There are a number of clinical methods to measure cardiac output, ranging from direct intracardiac catheterization to non-invasive measurement of the arterial pulse. Each method has advantages and drawbacks. Relative comparison is limited by the absence of a widely accepted "gold standard" measurement. Cardiac output can also be affected significantly by the phase of respiration – intra-thoracic pressure changes influence diastolic filling and therefore cardiac output. This is especially important during
mechanical ventilation, in which cardiac output can vary by up to 50% across a single respiratory cycle. Cardiac output should therefore be measured at evenly spaced points over a single cycle or averaged over several cycles. Invasive methods are well accepted, but there is increasing evidence that these methods are neither accurate nor effective in guiding therapy. Consequently, the focus on development of non-invasive methods is growing.
Doppler ultrasound This method uses
ultrasound and the
Doppler effect to measure cardiac output. The blood velocity through the heart causes a Doppler shift in the frequency of the returning ultrasound waves. This shift can then be used to calculate
flow velocity and volume, and effectively cardiac output, using the following equations: • Q = SV \times HR • SV = VTI \times CSA • CSA = \pi r^2 where: • CSA is the valve orifice cross sectional area, • r is the valve radius, and, • VTI is the
velocity time integral of the trace of the Doppler flow profile. Being non-invasive, accurate and inexpensive, Doppler ultrasound is a routine part of clinical ultrasound; it has high levels of reliability and reproducibility, and has been in clinical use since the 1960s.
Echocardiography Echocardiography is a non-invasive method of quantifying cardiac output using ultrasound. Two-dimensional (2D) ultrasound and Doppler measurements are used together to calculate cardiac output. 2D measurement of the diameter (d) of the aortic annulus allows calculation of the flow cross-sectional area (CSA), which is then multiplied by the VTI of the Doppler flow profile across the aortic valve to determine the flow volume per beat (
stroke volume, SV). The result is then multiplied by the heart rate (HR) to obtain cardiac output. Although used in clinical medicine, it has a wide test-retest variability. It is said to require extensive training and skill, but the exact steps needed to achieve clinically adequate precision have never been disclosed. 2D measurement of the aortic valve diameter is one source of noise; others are beat-to-beat variation in stroke volume and subtle differences in probe position. An alternative that is not necessarily more reproducible is the measurement of the pulmonary valve to calculate right-sided CO. Although it is in wide general use, the technique is time-consuming and is limited by the reproducibility of its component elements. In the manner used in clinical practice, precision of SV and CO is of the order of ±20%.
Transcutaneous Ultrasonic Cardiac Output Monitor (USCOM) uses
continuous wave Doppler to measure the Doppler flow profile VTI. It uses
anthropometry to calculate aortic and pulmonary valve diameters and CSAs, allowing right-sided and left-sided
Q measurements. In comparison to the echocardiographic method, USCOM significantly improves reproducibility and increases sensitivity of the detection of changes in flow. Real-time, automatic tracing of the Doppler flow profile allows beat-to-beat right-sided and left-sided
Q measurements, simplifying operation and reducing the time of acquisition compared to conventional echocardiography. USCOM has been validated from 0.12 L/min to 18.7 L/min in new-born babies, children and adults. The method can be applied with equal accuracy to patients of all ages for the development of physiologically rational haemodynamic protocols. USCOM is the only method of cardiac output measurement to have achieved equivalent accuracy to the implantable flow probe.
Transoesophageal The Transoesophageal Doppler includes two main technologies;
transoesophageal echocardiogram—which is primarily used for diagnostic purposes, and
oesophageal Doppler monitoring—which is primarily used for the clinical monitoring of cardiac output. The latter uses continuous wave Doppler to measure blood velocity in the
descending thoracic aorta. An ultrasound probe is inserted either orally or nasally into the oesophagus to mid-thoracic level, at which point the oesophagus lies alongside the descending
thoracic aorta. Because the transducer is close to the blood flow, the signal is clear. The probe may require re-focussing to ensure an optimal signal. This method has good validation, is widely used for fluid management during surgery with evidence for improved patient outcome, and has been recommended by the UK's National Institute for Health and Clinical Excellence (
NICE). Oesophageal Doppler monitoring measures the velocity of blood and not true
Q, therefore relies on a nomogram based on patient age, height and weight to convert the measured velocity into stroke volume and cardiac output. This method generally requires patient sedation and is accepted for use in both adults and children.
