Cardiac output

Cardiac output (CO), also denoted by the symbols <math>Q</math> and <math> \dot Q_{c} </math>, is a term in cardiac physiology that describes the volume of blood being pumped by the heart, in particular by a left or right ventricle, per unit time. CO values can be represented using many different physical units, such as dm3/min and L/min.

Since cardiac output is related to the amount of blood delivered to various parts of the body, it is an important indicator of how efficiently the heart is able to meet the demands of the body. For instance, infections are correlated with high CO and heart failure with low CO.

Along with stroke volume (SV), cardiac output is a global blood flow parameter of interest in hæmodynamics – the study of the flow of blood under external forces. The factors affecting stroke volume and heart rate also affect cardiac output. The figure to the right illustrates this dependency and lists a few of these factors. A detailed hierarchical illustration is provided in a subsequent figure.

There exist many methods to measure CO, both invasively and non-invasively, each with its own trade-offs. Unfortunately, there is a lack of a golden standard or reference measurement against which all of these methods can be compared.


The function of the heart is to drive blood through the circulatory system in a cycle that delivers oxygen, nutrients and chemicals to the cells of the body, and removes cellular waste. Since it pumps out whatever blood comes back into it from the venous system, it is effectively the amount of blood returning to the heart that determines how much blood the heart pumps out – its cardiac output, Q. Cardiac output is classically defined alongside stroke volume (SV) and the heart rate (HR) as:


<p style="margin:0; font-size:4pt;"> 






</dl>In standardizing what CO values are considered to be within normal range independent of the size of the subject's body, the accepted convention is to further index equation (1) using Body surface area (BSA), giving rise to the Cardiac index (CI). This is detailed in equation (2) below.


There are a number of clinical methods to measure Q, ranging from direct intracardiac catheterisation to non-invasive measurement of the arterial pulse. Each method has unique strengths and weaknesses. Relative comparison is limited by the absence of a widely accepted "gold standard" measurement. Q can also be affected significantly by the phase of respiration – intra-thoracic pressure changes influence diastolic filling and therefore Q. This is especially important during mechanical ventilation where Q can vary by up to 50%[citation needed] across a single respiratory cycle. Q should therefore be measured at evenly spaced points over a single cycle or averaged over several cycles.[citation needed]

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.[2][3][4]

Fick principle

Main article: Fick principle
File:Spirometry NIH.jpg
An illustration of how spirometry is done

The Fick principle, first described by Adolf Eugen Fick in 1870, assumes that the rate at which oxygen is consumed is a function of the rate of blood flow and the rate of oxygen picked up by the red blood cells. An application of the Fick principle involves calculating the oxygen consumed over a given period of time by measuring the oxygen concentration of venous blood and arterial blood. Q is calculated from these measurements as follows.

  • VO2 consumption per minute using a spirometer (with the subject re-breathing air) and a CO2 absorber
  • the oxygen content of blood taken from the pulmonary artery (representing mixed venous blood)
  • the oxygen content of blood from a cannula in a peripheral artery (representing arterial blood)

From these values, we know that:

<math>V_{O_2} = (Q \times C_A) - (Q \times C_V)</math>


  • CA is the oxygen content of arterial blood, and,
  • CV is the oxygen content of venous blood.

This allows us to say

<math> Q\ = \frac{{{V}_O}_2}{{C}_A - {C}_V} </math>

and therefore calculate Q. Note that (CA – CV) is also known as the arteriovenous oxygen difference.[5]

While considered to be the most accurate method for the measurement of Q, the Fick method is invasive and requires time for sample analysis. Further, accurate oxygen consumption samples are difficult to acquire. There have also been modifications to the Fick method where respiratory oxygen content is measured as part of a closed system and the consumed oxygen calculated using an assumed oxygen consumption index which is then used to calculate Q. Other variations use inert gases as tracers and measure the change in inspired and expired gas concentrations to calculate Q (Innocor, Innovision A/S, Denmark).

