Cardiac output

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Major factors influencing cardiac output – Cardiac output is influenced by heart rate and stroke volume, both of which are also variable.[1]

Cardiac Output, CO (Q or  \dot Q_{c} or CO ) is the volume of blood being pumped by the heart, in particular by a left or right ventricle in the time interval of one minute. CO may be measured in many ways, for example dm3/min (1 dm3 equals 1 litre).

CO = Stroke Volume, SV × Heart Rate, HR
Fig.1: Aortic blood pressure and aortic blood flow over one heartbeat interval: S = Systolic blood pressure; D = Diastolic blood pressure; MAP = Mean Arterial Pressure; SV = Stroke Volume; DN = dicrotic notch (aortic valve closure)

The relationship between the instantaneous values of aortic blood pressure and blood flow through the aortic valve over one heartbeat interval and their mean values are depicted in Fig.1. Their instantaneous values may be used in research; in clinical practice, their mean values, MAP and SV, are adequate.

CO value in all resting mammals of normal weight is a linear function of their weight with a slope of 0.1 l/min/kg.[2][3] 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 obese subject is more complex and one "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:

BSA[m²] = W0.425[kg] × H0725[cm] × 0.007184     {Eq.2}

The resulting indexed parameters are Stroke Index, SI (ml/beat/m²) defined as

SI[ml/beat/m²] = SV[ml]/BSA[m²]

and Cardiac Index, CI (l/min/m²), defined as

CI[l/min/m²] = CO[l/min]/BSA[m²]

These indexed blood flow parameter then exhibit normal range:

It is for the Stroke Index: 35 < SInormal < 65 ml/beat/m² and for the Cardiac Index: 2.8 < CInormal < 4.2 l/min/m².

The CO equation for indexed parameters than changes to

CI[l/min/m²] = (SI[ml/beat/m²] × HR[bpm])/1000

Clinical uses[edit]

Summary of Major factors influencing cardiac output – The primary factors influencing HR include autonomic innervation plus endocrine control. Not shown are environmental factors, such as electrolytes, metabolic products, and temperature. The primary factors controlling SV include preload, contractility, and afterload. Other factors such as electrolytes may be classified as either positive or negative inotropic agents.[1]

The function of the heart is to transport blood to deliver oxygen, nutrients and chemicals to the cells of the body to ensure their survival and proper function and to remove the cellular wastes. Since the heart is a 'demand pump', that 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 (Q). This, in turn, is controlled principally by the demand for oxygen by the cells of the body and the capacitance of the arterio-venous system. If the body has a high metabolic oxygen demand then the metabolically controlled flow through the tissues is increased, leading to a greater flow of blood back to the heart. This is also modified by the function of the vessels of the body as they actively relax and contract thereby increasing and decreasing the resistance to flow.[citation needed]

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 70 and 120 ml, a factor of only 1.7.[4][5][6]

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, the LV is filled with blood to the capacity known as end diastolic volume (EDV) during the filling phase or diastole. 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. Clearly the EF is dependent on the ventricular EDV which may vary with ventricular disease associated with ventricular dilatation. Even with LV dilatation and impaired contraction the Q may remain constant due to an increase in EDV.[citation needed]

Stroke Volume (SV) = EDV – ESV
Ejection Fraction (EF) = (SV / EDV) × 100%
Cardiac Output (Q) = SV × HR
Cardiac Index (CI) = Q / Body Surface Area (BSA) = SV × HR/BSA
HR is Heart Rate, expressed as BPM (Beats Per Minute)
BSA is Body Surface Area in square metres.

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. 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.[citation needed]

Measuring cardiac output[edit]

There are a number of clinical methods for measurement of Q ranging from direct intracardiac catheterisation to non-invasive measurement of the arterial pulse. Each method has unique strengths and weaknesses and 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, so there is an increasing focus on development of non-invasive methods.[7][8][9]

The Fick principle[edit]

Main article: Fick principle

The Fick principle was first described by Adolf Eugen Fick in 1870 and assumes that the rate at which oxygen is consumed is a function of the rate of blood flows and the rate of oxygen picked up by the red blood cells. The Fick principle involves calculating the oxygen consumed over a given period of time from measurement of the oxygen concentration of the venous blood and the arterial blood. Q can be calculated from these measurements:

  • 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:

VO2 = (Q×CA) - (Q×CV)


  • CA = Oxygen content of arterial blood
  • CV = Oxygen content of venous blood.

