We finally increased VAD speed during exercise to the maximum possible value and we observed a reduction of wedge pressure (-4.5 mmHg), a slight improvement of cardiac output (8.0 l/min) and a complete unloading of the native ventricle. The VAD can assure a proper hemodynamics at rest, but provides an insufficient unloading of the left ventricle and does not prevent wedge pressure from rising during exercise. Neither the VAD provides major benefits during exercise in terms of total cardiac output, which increases to a similar extend to an unassisted heart failure condition. VAD speed modulation can contribute to better unload the ventricle but the maximal flow reachable with the current devices is below the cardiac output observed in a healthy heart.Patients with a Ventricular Assist Device (VAD) are hemodynamically stable but show an impaired exercise capacity. Aim of this work is to identify and to describe the limiting factors of exercise physiology with a VAD. We searched for data concerning exercise in heart failure condition and after VAD implantation from the literature. Data were analyzed by using a cardiorespiratory simulator that worked as a collector of inputs coming from different papers. As a preliminary step the simulator was used to reproduce the evolution of hemodynamics from rest to peak exercise (ergometer cycling) in heart failure condition. Results evidence an increase of cardiac output of +2.8 l/min and a heart rate increase to 67% of the expected value. Then, we simulated the effect of a continuous-flow VAD at both rest and exercise. Total cardiac output increases of +3.0 l/min (+0.9 l/min due to the VAD and +2.1 l/min to the native ventricle). Since the left ventricle works in a non-linear portion of the diastolic stiffness line, we observed a consistent increase of pulmonary capillary wedge pressure (from 14 to 20 mmHg) for a relatively small increase of end-diastolic volume (from 182 to 189 cm Exercise incompetence is a hallmark of heart failure and is the result of multiple interacting cardiac and vascular impairments. In end-stage heart failure, medication is no longer adequate and heart transplantation or a mechanical assist device (VAD) are the only available treatment options besides palliation. VAD patients, although hemodynamically stable at rest, reach only 50% of the expected exercise capacity . SeveralThese studies evidence the complexity of the exercise pathophysiology in VAD patients and raise some questions about the interplay between the native ventricle and the VAD during exercise induced increases in cardiac output.Given these premises, this study is intended to provide a systematic overview of the knowledge on evolution of VAD patients’ hemodynamics during exercise. At present there are only a few studies describing exercise pathophysiology with a VAD ,3,7, andWe developed a lumped parameters cardiorespiratory simulator in LabVIEW 2014  able to reproduce exercise in different pathophysiological conditions.The left and right hearts are represented by a time-varying elastance model , the sysThe simulator also includes the baroreflex and metabolic peripheral controls, ventilation control and mechanics, gas exchange in lungs and tissues. A complete description of the model is provided in . A schemThe simulator is fed with hemodynamic and respiratory parameters at rest. Exercise is then simulated by increasing the oxygen consumption in the body, differentiating between exercising and non-exercising regions. Since the simulator does not include metabolic activity at cellular and tissue levels, the oxygen consumption and the respiratory quotient are actually two inputs of the simulator. Their values were taken from the literature and refer to HF patients characterized by a lower muscular oxygen extraction and an earlier anaerobic metabolism . Given the oxygen and carbon dioxide partial pressures in the upper body, a dedicated control calculates the minute ventilation. For increasing levels of workload, the minute ventilation augments with consequent increase of tidal volume and of ventilation frequency.The ventilated air is used in the gas exchange module at the level of the pulmonary circulation where the oxygenation and decarbonization of blood takes place.The simulator, arranged with all the components mentioned above, was already used to reproduce exercise physiology in healthy and in heart failure conditions and it was validated with data from literature .In the present work we further adapted the simulator to reproduce the hemodynamic physiology of exercise in severe heart failure condition, with and without a VAD. To this aim, some changes/improvements were implemented:The systole/diastole duration ratio changes according to the heart rate in such a way that diastole shortens more for a high heart rate and vice versa,A model of atrial contraction was implemented to reproduce the contribution of atrial chambers to fill the ventricles during late diastoleVenous