Additionally, the calculation of the arterial and venous oxygen content of the blood is a straightforward process. Almost all oxygen in the blood is bound to hæmoglobin molecules in the red blood cells. Measuring the content of hæmoglobin in the blood and the percentage of saturation of hæmoglobin (the oxygen saturation of the blood) is a simple process and is readily available to physicians. Using the fact that each gram of haemoglobin can carry 1.34 ml of O2, the oxygen content of the blood (either arterial or venous) can be estimated using the following formula.

<math> \begin{align}

\mathrm{Oxygen\ content\ of\ blood} &= \left [\mathrm{haemoglobin} \right] \left ( g/dl \right ) \ \times\ 1.34 \left ( ml\ \mathrm{O}_2 /\mathrm{g\ of\ haemoglobin} \right ) \\ &\times\ \mathrm{saturation\ of\ blood}\ \left ( \mathrm{percent} \right )\ +\ 0.0032\ \times\ \mathrm{partial\ pressure\ of\ oxygen} \left ( torr \right ) \end{align} </math>

Dilution methods

The cardiac output of the heart is the ratio of the amount of indicator injected, to the average concentration in the arterial blood after a single circulation through the heart.

This method was initially described using an indicator dye and assumes that the rate at which the indicator is diluted reflect the Q value. It measures the concentration of the dye at different points in the circulation, usually from an intravenous injection and then at a downstream sampling site, usually in a systemic artery. Once the concentrations are measured, Q is the ratio of the quantity of indicator dye injected, to the area under the dilution curve measured downstream. This is referred to as the Stewart-Hamilton equation, named after George N. Stewart (1897) and William F. Hamilton (1932).

<math>\mathrm{Cardiac\ output} = \frac{\mathrm{Quantity\ of\ Indicator}}{\int_0^\infty \mathrm{Concentration\ of\ Indicator}\cdot {dt}}</math>

When numerical methods are used to calculate this value computationally, the trapezoid rule is often used as an approximation of this integral.

Pulmonary artery thermodilution (trans-right-heart thermodilution)

File:Pulmonary artery catheter english.JPG
Diagram of Pulmonary artery catheter (PAC)

The indicator method was further developed by replacing the indicator dye with heated or cooled fluid. Since temperature change is measured at different sites in the circulation rather than dye concentration, this method is known as thermodilution. The pulmonary artery catheter (PAC) introduced to clinical practice in 1970, also known as the Swan-Ganz catheter, provides direct access to the right heart for thermodilution measurements. Continuous invasive cardiac monitoring in intensive care units has been all but phased out in an age of hospital-acquired infection. The PAC remains useful in right heart study done in cardiac catheterisation laboratories to date.

The PAC is balloon tipped and is inflated, which helps "sail" the catheter balloon through the right ventricle to occlude a smaller branch of the pulmonary artery system. The balloon is then deflated. The PAC thermodilution method involves injection of a small amount (10ml) of cold glucose at a known temperature into the pulmonary artery and measuring the temperature a known distance away (6–10 cm) using the same catheter with temperature sensors set apart at a known distance.

The historically significant Swan-Ganz multi-lumen catheter allows reproducible calculation of cardiac output from a measured time-temperature curve, also know as the thermodilution curve. Thermistor technology enabled the observation that low CO registers temperature change slowly, and inversely, high CO registers temperature change rapidly. The degree of change in temperature is directly proportional to the cardiac output. In this unique method, three or four repeated measurements or passes are usually averaged to improve accuracy.[6][7] Modern catheters are fitted with a heating filament which intermittently heats up and measures the thermodilution curve, providing serial Q measurement. However, these instruments average measurements over 2–9 minutes, depending on the stability of the circulation, and thus do not provide continuous monitoring.

PAC use is complicated by arrhythmias, infection, pulmonary artery rupture, and right heart valve damage. Recent studies in patients with critical illness, sepsis, acute respiratory failure and heart failure suggest that use of the PAC does not improve patient outcomes.[2][3][4] This clinical ineffectiveness may relate to its poor accuracy and sensitivity, which has been demonstrated by comparison with flow probes across a sixfold range of Qs.[8] PAC use is in decline as clinicians move to less invasive and more accurate technologies for monitoring hæmodynamics.