This allows us to say

 Q\ = \frac{{{V}_O}_2}{{C}_A - {C}_V}

and therefore calculate Q. Note that (CA – CV) is also known as the arteriovenous oxygen difference.[10]

While considered to be the most accurate method for Q measurement, Fick is invasive, requires time for the sample analysis, and 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 modifications use inert gas 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 hemoglobin molecules in the red blood cells. Measuring the content of hemoglobin in the blood and the percentage of saturation of hemoglobin (the oxygen saturation of the blood) is a simple process and is readily available to physicians. Using the fact that each gram of hemoglobin can carry 1.34 ml of O2, the oxygen content of the blood (either arterial or venous) can be estimated by the following formula:

 \mathrm{Oxygen\ content\ of\ blood} = \left [\mathrm{hemoglobin} \right] \left ( g/dl \right ) \ \times\ 1.34 \left ( ml\ \mathrm{O}_2 /\mathrm{g\ of\ hemoglobin} \right ) \times\ \mathrm{saturation\ of\ blood}\ \left ( \mathrm{percent} \right )\ +\ 0.0032\times\ \mathrm{partial\ pressure\ of\ oxygen} \left ( torr \right )

Dilution methods[edit]

The output of heart is equal to the amount of indicator injected divided by its 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 reflects the Q. The method measures the concentration of a dye at different points in the circulation, usually from an intravenous injection and then at a downstream sampling site, usually in a systemic artery. More specifically, the Q is equal to the quantity of indicator dye injected divided by the area under the dilution curve measured downstream (the Stewart (1897)-Hamilton (1932) equation):

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

The trapezoid rule is often used as an approximation of this integral.

Pulmonary artery thermodilution (trans-right-heart thermodilution)[edit]

The indicator method was further developed with replacement of the indicator dye by heated or cooled fluid and temperature change measured at different sites in the circulation rather than dye concentration; this method is known as thermodilution. The pulmonary artery catheter (PAC), also known as the Swan-Ganz catheter, was introduced to clinical practice in 1970 and provides direct access to the right heart for thermodilution measurements. Continuous invasive cardiac monitoring in the Intensive Care Unit has been all but phased out in an age of hospital acquired infection. Use of the PAC is still useful in right heart study in the cardiac catheterization laboratory today.

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 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 (The "thermodilution curve"). Enabled Thermistor technology allowed 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. Under this unique method, three or four repeated measurements or passes are usually averaged to improve accuracy.[11][12] Modern catheters are fitted with a heating filament which intermittently heats and measures the thermodilution curve providing serial Q measurement. However, these take an average of measurements made 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 use of the PAC does not improve patient outcomes.[7][8][9] 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.[13] PAC use is in decline as clinicians move to less invasive and more accurate technologies for monitoring hemodynamics.

Doppler ultrasound method[edit]

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 Doppler 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 = valve orifice cross sectional area; use pr²
  • r = valve radius
  • VTI = the velocity time integral of the trace of the Doppler flow profile

Doppler ultrasound is non-invasive, accurate and inexpensive and is a routine part of clinical ultrasound with high levels of reliability and reproducibility having been in clinical use since the 1960s.


Echocardiography is a noninvasive method of quantifying cardiac output using Ultrasound. Two dimensional (2D) ultrasound with Doppler measurements are used together to calculate Cardiac Output. 2D measurement of the diameter (d) of the aortic annulus allows calculation of the flow CSA (cross-sectional area) 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) which is then multiplied by the Heart Rate to obtain cardiac output. Although used in clinical medicine, it has a wide test-retest variability.[14] 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 the others. Measurement of the pulmonary valve to calculate right sided CO is an alternative, that is not necessarily more reproducible. While in wide general use the technique is time consuming and 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]

Transcutaneous Doppler: USCOM[edit]

The Ultrasonic Cardiac Output Monitor (USCOM,[13] 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 CSA's allowing both right and left sided Q measurements. This also significantly improves reproducibility compared with the echocardiographic method and therefore increases sensitivity for detection of changes in flow. Real time Automatic tracing of the Doppler flow profile allows for beat to beat right and left sided Q measurements significantly simplifying operation and reducing the time of acquisition compared with the conventional echocardiographic method. USCOM has been validated from 0.12 l/min to 18.7 l/min[15] in neonates,[16] children[17] and adults.[18] 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.[13] This accuracy has ensured high levels of clinical utility across a range of applications including sepsis, heart failure and hypertension.[19][20][21]

