50 Hz interference from the signal. Low-pass filters are used to eliminate higher-frequency artefacts from an EEG signal. The purpose of filtering is to reduce unwanted noise relative to the signal. When the frequency range of signal and noise overlap, some degree of signal degradation is inevitable.
Noise Originating from the Patient: Millivolt ECG potentials on the body surface are hundreds of times larger than microvolt EEG signals on the scalp. EMG signals may be even larger, and muscular activity, especially shivering, causes severe interference. Two features of electronic amplifier design substantially improve the EEG signal-to-noise ratio. ECG potentials are essentially the same across the scalp and are ignored by amplifiers with a high common mode rejection. EMG activity has a higher frequency content than the EEG signal, and may be minimized by a low-pass filter which attenuates the higher-frequency response of the amplifier to a level which attenuates the EMG signals and does not interfere with the characteristics of the EEG.
Noise Originating from the Patient–Electrode Interface: Recording electrodes do not behave as passive conductors. All skin–metal electrode systems employ a metal surface in contact with an electrolyte solution. Polarization describes the interaction between metal and electrolyte which generates a small electrical gradient. Electrodes comprising metal plated with one of its own salts, e.g. silver—silver chloride, avoid this problem because current in each direction does not significantly change the electrolyte composition. Mechanical movement of recording electrodes may also cause significant potential gradients – alteration in the physical dimensions of the electrode changes the cell potential and skin–electrode impedance. Differences in potential between two electrodes connected to a differential amplifier are amplified and asymmetry of electrode impedance seriously impairs the common mode rejection ratio of the recording amplifier.
Noise Originating Outside the Patient: Electrical interference. Mains frequency interference with the recording of biological potentials may be troublesome, particularly in electromagnetically noisy clinical environments. Patients function physically as large unscreened conductors and interact with nearby electrical sources through the processes of capacitive coupling and electromagnetic induction.
Capacitance permits alternating current to pass across an air gap. A live mains conductor and nearby patient behave as the two plates of a capacitor. The very small mains frequency current which flows through the patient is of no clinical significance but confounds the detection and amplification of biological potentials, creating unwanted interference in the recording. Capacitive coupled interference is minimized by reducing the capacitance and the alternating potential difference. This is achieved by moving the patient away from the source of interference and by screening mains-powered equipment with a conductive surround which is maintained at earth potential by a low-resistance earth connection and by surrounding leads with a braided copper screen – stray capacitances couple with the screen instead of the lead.
Alternating currents in a conductor generate a magnetic flux. This induces voltages in any nearby conductors which lie in the changing magnetic flux, including the patient or signal leads to the amplifier, which function as inefficient secondary transformers. This source of interference is minimized by keeping patients as far as possible from powerful sources of electromagnetic flux, especially mains transformers. Electromagnetic inductance may be minimized by ensuring that all patient leads are the same length, closely bound or twisted together until very close to the electrodes. This ensures that the induced signals are identical in all leads and therefore susceptible to common mode rejection.
The importance of low electrode impedance. Low electrode impedance may exaggerate the effects of surrounding electrical interference. Capacitive and inductive coupling produce very small currents in the recording leads. If the electrode impedance is low, the potential at the amplifier input must remain close to the potential at the skin surface, so that minimal interference results. If electrode impedance is high, the small induced currents may create a significant potential difference across that impedance, leading to severe 50 Hz interference.
Radiofrequency interference from diathermy is a severe problem for the recording of biological potentials. ECG amplifiers may be provided with some protection by filtering the signal before it enters the isolated input circuit, filtering the power supply to block mains-borne radiofrequencies and enclosing the electronic components in a double screen, the outer earthed and the inner at amplifier potential.
Pressure is a mechanical signal fundamental to measurement and monitoring in anaesthesia. Several physical principles and a wide range of instruments are used to measure pressure. Liquid column manometers display pressure according to the height of a column of fluid relative to a predefined zero-point, and the density of the fluid. Mechanical pressure gauges are used widely, particularly in high-pressure gas supplies; pressure-dependent mechanical movement is amplified by a gearing mechanism which drives a pointer across a scale.