Doppler ultrasound

File:VTI LVOT.png
Doppler signal in the left ventricular outflow tract: Velocity Time Integral (VTI)

This method uses ultrasound and the Doppler effect to measure Q. 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 Q, using the following equations:

  • Q = SV × HR
  • SV = VTI × CSA
  • CSA = π r²


  • 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 with high levels of reliability and reproducibility. It has been in clinical use since the 1960s.


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.[9] 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, and beat-to-beat variation in stroke volume and subtle differences in probe position are others. 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%.[citation needed]


The Ultrasonic Cardiac Output Monitor (USCOM,[8] Uscom Ltd, Sydney, Australia) uses Continuous Wave Doppler (CW) to measure the Doppler flow profile VTI, as in echocardiography, but uses anthropometry to calculate aortic and pulmonary valve diameters and CSAs, allowing both right-sided and left-sided Q measurements. This also significantly improves reproducibility compared to the echocardiographic method and therefore increases sensitivity toward the detection of changes in flow. Real time automatic tracing of the Doppler flow profile allows for beat-to-beat right-sided and left-sided Q measurements, significantly simplifying operation and reducing the time of acquisition compared to the conventional echocardiographic method. USCOM has been validated from 0.12 l/min to 18.7 l/min[10] in neonates,[11] children[12] and adults.[13] This means the method can be applied with equal accuracy to neonates, children and adults for the development of physiologically rational haemodynamic protocols. USCOM is the only method of cardiac output measurement to have achieved equivalent accuracy to the gold standard implantable flow probe.[8] This accuracy has ensured high levels of clinical utility across a range of applications including sepsis, heart failure and hypertension.[14][15][16]


A Transoesophageal echocardiogram (BrE: TOE, AmE: TEE) probe.

Transoesophageal Doppler (TOD), is a term encompassing two main technologies: Transoesophageal Echocardiogram (TOE/TEE), which is primarily used for diagnostic purposes, and (what is commonly termed) oesophageal Doppler (ODM/EDM), primarily used for the clinical monitoring of cardiac output. The latter utilises CW ultrasound and the Doppler effect 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, however 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,[17][18][19][20][21][22][23][24] and has been recommended by the UK's National Institute for Health and Clinical Excellence (NICE).[25] One limitation is that ODM measures the velocity of blood and not true Q, therefore relies on a nomogram[26] 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 paediatrics.

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, e.g., compliance, impedance, etc.

Physiologic or therapeutic changes in vessel diameter are assumed to reflect changes in Q. Put simply, PP methods measure the combined performance of the heart and the vessels, thus limiting the application of PP methods for measurement of Q. This can be partially compensated for by intermittent calibration of the waveform to another Q measurement method and 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.


A sphygmomanometer used to measure pulse pressure (PP) non-invasively

The sphygmomanometer or cuff blood pressure device was introduced to clinical practice in 1903, allowing non-invasive measurements of blood pressure and providing the common PP waveform values of peak systolic and diastolic pressure, which can be used to calculate mean arterial pressure (MAP) and pulse pressure (PP). The pressure in the arteries, measured by sphygmomanometry, is often used as an indicator of the function of the heart. The pressure pulses in the heart are conducted to the arteries, and the arterial pressure is assumed to reflect the function of the heart or the Q. However, no account is made of the elasticity of the arterial bed, or its impact on the pressure signal.

  • The pressure in the heart rises as blood is forced into the aorta
  • The more stretched the aorta, the greater the pulse pressure (PP)
  • In healthy young subjects, each additional 2 ml of blood results in a 1 mmHg rise in pressure
  • Therefore:
SV = 2 ml × Pulse Pressure
Q = 2 ml × Pulse Pressure × HR

By resting a more sophisticated pressure sensing device, a tonometer, against the skin surface and sensing the pulsatile artery, continuous PP wave forms can be acquired non-invasively, and an analysis made of these pressure signals. However, as the heart and vessels function independently, and sometimes paradoxically, the changes in PP both reflect and mask changes in Q. So these measures represent combined cardiac and vascular function. A similar system that relies on arterial pulse is the pressure recording analytical method (PRAM).