vessels are represented as collapsible tubes whose pressure-volume relationship becomes nonlinear for high values of transmural pressureThe effect of the increase of pulmonary venous capillary wedge pressure on the pulmonary vasculature was implementedA VAD model based on the pressure-flow characteristics of HeartMate II (Thoratec Corporation) was added between the left ventricle and the ascending aorta of the simulatorMore details about how these improvements were implemented are reported in S1.Parameters such as resistances, compliances, ventricular contractility represent the inputs of the simulator. The outputs of the simulator are: pressure, flow, volume, O2 and CO2 concentration profiles in every section of the cardiovascular system.A detailed description of the procedure for the determination of the input parameters was provided in . Here weWe here describe how the systolic and diastolic properties of the ventricle were represented.Plv is the left ventricular pressure, Pintr is the intrathoracic pressure, Vlv is the left ventricular volume, V0 is the zero pressure filling volume, t is time ranging from 0 s to 60/HR s, Elvs is the end-systolic elastance and vc is an adimensional ventricular contraction function ranging from 0 during diastole to 1 ad end systole (see the appendix for more information). For the right ventricle a similar equation was implemented.The contractility of both the left and right ventricles was implemented by a time-varying elastance model :Plv(t)=Elvs was reduced to 0.5 mmHg/cm3 and V0 was set to 50 cm3. For HF+VAD we used the same value for Elvs but V0 was reduced to 0 cm3 to account for the unloading effect of the VAD on the left ventricle. For the right ventricle a similar equation was implemented. Right ventricular contractility (Ervs) was taken from [To simulate the impairment of the systolic function, typical for HF, ken from  and refeal and bl are constant parameters describing the stiffness properties of the ventricle during diastole. For HF, al and bl were taken from [The left ventricular stiffness was represented with an exponential function describing the relation between pressures and volumes:ken from  and refeal and bl were estimated fitting the data of 10 patients with a continuous-flow assist device [For HF+VAD, parameters t device : data reThree patients were excluded from the analysis since they had a very different stiffness in comparison to the rest of the group. This difference might be due to large variety of VAD therapy duration (668 ± 596 days), so that in some patients it is likely to assume that the VAD was not left in place for sufficient time to reduce left ventricular dilatation .Ras (Rap) is the systemic (pulmonary) arterial resistance, Pas is the mean systemic arterial pressure, Pcv is the mean central venous pressure, Pap is the mean pulmonary arterial pressure, Pwedge is the pulmonary capillary wedge pressure and CO is the total cardiac output.Systemic and pulmonary vascular resistances were estimated from literature data relative to rest condition as follows:Ras and Rap in HF were calculated averaging data from [Ras was calculated considering data from [Rap averaging data from [Ras was then split into the resistance of upper body, kidneys, splanchnic circulation and lower limbs as reported in [ata from ,16,17,18ata from ,5,13, RaTwo main controls are implemented for the vasculature: the metabolic and the baroreflex controls. We will describe them here briefly.th vascular region has a specific reference value (CO2iRef). If the actual venous oxygen content falls below CO2iRef, hypoxic vasodilation occurs locally. Heart failure patients are chronically exposed to low values of perfusion, therefore their tissues are used to lower values of oxygen concentration. To take this phenomenon into account we set a lower CO2iRef for HF and for HF+VAD in all the vascular compartments : if the actual aortic pressure is lower than Pset vasoconstriction occurs, heart rate and contractility increase and venous tone reduces and vice versa. From a physiological point of view, Pset represents the equilibrium value and it was characterized in the simulator at rest condition: for HF we averaged the mean systemic pressure at rest taken from [Pset automatically increases due to the “baroreflex resetting phenomenon” that progressively shifts the baroreflex control towards higher pressure levels.The baroreflex control works around a set-point value  [Rap at rest and this value was kept constant during exercise simulations.Concerning the pulmonary vasculature, the hemodynamic data of HF patients show a slight decrease of g·s/cm3) ,17,18. WTwo simulation sets were performed: one for the HF and the other for the HF+VAD.To simulate exercise in HF we started by characterizing the simulator at rest. For this purpose we fed hemodynamic and respiratory parameter values at rest into the simulator as explained in the previous paragraphs. The simulator was left running