For most physiological pressure measurements, diaphragm gauges are used – a flexible diaphragm moves according to the applied pressure. Mechanical display of diaphragm movement is limited by poor sensitivity to small pressures, inertia to changing pressure and a narrow range of linear response. In modern diaphragm gauges used for sensing dynamic pressures, movement of the diaphragm is sensed by a device which converts the mechanical energy imparted to the diaphragm into electrical energy.
The first step in transduction is movement of the diaphragm caused by the relationship to applied pressure. This depends on the stiffness of the diaphragm and substantially determines the operating characteristics of the transducer. Linearity of amplitude and frequency response are improved by using small stiff diaphragms which require a more sensitive mechanism for sensing diaphragm movement.
Wire strain gauges are based on the principle that stretching or compression of a wire changes the electrical resistance. Changes in capacitance or inductance have also been coupled to movement of a diaphragm. Silicon strain gauges use the changes in resistance in a thin slice of silicon crystal which occur when it is compressed or expanded. They are very sensitive and suitable for incorporation into a small stiff diaphragm with excellent frequency response, but non-linearity and temperature dependence are difficult technical problems.
Optical transduction senses movement of the diaphragm by reflecting light from the silvered back of the convex diaphragm on to a photocell. Applied pressure causes the silvered surface to become more convex. This causes the reflected light beam to diverge, reducing the intensity of reflected light sensed by the photoelectric cell. This design is used in a fibreoptic cardiac catheter for intravascular pressure measurement. These miniature pressure transducers are expensive but have a high-frequency response and fibreoptic light sources eliminate the risk of microshock.
The principal aim of an anaesthetist is to ensure the delivery of oxygen to the patient’s tissues. In physiological terms, oxygen delivery is the product of the cardiac output, the concentration of haemoglobin and its oxygen saturation. Clinically, if the patient is pink, with a normal volume pulse and has warm extremities, then these aims are being met. When combined with a urine output of greater than 0.5 mL h–1 it is unlikely that the patient has any cardiovascular problems. A further confirmatory test, especially useful in children, is the capillary refill time. When an extremity is compressed for 5 s, if capillary refill occurs in less than 1.5 s, cardiac output is adequate. If the refill time is greater than 5 s, then shock is likely to be present.
The need for direct patient observation cannot be overestimated. Literally having a ‘finger on the pulse’ and being able to see the patient are the most important safety factors. While factors such as drapes and dimmed theatre lights may make direct observation difficult, there should not be complete reliance on electronic monitoring.
The electrocardiogram (ECG) is a well-established measure of myocardial electrical activity. The synchronous depolarization and prolonged action potentials in cardiac muscle summate to generate a potential field of high amplitude. This potential difference is detected between two electrodes placed on the body surface. In the three-lead system commonly in use, the third lead is used as a reference electrode. The voltage changes are very small (1 mV in amplitude with a frequency response of 0.05–100 Hz) and require amplification before being displayed as the familiar waveform.
Different lead positions detect electrical activity from different parts of the myocardium. The commonest position of the electrodes used in the operating theatre is the CM5 arrangement, as this is the best position to detect ischaemia of the left ventricle (Fig. 16.2).
The ECG is a standard monitor used on all anaesthetized patients. The visible waveform allows the cardiac rhythm to be identified and may often be printed for further analysis. Alarms may be set to identify arrhythmias and brady/tachycardias. Many monitors are able to display a numerical value of the heart rate in addition to a measure of any ST segment depression/elevation produced by cardiac ischaemia/infarction. This may be displayed as a trend over time and the success of treatment observed.
Unfortunately, the relatively small voltages measured are easily swamped by skeletal muscle activity or surgical diathermy, often leading to false alarms. The signal may also be severely degraded if the gel of the electrodes has been allowed to dry out or if the weight of the leads is allowed to pull on the electrodes. In addition, the monitor only identifies ischaemia in a single area; multiple lead systems are required to monitor the whole myocardium. While the ECG has become a standard monitor, it adds little to the information provided by palpating the pulse. It must be remembered that electrical activity does not always produce a cardiac output. Complications are rare, although the electrode adhesive may produce skin damage in susceptible patients.