Finapres methodology

In 1967, the Czech physiologist Jan Peñáz invented and patented the volume clamp method to measure continuous blood pressure. The principle of the volume clamp method is to provide equal pressures, dynamically, on either side of the wall of an artery. By clamping the artery to a certain volume, inside pressure (intra-arterial pressure) balances outside pressure (finger cuff pressure). He decided that 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. Among them are the use of modulated infra-red light in the optical system inside the sensor, the light-weight, easy-to-wrap finger cuff with velcro fixation, a new pneumatic proportional control valve principle, and last but not the least, the invention of a set point strategy for the determination and tracking of 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 turned out to be surprisingly[according to whom?] accurate, robust and reliable and was never changed since its invention.

The Finapres methodology was developed to use this information to accurately calculate arterial pressure from finger cuff pressure data. A generalised algorithm to correct for the pressure level difference between the finger and brachial sites within an individual patient was developed. This correction worked under all of the circumstances that 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.[citation needed]

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:

  1. pressure and flow are inter-related at each site in the arterial system by their so-called characteristic impedance, and,
  2. at the proximal aortic site, the 3-element Windkessel model of this impedance can be modeled with sufficient accuracy in an individual patient with known age, gender, height and weight.

Recent work comparing non-invasive peripheral vascular monitors suggests modest clinical utility restricted to patients with normal and invariant circulation.[27]


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 all affect the accuracy of the readings. Invasive PP measurements can be calibrated or uncalibrated.

Calibrated PP – PiCCO, LiDCO
File:Lithium chloride.jpg
Lithium chloride, in salt form, on a petri dish

PiCCO (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, as 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 the case of PiCCO, transpulmonary thermodilution is used as the calibrating technique. Transpulmonary thermodilution 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. 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 co-workers.[28] Transpulmonary thermodilution spans right heart, pulmonary circulation and left heart. This allows further mathematical analysis of the thermodilution curve, giving measurements of cardiac filling volumes (GEDV), intrathoracic blood volume, and extravascular lung water. While transpulmonary thermodilution allows for less invasive Q calibration, the method is also less accurate than PA thermodilution and still requires a central venous and arterial line with the accompanied infection risks.

In the case of 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.[29] Calibration events are limited in frequency because it involves injection of lithium chloride, and can be subject to error 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
File:Frank Starling’s curve.png
Cardiac function curve in Frank–Starling's law, illustrating stroke volume (SV) as a function of preload

FloTrac/Vigileo (Edwards Lifesciences LLC, U.S.A.) is an uncalibrated haemodynamic monitor based on pulse contour analysis. It estimates cardiac output (Q) utilizing a standard arterial catheter with a manometer located in the femoral or radial artery. The device consists of a special 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 utilises an algorithm that is based on Frank–Starling law of the heart, that 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 time of 20 seconds, and a vascular tone factor (Khi, or χ) to generate stroke volume. The equation in simplified form is as follows: <math display="inline">SV = \mathrm{std}(AP) \cdot \chi</math>, or, <math display="inline">BP \cdot k \mathrm{\ (constant)}</math>. Khi is designed to reflect arterial resistance, and compliance is a multivariate polynomial equation that continuously quantifies arterial compliance and vascular resistance. Khi does so 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. Cardiac output (Q) is then derived using equation (1). Only perfused beats that generate an arterial waveform are counted for in HR.[citation needed]

This system estimates Q using an existing arterial catheter with variable accuracy and precision. While these arterial monitors do not require intracardiac catheterisation from a pulmonary artery catheter, they do require an arterial line and are therefore invasive. As with other arterial waveform systems, the short time required for set up and data acquisition are additional benefits of this technology. Disadvantages include its inability to provide data regarding right-sided heart pressures, or mixed venous oxygen saturation.[30][31] Intrinsic to all arterial waveform technologies is the measurement of Stroke Volume Variation (SVV), which predicts volume responsiveness. 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.[32]

Further, arterial monitoring systems are unable to predict changes in vascular tone and end up estimating 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,[33] and of unproven benefit.[34] Arterial pressure monitoring is limited in patients off ventilation, in atrial fibrillation, in patients on vasopressors, and in patients with a dynamic autonomic system such as in sepsis.[29]

Uncalibrated, pre-estimated demographic data-free — PRAM

Pressure Recording Analytical Method (PRAM), estimates Q from the analysis of the pressure wave profile obtained from an arterial catheter (radial or femoral access). This PP waveform can then be used to determine Q, similar to FloTrac. As the waveform is sampled at 1000 Hz, the detected pressure curve can be measured to calculate the real (relative to the patient under examination) and actual (beat-to-beat) stroke volume. Unlike FloTrac, neither constant values of impedance from external calibration, nor form pre-estimated in vivo or in vitro data, are needed.