until a steady condition was reached and data were collected. Then we increased oxygen consumption up to 11.5 ml/min/kg (bicycle workload of 56 watts) which is the peak level for severe HF according to ,18. The The same procedure was repeated for HF+VAD. We first set the VAD speed at 9500 rpm and collected the data at rest condition. Then we increased oxygen uptake up to 15.2 ml/min/kg (bicycle workload of 80 watts) which is the peak level according to . The simSimulation results refer to pressures and flows in the cardiorespiratory simulator, reported as average values, calculated over 15 heart cycles to reduce the ventilation effect. As pulmonary capillary wedge pressure, we considered the average value of left atrial pressure obtained from the simulator.In addition, continuous pressure and volume data of the left ventricle were recorded.In 3, the end diastolic volume (VED) from 320 to 355 cm3. Concerning ventilation parameters we observe an increase of minute ventilation from 9.2 to 32.9 l/min and an increase of central arteriovenous oxygen difference from 7.0 to 13.1 ml/dl.In HR from 78 to 115 and 114 bpm, total cardiac output from 4.3 to 7.3 and 8.0 l/min, mean systemic arterial pressure from 88 to 84 and 96 mmHg, mean pulmonary arterial pressure from 22 to 33 and 30 mmHg, wedge pressure from 14 to 20 and 16 mmHg, right atrial pressure from 8 to 15 and 15 mmHg. Left ventricular end systolic volume changes from 154 to 147 and 150 cm3 and the end diastolic volume from 182 to 189 and 183 cm3. Minute ventilation increases from 9.2 to 44 and 44 l/min and central arteriovenous oxygen difference increases from 6.0 to 15.1 and 13.7 ml/dl. Left ventricular flow goes from 0 to 2.1 and 0 l/min, VAD flow changes from 4.3 to 5.2 and 8.0 l/min.Parameters change in order as it follows: 3). In the HF+VAD condition ventricular volumes are in general narrower thanks to the unloading effect of the VAD. At rest condition, the left ventricle does not eject and its peak pressure is below the systemic arterial pressure. During exercise, Elvs increases and preload also increases, these two conditions make the left ventricle ejecting again. When the speed is increased to 12000 rpm, the VAD goes back to full support again and the left ventricle stops ejecting.In dPlv/dVlv slope of 0.26 mmHg/cm3 for HF and of 0.48 mmHg/cm3 for HF+VAD.In Exercise is a particular condition in which the heart and the circulation adapt to accommodate a higher cardiac output to better oxygenate peripheral regions. Venous return increases and the natural reaction of a healthy heart is to increase contractility and heart rate so that the end-diastolic ventricular volumes stay more or less the same as rest values.In a failing heart, the increase of contractility is not so extensive or even absent so the myocardium resort to using the Frank Starling mechanism the ventricle increases its end diastolic volume in order to eject a larger stroke volume ,21.In case of a dilated heart, as we simulated in the present work, we observed an increase of both end-diastolic and end-systolic volumes see Figs  and 5. SConcerning the capillary wedge pressure, previous studies ,21 showeSystemic arterial pressure slightly increases in our simulator during exercise . This iset al. [Some comments need to be addressed with respect to HR: according to our simulations HR at peak exercise is less than 80% of the expected value (considered as the threshold for chronotropic incompetence). Chronotropic incompetence is recognized to be one of the most important indicators of exercise limitation in HF and a critical variable for patients’ performance . The simet al. ). The reVe) increases up to 32.9 l/min for a carbon dioxide removal (VCO2) of 1022 ml/ min. This corresponds to a Ve/VCO2 of 32 ml/ml, a value higher than the physiological threshold of 30 ml/ml and associated with ventilation inefficiency [Finally, according to results in ficiency .In the previous paragraph we provided a descriptive overview of exercise impairments in a HF condition and validated the simulator with data from literature. In our opinion this is a necessary step before reproducing VAD pathophysiology since, as we already stated, VAD and HF conditions share similar cardiovascular and ventilation impairments during exercise .By comparing Figs But during exercise the beneficial effects of the VAD are less evident: the VAD is not capable to refrain wedge pressure and the pulmonary pressure to increase and cardiac output augmentation is just comparable to the one observed in HF (+2.8 l/min in HF and +3.0 l/min in HF+VAD).The increase of wedge pressure was already shown by Hayward et al , but theThe ventricle with a VAD has stiffer elastic properties than failed unassisted ventricles. This means that for the same ventricular volume, wedge pressure is higher in HF+VAD than in HF.dPlv/dVlv relationship, shown in 3) in HF+VAD. These results are in agreement with the echo measurements reported in [The native