Pulse pressure methods Pulse pressure (PP) methods measure the pressure in an artery over time to derive a waveform and use this information to calculate cardiac performance. However, any measure from the artery includes changes in pressure associated with changes in arterial function, for example compliance and impedance. Physiological or therapeutic changes in vessel diameter are assumed to reflect changes in
Q. PP methods measure the combined performance of the heart and the blood vessels, thus limiting their application for measurement of
Q. This can be partially compensated for by intermittent calibration of the waveform to another
Q measurement method then monitoring the PP waveform. Ideally, the PP waveform should be calibrated on a beat-to-beat basis. There are invasive and non-invasive methods of measuring PP.
Finapres methodology In 1967, the Czech physiologist Jan Peňáz invented and patented the
volume clamp method of measuring continuous blood pressure. The principle of the volume clamp method is to dynamically provide equal pressures, on either side of an artery wall. By clamping the artery to a certain volume, inside pressure—intra-arterial pressure—balances outside pressure—finger cuff pressure. Peñáz decided the finger was the optimal site to apply this volume clamp method. The use of finger cuffs excludes the device from application in patients without
vasoconstriction, such as in sepsis or in patients on vasopressors. In 1978, scientists at BMI-TNO, the research unit of
Netherlands Organisation for Applied Scientific Research at the
University of Amsterdam, invented and patented a series of additional key elements that make the volume clamp work in clinical practice. These methods include the use of modulated infrared light in the optical system inside the sensor, the lightweight, easy-to-wrap finger cuff with
velcro fixation, a new pneumatic proportional control valve principle, and a set point strategy for the determining and tracking the correct volume at which to clamp the finger arteries—the Physiocal system. An acronym for physiological calibration of the finger arteries, this Physiocal tracker was found to be accurate, robust and reliable. The Finapres methodology was developed to use this information to calculate arterial pressure from finger cuff pressure data. A generalised algorithm to correct for the pressure level difference between the finger and brachial sites in patients was developed. This correction worked under all of the circumstances it was tested in—even when it was not designed for it—because it applied general physiological principles. This innovative brachial pressure waveform reconstruction method was first implemented in the Finometer, the successor of Finapres that BMI-TNO introduced to the market in 2000. The availability of a continuous, high-fidelity, calibrated blood pressure waveform opened up the perspective of beat-to-beat computation of integrated haemodynamics, based on two notions: pressure and flow are inter-related at each site in the arterial system by their so-called characteristic impedance. At the proximal aortic site, the 3-element
Windkessel model of this impedance can be modelled with sufficient accuracy in an individual patient with known age, gender, height and weight. According to comparisons of non-invasive peripheral vascular monitors, modest clinical utility is restricted to patients with normal and invariant circulation.
Invasive Invasive PP monitoring involves inserting a
manometer pressure sensor into an artery—usually the
radial or
femoral artery—and continuously measuring the PP waveform. This is generally done by connecting the catheter to a signal processing device with a display. The PP waveform can then be analysed to provide measurements of cardiovascular performance. Changes in vascular function, the position of the catheter tip or damping of the pressure waveform signal will affect the accuracy of the readings. Invasive PP measurements can be calibrated or uncalibrated.
Calibrated PP – PiCCO, LiDCO (
PULSION Medical Systems AG, Munich, Germany) and PulseCO (LiDCO Ltd, London, England) generate continuous
Q by analysing the arterial PP waveform. In both cases, an independent technique is required to provide calibration of continuous
Q analysis because arterial PP analysis cannot account for unmeasured variables such as the changing compliance of the vascular bed. Recalibration is recommended after changes in patient position, therapy or condition. In PiCCO, transpulmonary thermodilution, which uses the Stewart-Hamilton principle but measures temperatures changes from central venous line to a central arterial line, i.e., the femoral or axillary arterial line, is used as the calibrating technique. The
Q value derived from cold-saline thermodilution is used to calibrate the arterial PP contour, which can then provide continuous
Q monitoring. The PiCCO algorithm is dependent on blood pressure waveform morphology (mathematical analysis of the PP waveform), and it calculates continuous
Q as described by Wesseling and colleagues. Transpulmonary thermodilution spans right heart, pulmonary circulation and left heart, allowing further mathematical analysis of the thermodilution curve and giving measurements of cardiac filling volumes (End-diastolic volume|), intrathoracic blood volume and extravascular lung water. Transpulmonary thermodilution allows for less invasive
Q calibration but is less accurate than PA thermodilution and requires a central venous and arterial line with the accompanied infection risks. In LiDCO, the independent calibration technique is
lithium chloride dilution using the Stewart-Hamilton principle. Lithium chloride dilution uses a peripheral vein and a peripheral arterial line. Like PiCCO, frequent calibration is recommended when there is a change in Q. Calibration events are limited in frequency because they involve the injection of lithium chloride and can be subject to errors in the presence of certain muscle relaxants. The PulseCO algorithm used by LiDCO is based on pulse power derivation and is not dependent on waveform morphology.