PRAM has been validated against the considered gold standard methods in stable condition[35] and in various haemodynamic states.[36] It can be used to monitor pediatric[37] and mechanically supported[38] patients.

Apart from generally monitored haemodynamic values and fluid responsiveness parameters, an exclusive reference is also provided by PRAM: Cardiac Cycle Efficiency (CCE). Expressed by a pure number, ranging from 1 (best) to -1 (worst), 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 course.[39]

Impedance cardiography

Impedance cardiography (often abbreviated as ICG, or Thoracic Electrical Bioimpedance (TEB)) is a method that measures changes in impedance across the thoracic region over the cardiac cycle. Lower impedance indicates greater intrathoracic fluid volume and blood flow. Therefore, 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.[40]

Both invasive and non-invasive approaches are being used.[41] The non-invasive approach has achieved some acceptance with respect to its reliability and validity,[42][43][44][45] although there is not complete agreement on this point.[46] The clinical use of this approach in the diagnosis, prognosis and therapy of a variety of diseases continues.[47]

Non-invasive ICG equipment includes the Bio-Z Dx,[48] the Niccomo,[49] and TEBCO products by BoMed.[50][51]

Ultrasound dilution method

Ultrasound dilution (UD) uses body temperature normal saline (NS) as an indicator introduced into an extracorporeal loop to create an atriovetricular (AV) circulation, with an ultrasound sensor used to measure the dilution and then calculate cardiac output using a proprietary algorithm. A number of other haemodynamic variables can also be calculated such as total end-diastole volume (TEDV), central blood volume (CBV) and active circulation volume (ACVI).[citation needed]

The UD method was firstly introduced in 1995.[52] It was used extensively to measure flow and volumes with extracorporeal circuit conditions such as ECMO[53][54] and Haemodialysis,[55][56] leading more than 150 peer reviewed publications. It has now been adapted to intensive care units (ICU) as the COstatus device.[57]

The UD method is based on ultrasound indicator dilution.[58] 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), temperature, etc. 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.[citation needed]

UD requires establishment of an extracorporeal circulation through its unique AV loop with two preexisting 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 right heart 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. It 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 a relatively safe, although invasive, and reproducible tool.[citation needed]

Electrical Cardiometry

File:4 electrodes.jpg
Electrode array that measures Thoracic electrical bioimpedance (TEB)

Electrical Cardiometry is a non-invasive method similar to Impedance cardiography, in the fact that both methods measure thoracic electrical bioimpedance (TEB). The underlying model is what differs, being that 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 market for use in adults, pediatrics, and neonates. Electrical cardiometry monitors have shown promise in postoperative cardiac surgical patients, in both haemodynamicially stable and unstable cases.[59]

Magnetic Resonance Imaging

Velocity-encoded phase contrast Magnetic Resonance Imaging (MRI)[60] 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 with a beaker and timer,[61] and less variable than both the Fick principle[62] and thermodilution.[63]

Velocity-encoded MRI is based on detection of changes in the phase of proton precession. These changes are proportional to the velocity of the movement of those protons through a magnetic field with a known gradient. When using velocity-encoded MRI, the result of the MRI scan 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 where 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, and 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 to graph flow versus time. 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, and thereby Q can be calculated using equation (1). 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.[64]

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 currently routinely used as a part of clinical cardiac MRI examinations.[65]


File:2036 Summary of Factors in Cardiac Output.jpg
Hierarchical summary of major factors influencing cardiac output.

Cardiac output is primarily controlled by the oxygen requirement of tissues in the body. In contrast to other pump systems, this makes the heart a demand pump that does not regulate its own output.[66] When the body has a high metabolic oxygen demand, the metabolically controlled flow through the tissues is increased, leading to a greater flow of blood back to the heart. This, in turn, leads to higher cardiac output.