An adequate arterial pressure is essential for tissue perfusion; even when perfusion is adequate, hypotension may lead to renal failure. Indirect methods of measuring arterial pressure do not depend on contact between arterial blood and the system for signal recognition and transduction. Arterial pressure may be most rapidly estimated by palpating a pulse, although this method is too unreliable as a single technique. The majority depend on signals generated by the occlusion of a major artery using a cuff, known as the Riva-Rocci method. Systolic pressure can be estimated by the return of a palpable distal pulse; auscultation of the Korotkoff sounds can determine systolic and diastolic pressures. These methods, however, are too time-consuming during anaesthesia and often impossible because of poor access to the patient’s arm.
The oscillometric measurement of arterial pressure estimates arterial pressure by analysis of the pressure oscillations which are produced in an occluding cuff by pulsatile blood flow in the underlying artery during deflation of the cuff (Fig. 16.3). The original automatic oscillometers used two cuffs. The upper cuff was inflated to occlude the arterial flow and then gradually deflated. As the blood flow began to pass under the upper cuff, the small changes in volume were detected by the lower cuff with an electromechanical pressure transducer. Modern machines use a single cuff with two tubes for inflation/measurement. During slow deflation, each pulse wave produces a pressure transient in the cuff which may be distinguished from the slowly decreasing ambient pressure in the cuff. Above systolic pressure, the transients are small, but suddenly increase in magnitude when the cuff pressure reaches the systolic point. As the cuff pressure decreases further, the amplitude reaches a peak and then starts to diminish. The mean arterial pressure correlates closely with the lowest cuff pressure at which the maximum amplitude is maintained. As the cuff pressure reaches diastolic pressure, the transients abruptly diminish in amplitude. To avoid high cuff pressures and long deflation times, monitors inflate the cuff to just above a normal systolic pressure and then slowly decrease the pressure until a pulse is detected. Consequently, estimates of diastolic pressure can be unreliable. If a pulse is not detected, the cuff is then inflated to a higher pressure. This process may be repeated several times before a measurement is made.
FIGURE 16.3 Diagram showing: (A) Relationship between cuff pressure and intra-arterial pressure as cuff pressure decreases during oscillometry; (B) the signal created by the relative pressure changes in A. The sharp spikes of pressure in B are created by the walls of the artery opening and closing. These spikes are detected by a transducer first when the cuff pressure is just below systolic arterial pressure; their amplitude reaches a peak at mean arterial pressure and they cease when the cuff pressure is below diastolic pressure.
Commercial instruments incorporate mechanisms for improving the reliability of the measurement. For example, at each successive plateau pressure during the controlled deflation, successive pressure fluctuations are compared and accepted only if they are similar. All automatic oscillometric instruments require a regular cardiac cycle with no great differences between successive pulses. Accurate and consistent readings may be impossible in patients with an irregular rhythm, particularly atrial fibrillation.
Clinical studies comparing automatic oscillometric instruments with direct arterial pressure have demonstrated good correlation for systolic pressure with a tendency to overestimate at low pressures and underestimate at high pressures. Mean and diastolic pressures were less reliable. The 95% confidence interval for all three indices exceeded 15 mmHg. The disadvantages of automated oscillometry are shown in Table 16.3.
Delayed measurement with arrhythmias or patient movement
Inaccuracy with systolic pressure < 60 mmHg
Inaccurate if the wrong size cuff used
May be inaccurate in obese patients
Discomfort in awake patients
Skin and nerve damage in prolonged use
Delay in injected drugs reaching the circulation
Backflow of blood into i.v. cannulae
Pulse oximeter malfunction as cuff is inflated
Other Techniques: Measuring devices depend on the detection of movement of the arterial wall using changes in pressure or sound below audible frequencies and detection of blood flow using the Doppler shift of an ultrasound signal, or plethysmography. Several other techniques have been used to measure arterial pressure, but have failed to find widespread usage. These include the Penaz technique, which measures the effect of external pressure on the blood flow through a finger, and other devices relying on pressure measurements over an artery, Doppler probes or detection of Korotkoff sounds with a microphone.