Transoesophageal Doppler: TOD[edit]

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,[22][23][24][25][26][27][28][29] and has been recommended by the UK's National Institute for Health and Clinical Excellence (NICE).[30] One limitation is that ODM measures the velocity of blood and not true Q, therefore relies on a nomogram[31] 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[edit]

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 the changes in pressure associated with changes in arterial function (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:

Non-invasive PP – Sphygmomanometry and Tonometry[edit]

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 analysis made of these pressure signals. However as the heart and vessels function independently and sometimes paradoxically the changes in the PP both reflect and mask changes in Q. So these measures represent combined cardiac and vascular function only. A similar system that uses the arterial pulse is the pressure recording analytical method (PRAM).

Finapres methodology[edit]

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: inside pressure (= intra-arterial pressure) equals outside pressure (= finger cuff pressure) by clamping the artery to a certain volume. 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 patients on vasopressors.

In 1978 scientists at BMI-TNO, the research unit of Netherlands Organization for Applied Scientific Research at The University of Amsterdam, invented and patented a series of additional key elements to make the volume clamp work in clinical practice, among them: 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 least the invention of a setpoint 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 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 the finger cuff pressure data. A generalized algorithm to correct for the pressure level difference between the finger and brachial sites within an individual patient was developed and this correction worked under all circumstances that it was tested, even when it was not designed for it, since it applied general physiological principles. The first implementation of this innovative brachial pressure waveform reconstruction was in the Finometer, the successor of Finapres that BMI-TNO introduced in 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 hemodynamics, based on two notions:

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

Recent work comparing nonivasive peripheral vascular monitors suggests modest clinical utility restricted to patients with normal and invariant circulation.[32]

Invasive PP[edit]

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 usually done by connecting the catheter to a signal processing and display device. 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[edit]

PiCCO (PULSION Medical Systems AG, Munich, Germany) and PulseCO (LiDCO Ltd, London, England) generate continuous Q by analysis of the arterial PP waveform. In both cases, an independent technique is required to provide calibration of the 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. femoral or axillary) arterial line. The Q derived from this 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 (i.e. mathematical analysis of the PP waveform) and calculates continuous Q as described by Wesseling and co-workers.[33] 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 attendant 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 to a peripheral arterial line. Like PiCCO frequent calibration is recommended when there is a change in Q.[34] 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[edit]

FloTrac/Vigileo (Edwards Lifesciences LLC, U.S.A.) is an uncalibrated pulse contour analysis-based hemodynamic monitor that 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 wave (AP) (over a sampled period of time of 20 seconds) and a vascular tone factor (Khi) to generate stroke volume. The equation in simplified form is as follows: SV=std(AP) * Khi or BP x k(constant). Khi is conceived 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 morphologic change of the 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 the arterial pressure waveform, the effect of vascular tone is assessed allowing calculation of SV. Cardiac Output (Q) is then derived utilizing the equation Q=HR*SV. Only perfused beats that generate an arterial waveform are counted for HR.[citation needed]

This system estimates Q using an existing arterial catheter with variable accuracy and precision. While these invasive arterial monitors do not require intracardiac catheterisation from a pulmonary artery catheter, they do require an arterial line and are invasive. As with the 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.[35][36] Intrinsic to all arterial waveform technologies is the measurement of Stroke Volume Variation (SVV) which predicts volume responsiveness and is used for managing fluid optimization in high risk surgical or critically ill patients. A Physiologic Optimization Program based on hemodynamic principles that incorporates the data pairs SV and SVV has been published.[37] Further arterial monitoring systems are unable to predict changes in vascular tone and so 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,[38] and of unproven benefit.[39] 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.[34]

Uncalibrated, pre-estimated demographic data-free — PRAM[edit]

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 similarly 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, no constant values of impedance deriving from an external calibration neither form pre-estimated in vivo/in vitro data are needed.

PRAM has been validated against the considered gold standard methods in stable condition[40] and in various hemodynamic states;[41] it can be used to monitor pediatric[42] and mechanically supported[43] patients.