Another factor that controls cardiac output is the capacitance, also known as compliance, of the arterio-vascular channels that carry the blood. As the vessels of the body actively relax and contract, the resistance to blood flow increases and decreases respectively. Thin-walled veins have about eighteen times the capacitance of thick-walled arteries because they are able to carry more blood by virtue of being more distensible.[67]

From this formula, it is clear that the factors affecting stroke volume and heart rate also affect cardiac output. The figure to the right illustrates this dependency and lists a few of these factors. A more detailed hierarchical illustration is provided in a subsequent figure.

Equation (1) reveals HR and SV to be the primary determinants of cardiac output Q. A detailed representation of these factors is illustrated in the figure to the right. The primary factors that influence HR are autonomic innervation plus endocrine control. Not shown are environmental factors, such as electrolytes, metabolic products, and temperature. The determinants of SV during the cardiac cycle are the contractility of the heart muscle, the degree of preload of myocardial distention prior to shortening, and the afterload during ejection.[68] Other factors such as electrolytes may be classified as either positive or negative inotropic agents.[69]

Cardiac response

Table 3: Cardiac response to decreasing blood flow and pressure due to decreasing cardiac output[1]
Baroreceptors (aorta, carotid arteries, venae cavae, and atria) Chemoreceptors (both central nervous system and in proximity to baroreceptors)
Sensitive to Decreasing stretch[1] Decreasing O2 and increasing CO2, H+, and lactic acid[1]
Target Parasympathetic stimulation suppressed[1] Sympathetic stimulation increased[1]
Response of heart Increasing heart rate and increasing stroke volume[1] Increasing heart rate and increasing stroke volume[1]
Overall effect Increasing blood flow and pressure due to increasing cardiac output; haemostasis restored[1] Increasing blood flow and pressure due to increasing cardiac output; haemostasis restored[1]
Table 4: Cardiac response to increasing blood flow and pressure due to increasing cardiac output[1]
Baroreceptors (aorta, carotid arteries, venae cavae, and atria) Chemoreceptors (both central nervous system and in proximity to baroreceptors)
Sensitive to Increasing stretch[1] Increasing O2 and decreasing CO2, H+, and lactic acid[1]
Target Parasympathetic stimulation increased[1] Sympathetic stimulation suppressed[1]
Response of heart Decreasing heart rate and decreasing stroke volume[1] Decreasing heart rate and decreasing stroke volume[1]
Overall effect Decreasing blood flow and pressure due to decreasing cardiac output; haemostasis restored[1] Decreasing blood flow and pressure due to decreasing cardiac output; haemostasis restored[1]

Vascular resistance

The vascular beds are a dynamic and connected part of the circulatory system against which the heart must pump to transport the blood. Q is influenced by the resistance of the vascular bed against which the heart is pumping. For the right heart this is the pulmonary vascular bed, creating Pulmonary Vascular Resistance (PVR), while for the systemic circulation this is the systemic vascular bed, creating Systemic Vascular Resistance (SVR). The vessels actively change diameter under the influence of physiology or therapy, vasoconstrictors decrease vessel diameter and increase resistance, while vasodilators increase vessel diameter and decrease resistance. Put simply, increasing resistance decreases Q; conversely, decreased resistance increases Q.

This can be explained mathematically:

By simplifying Darcy's Law, we get the equation that

Flow = Pressure/Resistance

When applied to the circulatory system, we get:


Where MAP is the Mean Aortic (or Arterial) Blood Pressure in mmHg, RAP is the Mean Right Atrial Pressure in mmHg, and TPR is the Total Peripheral Resistance in dynes·sec/cm5.

However, as MAP>>RAP, and RAP is approximately 0, this can be simplified to:


For the right heart Q ≈ MAP/PVR, while for the left heart Q ≈ MAP/SVR.

Physiologists will often re-arrange this equation, making MAP the subject, to study the body's responses.