To measure arterial pressure directly, a cannula (usually 20–22G parallel-sided Teflon) must first be inserted into an artery (usually the radial because occlusion of the artery may be compensated for by flow through the ulnar artery). As fluids are incompressible, the pressure in the artery is transmitted directly to a transducer, which converts pressure into an electrical signal which is displayed by the monitor. The cost and complexity of pressure transducers are compensated by convenience, accuracy, continuity of measurement and an electrical output which may be processed, stored and displayed according to the requirements. The transducer should be at the level of the left ventricle and the transducer opened to the atmosphere to provide a zero reading before use. Monitors usually display systolic and diastolic pressures as well as the mean pressure, calculated automatically by integrating the area under the pressure waveform. The waveform provides useful additional information: a rapidly appreciated estimate of pressure, a qualitative assessment of the adequacy of the frequency response and damping, and an assessment of relative hypovolaemia during positive pressure as identified by the variability or ‘swing’ in the waveform.
The advantage of direct measurements is a real-time measure of arterial pressure, which is essential when administering drugs such as vasopressors to critically ill patients (Table 16.4). Such measurement systems also provide a means for obtaining samples for arterial blood gas analysis and other blood tests. The use of arterial cannulae has therefore become standard practice for severely ill patients, both in the operating theatre and in the intensive care unit.
Accuracy of pressure measurement
Beat-by-beat observation of changes when blood pressure is variable or when vasoactive drugs are used
Accuracy at low pressures
Ability to obtain frequent blood samples
However, errors are common as a result of malpositioning of the transducers and failure to zero the transducer before use. For example, if the operating table is moved upwards while the transducer remains static, the difference in height artificially increases the pressure reading. Further, while modern disposable sets are usually reliable and accurate, they may occasionally malfunction. Unusual readings should therefore be checked against a reading from a non-invasive monitor. Complications relating to arterial cannulae are shown in Table 16.5.
Requires skill to insert
Pain on insertion
Arterial damage and thrombosis
Embolization of thrombus or air
Ischaemia to tissues distal to puncture site
Inadvertent injection of drugs
Late development of fistula or aneurysm
Resonant Frequency and Damping: Fourier showed that all complex waveforms may be described as a mixture of simple sine waves of varying amplitude, frequency and phase. These consist of a fundamental wave, in this case at the pulse frequency, and a series of harmonics. The lower harmonics tend to have the greatest amplitude and a reasonable approximation to the arterial pressure waveform may be obtained by accurate reproduction of the fundamental and first 10 harmonics. In other words, to reproduce an arterial waveform at 120 beats min–1 accurately would require transduction with a linear frequency response up to a frequency of at least (120 × 10)/60 = 20 Hz. Accurate reproduction of a waveform requires that both the amplitude and phase difference of each harmonic are faithfully reproduced. This requires a transduction system with a natural frequency higher than the significant frequency components of the system, and the correct amount of damping.
The fluid and diaphragm of the transducer constitute a mechanical system which oscillates in simple harmonic motion at the natural resonant frequency. This determines the frequency response of the measurement system (Fig. 16.4). The resonant frequency of a catheter-transducer measuring system is highest, and the frictional resistance to fluid flow which dampens the frequency response is lowest, when the velocity of movement of fluid in the catheter is minimized. This is achieved with a stiff, low-volume displacement diaphragm and a short, wide, rigid catheter.
Determination of the Resonant Frequency and Damping: The resonant frequency and the effects of damping may be estimated by applying a step change in pressure to the catheter-transducer system and recording the response (Fig. 16.5). The underdamped system responds rapidly but overshoots and oscillates close to the natural resonant frequency of the system; frequency components of the pressure wave close to the resonant frequency are exaggerated. By contrast, the overdamped system responds slowly and the recorded signal decreases slowly to reach the baseline, with no overshoot. High-frequency oscillations are damped, underestimating the true pressure changes. These extremes are undesirable.