Statistical analysis of arterial pressure – FloTrac/Vigileo FloTrac/Vigileo (
Edwards Lifesciences) is an uncalibrated, haemodynamic monitor based on pulse contour analysis. It estimates cardiac output (
Q) using a standard arterial catheter with a manometer located in the femoral or radial artery. The device consists of a high-fidelity pressure transducer, which, when used with a supporting monitor (Vigileo or EV1000 monitor), derives left-sided cardiac output (
Q) from a sample of arterial pulsations. The device uses an algorithm based on the
Frank–Starling law of the heart, which states pulse pressure (PP) is proportional to stroke volume (SV). The algorithm calculates the product of the standard deviation of the arterial pressure (AP) wave over a sampled period of 20 seconds and a vascular tone factor (Khi, or χ) to generate stroke volume. The equation in simplified form is: SV = \mathrm{std}(AP) \cdot \chi, or, BP \cdot k \mathrm{\ (constant)}. Khi is designed to reflect arterial resistance; compliance is a multivariate polynomial equation that continuously quantifies arterial compliance and vascular resistance. Khi does this by analyzing the morphological changes of arterial pressure waveforms on a bit-by-bit basis, based on the principle that changes in compliance or resistance affect the shape of the arterial pressure waveform. By analyzing the shape of said waveforms, the effect of vascular tone is assessed, allowing the calculation of SV.
Q is then derived using equation (). Only perfused beats that generate an arterial waveform are counted for in HR. This system estimates Q using an existing arterial catheter with variable accuracy. These arterial monitors do not require intracardiac catheterisation from a pulmonary artery catheter. They require an arterial line and are therefore invasive. As with other arterial waveform systems, the short set-up and data acquisition times are benefits of this technology. Disadvantages include its inability to provide data regarding right-sided heart pressures or mixed venous oxygen saturation. The measurement of Stroke Volume Variation (SVV), which predicts volume responsiveness is intrinsic to all arterial waveform technologies. It is used for managing fluid optimisation in high-risk surgical or critically ill patients. A physiologic optimization program based on haemodynamic principles that incorporates the data pairs SV and SVV has been published. Arterial monitoring systems are unable to predict changes in vascular tone; they estimate changes in vascular compliance. The measurement of pressure in the artery to calculate the flow in the heart is physiologically irrational and of questionable accuracy, and of unproven benefit. Arterial pressure monitoring is limited in patients off-ventilation, in atrial fibrillation, in patients on vasopressors, and in those with a dynamic autonomic system such as those with sepsis. and in various haemodynamic states. It can be used to monitor pediatric and mechanically supported patients. Generally monitored haemodynamic values, fluid responsiveness parameters and an exclusive reference are provided by PRAM: Cardiac Cycle Efficiency (CCE). It is expressed by a pure number ranging from 1 (best) to -1 (worst) and it indicates the overall heart-vascular response coupling. The ratio between heart performance and consumed energy, represented as CCE "stress index", can be of paramount importance in understanding the patient's present and future courses.
Impedance cardiography Impedance cardiography (often abbreviated as ICG, or Thoracic Electrical Bioimpedance (TEB)) measures changes in
electrical impedance across the thoracic region over the cardiac cycle. Lower impedance indicates greater intrathoracic fluid volume and blood flow. By synchronizing fluid volume changes with the heartbeat, the change in impedance can be used to calculate stroke volume, cardiac output and systemic vascular resistance. Both invasive and non-invasive approaches are used. The reliability and validity of the non-invasive approach has gained some acceptance, although there is not complete agreement on this point. The clinical use of this approach in the diagnosis, prognosis and therapy of a variety of diseases continues. Non-invasive ICG equipment includes the Bio-Z Dx, the Niccomo, and TEBCO products by BoMed.