A part to generally monitored hemodynamic values and to fluid responsiveness parameters, an exclusive reference is also provided by PRAM: Cardiac Cycle Efficiency (CCE). Expressed by a pure number ranging from 1 (the best) and -1 (the worse) it indicates the overall heart-vascular response coupling; the ratio between the heart performed and consumed energy, represented as CCE "stress index", can be of paramount importance in understanding patient present and next future course.[44]

Impedance cardiography[edit]

Impedance cardiography (often related as ICG or TEB) is a method that measures changes in impedance across the thoracic region over the cardiac cycle. Lower impedance indicates greater the intrathoracic fluid volume and blood flow. Therefore, by synchronizing fluid volume changes with heartbeat, the change in impedance can be used to calculate stroke volume, cardiac output, and systemic vascular resistance.[45]

Both invasive and non-invasive approaches are being used.[46] The noninvasive approach has achieved some acceptance with respect to its reliability and validity.[47][48][49][50] although there is not complete agreement on this point.[51] The clinical use of this approach in a variety of diseases continues.[52]

Noninvasive ICG equipment includes the Bio-Z Dx,[53] the niccomo[54] and TEBCO products by BoMed[55][56]

Ultrasound dilution method[edit]

Ultrasound dilution (UD) uses body temperature normal saline (NS) as an indicator introduced into an extracorporeal loop to create an AV circulation, with an ultrasound sensor used to measure the dilution and then calculate cardiac output using a proprietary algorithm. A number of other hemodynamic 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.,[57] and it was used extensively to measure flow and volumes with extracorporeal circuits condition such as ECMO[58][59] and Hemodialysis,[60][61] leading more than 150 peer reviewed publications, and now it has adapted to intensive care units (ICU) settings as COstatus.[62]

The UD method is based on ultrasound indicator dilution.[63] 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 produce 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 A-V loop, it is detected by the venous clamp-on sensor on the AV 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 by the classic Stewart-Hamilton equation. It is a non-invasive procedure only by connection the AV loop and two lines of 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[edit]

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 or patients (is currently US market approved for use in adults, pediatrics, and neonates). Electrical Cardiometry monitors have shown promise in postoperative cardiac surgical patients (both hemodynamicially stable and unstable).[64]

Magnetic Resonance Imaging[edit]

Velocity encoded phase contrast Magnetic Resonance Imaging (MRI)[65] 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[66] and less variable than both the Fick principle[67] and thermodilution.[68]

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 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 hence 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 as the product of stroke volume and heart rate. MRI is typically used to quantify the flow over one cardiac cycle as the average of several heart beats, but it is also possible to quantify the stroke volume in real time on a beat-for-beat basis.[69]

While MRI is an important research tool for accurately measuring Q, it is currently not clinically used for hemodynamic monitoring in the emergency or intensive care setting. Cardiac output measurement by MRI is currently routinely used as a part of clinical cardiac MRI examinations.[70]

Cardiac output and vascular resistance[edit]

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 = Mean Aortic (or Arterial) Blood Pressure in mmHg,

RAP = Mean Right Atrial Pressure in mmHg and

TPR = Total Peripheral Resistance in dynes-sec-cm-5.

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

Cardiac output and respiration[edit]

Q is affected by the phase of respiration with intra-thoracic pressure changes influencing diastolic heart filling and therefore Q. Breathing in reduces intra-thoracic pressure, filling the heart and increasing Q, while breathing out increases intra-thoracic pressure, reduces heart filling and Q. This respiratory response is called stroke volume variation and can be used as an indicator of cardiovascular health and disease. These respiratory changes are important, particularly during mechanical ventilation, and Q should therefore be measured at a defined phase of the respiratory cycle, usually end-expiration.[citation needed]


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; hemostasis restored[1] Increasing blood flow and pressure due to increasing cardiac output; hemostasis 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; hemostasis restored[1] Decreasing blood flow and pressure due to decreasing cardiac output; hemostasis restored[1]

Combined cardiac output[edit]

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.[71]

Cardiac input[edit]

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.

Example values[edit]

Measure Typical value Normal range
end-diastolic volume (EDV) 120 mL[72][non-primary source needed] 65–240 mL[72][non-primary source needed]
end-systolic volume (ESV) 50 mL[72][non-primary source needed] 16–143 mL[72][non-primary source needed]
stroke volume (SV) 70 mL 55–100 mL
ejection fraction (Ef) 58% 55–70%[73]
heart rate (HR) 75 bpm 60–100 bpm[74]
cardiac output (CO)oo 5.25 L/minute 4.0–8.0 L/min[75]


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    • End-systolic volume (left ventricular) – average 50.1 and range, 16 – 143 mL:
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