Optimal Damping: Optimal damping maximizes the frequency response of the system, minimizes resonance and represents the best compromise between speed of response and accuracy of transduction. A small overshoot represents approximately 7% of the step change in pressure, with the pressure then following the arterial waveform (Fig. 16.5).
Damping is relatively unimportant when the frequencies being recorded are less than two-thirds of the natural frequency of the catheter-transducer system. Modern transducer systems using small compliance transducers connected to a short, stiff catheter, with a minimum of constrictions or connections, approximate to this ideal. The system also includes a pressurized bag of saline which produces a flow of 1–3 mL h–1 through a restrictor to prevent clot formation, as well as the facility to allow a higher flow rate to flush the system, for example after a blood sample has been taken. Air bubbles in the system, clotting or kinking in the vascular catheter, and arterial spasm lower the natural resonant frequency and increase the damping.
In clinical practice, the resonant frequency of the whole system is uncomfortably close to the frequency content of the signal, and accurate measurements require optimal damping. However, damping is difficult to measure and control, and is poor compensation for an inadequate frequency response in the pressure recording system. Adjustment of damping is difficult to achieve and mechanical methods which include inserting constrictions or a compliant tube into the system to increase damping further reduce the resonant frequency. Electronic damping of the electrical output from the transducer cannot correct for non-linear amplification and attenuation of frequencies in the pressure wave before transduction.
The accuracy of pressure measurements, particularly using indirect methods, needs to be considered. Invasive direct measurement of arterial pressure is the usual standard for comparison. However, the catheter-transducer system must be carefully set up and tested for optimal performance and this is hard to achieve in clinical practice. Arterial pressure varies throughout the arterial tree and the measured pressure depends on the site of measurement. As the pulse wave travels from the ventricle to peripheral arteries, changes in vessel diameter and elasticity affect the pressure waveform, which becomes narrower with increased amplitude. Differences in arterial pressure between limbs are common, particularly in patients with arterial disease.
Indirect methods using an occluding cuff make intermittent measurements, with the systolic and diastolic readings reflecting the conditions in the artery at two instants at which the end-points are detected. By contrast, direct pressure measurements are the average of a number of cycles, more precisely reflecting mean pressures. Indirect measurements may be compromised by taking a small number of infrequent samples from a variable signal.
Central venous pressure is often considered a measure of the amount of blood within the venous system; a pressure less than normal (2–3 mmHg) indicates hypovolaemia and a higher pressure indicates volume overload. While such a view is reliable for healthy patients with acute blood loss, it is not so simple in other circumstances. For example, patients with damage to the right side of the heart may have raised central venous pressure even when the filling pressure of the left side of the heart is low. Single measurements rarely provide an accurate reflection of the fluid status of the patient. However, repeated measurements taken while a fluid challenge is given can be informative.
Long catheters inserted via the antecubital fossa are relatively easy and safe to insert but are of small diameter. Catheters inserted via the basilic or cephalic vein are sometimes difficult to advance past the shoulder. It is also difficult to determine if the tip of the catheter is within a central vein without X-ray imaging. Thrombosis of the veins is common if the catheter is left in situ for more than 24 h.
Femoral venous catheters are inserted just below the inguinal ligament. They are also relatively easy to insert and may be of large gauge to allow rapid transfusion of fluids. This route is often chosen in children. However, the site of insertion is often within a skin fold, making skin sepsis more likely.
Internal jugular catheters are used most commonly because the vein is superficial, of larger diameter, and easily managed. This is the route which is often most appropriate for use in an emergency. However, the insertion point is adjacent to several vital structures, including the carotid artery, lung, brachial plexus and cervical spine, with the result that direct needle trauma to these structures can occur. Current guidelines recommend the use of an ultrasound probe for insertion of a catheter via the internal jugular route to improve the accuracy of insertion and to minimize complications.