Ultrasound dilution Ultrasound dilution (UD) uses body-temperature normal saline (NS) as an indicator introduced into an extracorporeal loop to create an atrioventricular (AV) circulation with an ultrasound sensor, which is used to measure the dilution then to calculate cardiac output using a proprietary algorithm. A number of other haemodynamic variables, such as total end-diastole volume (TEDV), central blood volume (CBV) and active circulation volume (ACVI) can be calculated using this method. The UD method was firstly introduced in 1995. It was extensively used to measure flow and volumes with extracorporeal circuit conditions, such as
ECMO and
Haemodialysis, leading more than 150 peer reviewed publications. UD has now been adapted to
intensive care units (ICU) as the COstatus device. The UD method is based on ultrasound indicator dilution. Blood ultrasound velocity (1560–1585 m/s) is a function of total blood protein concentration—sums of proteins in plasma and in red blood red cells—and temperature. Injection of body-temperature normal saline (ultrasound velocity of saline is 1533 m/s) into a unique AV loop decreases blood ultrasound velocity, and produces dilution curves. UD requires the establishment of an extracorporeal circulation through its unique AV loop with two pre-existing arterial and central venous lines in ICU patients. When the saline indicator is injected into the AV loop, it is detected by the venous clamp-on sensor on the loop before it enters the patient's heart's right atrium. After the indicator traverses the heart and lung, the concentration curve in the arterial line is recorded and displayed on the COstatus HCM101 Monitor. Cardiac output is calculated from the area of the concentration curve using the Stewart-Hamilton equation. UD is a non-invasive procedure, requiring only a connection to the AV loop and two lines from a patient. UD has been specialised for application in pediatric ICU patients and has been demonstrated to be relatively safe although invasive and reproducible.
Electrical cardiometry Electrical cardiometry is a non-invasive method similar to Impedance cardiography; both methods measure thoracic electrical bioimpedance (TEB). The underlying model differs between the two methods; Electrical cardiometry attributes the steep increase of TEB beat-to-beat to the change in orientation of red blood cells. Four standard ECG electrodes are required for measurement of cardiac output. Electrical Cardiometry is a method trademarked by Cardiotronic, Inc., and shows promising results in a wide range of patients. It is currently approved in the US for use in adults, children and babies. Electrical cardiometry monitors have shown promise in postoperative cardiac surgical patients, in both haemodynamically stable and unstable cases.
Magnetic resonance imaging Velocity-encoded phase contrast Magnetic resonance imaging (MRI) is the most accurate technique for measuring flow in large vessels in mammals. MRI flow measurements have been shown to be highly accurate compared to measurements made with a beaker and timer, and less variable than the Fick principle and thermodilution. Velocity-encoded MRI is based on the detection of changes in the phase of proton
precession. These changes are proportional to the velocity of the protons' movement through a magnetic field with a known gradient. When using velocity-encoded MRI, the result is two sets of images, one for each time point in the cardiac cycle. One is an anatomical image and the other is an image in which the signal intensity in each
pixel is directly proportional to the through-plane velocity. The average velocity in a vessel, i.e., the
aorta or the
pulmonary artery, is quantified by measuring the average signal intensity of the pixels in the cross-section of the vessel then multiplying by a known constant. The flow is calculated by multiplying the mean velocity by the cross-sectional area of the vessel. This flow data can be used in a flow-versus-time graph. The area under the flow-versus-time curve for one
cardiac cycle is the stroke volume. The length of the cardiac cycle is known and determines heart rate;
Q can be calculated using equation (). MRI is typically used to quantify the flow over one cardiac cycle as the average of several heart beats. It is also possible to quantify the stroke volume in real-time on a beat-for-beat basis. While MRI is an important research tool for accurately measuring
Q, it is currently not clinically used for haemodynamic monitoring in emergency or intensive care settings. , cardiac output measurement by MRI is routinely used in clinical cardiac MRI examinations.
Dye dilution method The dye dilution method is done by rapidly injecting a dye,
indocyanine green, into the right atrium of the heart. The dye flows with the blood into the aorta. A probe is inserted into the aorta to measure the concentration of the dye leaving the heart at equal time intervals [0,
T] until the dye has cleared. Let
c(
t) be the concentration of the dye at time
t. By dividing the time intervals from [0,
T] into subintervals Δ
t, the amount of dye that flows past the measuring point during the subinterval from t=t_{i-1} to t=t_i is: (concentration)(volume)=c(t_i)(F\Delta t) where F is the rate of flow that is being calculated. The total amount of dye is: \sum_{i=1}^nc(t_i)(F\Delta t)=F \sum_{i=1}^nc(t_i)(\Delta t) and, letting n\rightarrow\infty, the amount of dye is: A=F\int_{0}^{T} c(t)dt Thus, the cardiac output is given by: F=\frac{A}{\int_{0}^{T} c(t)dt} where the amount of dye injected A is known, and the integral can be determined using the concentration readings. The dye dilution method is one of the most accurate methods of determining cardiac output during exercise. The error of a single calculation of cardiac output values at rest and during exercise is less than 5%. This method does not allow measurement of 'beat to beat' changes, and requires a cardiac output that is stable for approximately 10 s during exercise and 30 s at rest. ==Factors influencing cardiac output==