As has already been stated, Q is also the product of the heart rate (HR) and the stroke volume (SV), which allows us to say:

Q ≈ (HR × SV) ≈ MAP / TPR

Clinical significance

When Q increases in a healthy but untrained individual, most of the increase can be attributed to an increase in heart rate (HR). Change of posture, increased sympathetic nervous system activity, and decreased parasympathetic nervous system activity can also increase cardiac output. HR can vary by a factor of approximately 3, between 60 and 180 beats per minute, while stroke volume (SV) can vary between Script error: No such module "convert"., a factor of only 1.7.[70][71][72]

Diseases of the cardiovascular system are often associated with changes in Q, particularly the pandemic diseases of hypertension and heart failure. Cardiovascular disease can be associated with increased Q as occurs during infection and sepsis, or decreased Q, as in cardiomyopathy and heart failure.[68] Sometimes, in the presence of ventricular disease associated with dilatation, EDV may vary. An increase in EDV could counterbalance LV dilatation and impaired contraction. From equation (3), the resulting cardiac output Q may remain constant.

The ability to accurately measure Q is important in clinical medicine as it provides for improved diagnosis of abnormalities, and can be used to guide appropriate management.[73]

Example values

Measure Typical value Normal range
end-diastolic volume (EDV) 120 mL[74][non-primary source needed] 65–240 mL[74][non-primary source needed]
end-systolic volume (ESV) 50 mL[74][non-primary source needed] 16–143 mL[74][non-primary source needed]
stroke volume (SV) 70 mL 55–100 mL
ejection fraction (Ef) 58% 55–70%[75]
heart rate (HR) 75 bpm 60–100 bpm[76]
cardiac output (CO)oo 5.25 L/minute 4.0–8.0 L/min[77]

Related measurements

Ejection fraction

A parameter related to SV is ejection fraction (EF). EF is the fraction of blood ejected by the left ventricle (LV) during the contraction or ejection phase of the cardiac cycle or systole. Prior to the start of systole, during the filling phase or diastole, the LV is filled with blood to the capacity known as end diastolic volume (EDV). During systole, the LV contracts and ejects blood until it reaches its minimum capacity known as end systolic volume (ESV), it does not empty completely. The following equations help translate the effect of EF and EDV on cardiac output Q, via SV.


<p style="margin:0; font-size:4pt;"> 







Cardiac input

Cardiac input (CI) is the inverse operation of cardiac output. As cardiac output implies the volumetric expression of ejection fraction, cardiac input implies the volumetric injection fraction (IF).

IF = end diastolic volume (EDV) / end systolic volume (ESV)

Cardiac input is a readily imaged mathematical model of diastole.[clarification needed]

Cardiac index

See also: Cardiac index

In all resting mammals of normal weight, CO value is a linear function of their weight with a slope of 0.1 l/min/kg.[78][79] Fat has about 65% of oxygen demand per weight in comparison to other lean body tissues. As a result, the calculation of normal CO value in an obese subject is more complex and a single common "normal" value of SV and CO for adults cannot exist. All blood flow parameters have to be indexed. The accepted convention is to index them by the Body Surface Area, BSA [m²], by DuBois & DuBois Formula, a function of height and weight:

<math display="block">BSA_{[m^2]} = W^{0.425}_{[kg]} \times H^{0725}_{[cm]} \times 0.007184</math>

The resulting indexed parameters are Stroke Index (SI) and Cardiac Index (CI). Stroke Index, measured in ml/beat/m², is defined as

<math display="block">SI_{[ml/beat/{m}^2]} = \frac{SV_{[ml]}}{BSA_{[m^2]}}</math>

Cardiac Index, measured in l/min/m², is defined as<math display="block">CI_{[l/min/{m}^2]} = \frac{CO_{[l/min]}}{BSA_{[{m}^2]}}</math>

The CO equation (1) for indexed parameters then changes to the following.


<p style="margin:0; font-size:4pt;"> 






</dl>The normal range for these indexed blood flow parameters are between 35 and 65 ml/beat/m² for SI and between 2.8 and 4.2 l/min/m² for CI.[citation needed]

Combined cardiac output

Combined cardiac output (CCO) is the sum of the outputs of the right and left side of the heart. It is useful in fetal circulation, where cardiac outputs from both sides of the heart partly work in parallel by the foramen ovale and ductus arteriosus, both directly supplying the systemic circulation.[80]


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