ventricle under VAD works in a very non-linear portion of the diastolic stiffness. The linearization of the orted in ,5,13 shoThe two observations above are the result of the so called “reverse-remodeling” operated by the VAD. The hemodynamic unloading of VAD on the ventricle stops the vicious cycle of ventricular dysfunction and dilatation observed during heart failure . Fig 2-BCO, the simulator shows that , but during exercise the ventricle provides a significant contribution to total cardiac output (2.1 l/min). The VAD is capable to increase its flow only by +0.9 l/min at exercise.Given this limited increase of CO and the modest contribution of the VAD to it, we investigated if VAD speed modulation could provide some benefits in terms of hemodynamics during exercise. So far the influence of VAD speed increase was investigated in several studies, some reporting significant effects on hemodynamics ,29 and sIt is also important to remark that most of the previous studies investigated a relatively small increase of VAD speed and therefore the question arises whether a more important VAD speed change would have provided a better patient performance. In this study, thanks to the use of a simulator, we could investigate the effects of a VAD running at the maximal possible speed (12000 rpm for the HeartMate II).By increasing the VAD speed we could obtain a higher total cardiac output and an increase of VAD flow compared to exercise at baseline speed . Basically, VAD speed modulation has a large effect on left ventricular unloading but affects CO only mildly.VAD speed also decreases wedge pressure by -4.5 mmHg see , which sThe increase of VAD speed does not affect ventilation substantially but it reduces central arteriovenous oxygen difference of -1.4 ml/dl . This meTo summarize, by combining the literature observations and our simulation results we could speculate that exercise in VAD patients is a very peculiar condition both during systole and diastole. The native ventricle becomes stiffer after the VAD implantation so that the wedge pressure rises even for small changes of ventricular volumes. One could also argue that the decrease in end-diastolic volume, operated by the VAD, is beneficent to decrease systolic wall tension, but is contra-productive by increasing diastolic chamber stiffness which becomes the major obstacle to increased filling and thus stroke volumes.The VAD effect on ventricular stiffness, might not be accompanied by a significant improvement in the myocyte function and in the overall myocardial contractile strength . Hence aThe VAD also shows an inadequate capability in improving its flow during exercise. Our simulations show that in a failed unassisted ventricle CO increases of +2.8 l/min (from 3.7 to 6.5 l/min). Similarly, in a failed assisted ventricle CO increases of +3.0 l/min (from 4.3 to 7.3 l/min), +2.1 l/min due to the left ventricle and the remaining +0.9 l/min due to the VAD. There results evidence that VAD is not capable to properly increase CO neither to unload the left ventricle during exercise.By increasing VAD speed to the highest limit we were just able to re-establish a condition of full support again but total cardiac output reached a plateau of 8.0 l/min. This indicates that VAD speed modulation can produce a better unloading of the left ventricle, but it has a moderate effect on total cardiac output. Future studies should be conducted to investigate whether an even higher support, at present not available with the current VAD technology, could positively affect cardiac output or whether other factors like right ventricular function, peripheral resistance etc. would limit cardiac output anyway. In addition, it is worth to underline that we simulated competent heart valves. As future step, the role of aortic insufficiently, a rather common complications in VAD patients, should be investigated especially in the hypothesis of a VAD speed modulation.Observations of hemodynamic changes in patients with VAD during exercise were collected. A cardio-respiratory simulator was updated with these fragmented data allowing the most reliable simulation of exercise physiology of patient with heart failure and VAD. The simulator indicates that exercise capacity remains limited in VAD patients mainly because of an increased left ventricular stiffness and a non-optimal unloading of the left ventricle at peak exercise.S1 TableList of additional parameters used to characterize the simulator in HF and in VAD rest conditions.(DOCX)Click here for additional data file.


Surgical myectomy is regarded as a gold standard treatment for Hypertrophic Cardiomyopathy (HCM) and may be conducted using different approaches, depending on the location of the thickened heart muscle, to widen the outflow tract and thereby reduce outflow drag to favorably alter the natural course of HCM. We present a novel approach for patient-specific visualization of precise locations of left ventricular (LV) hypertrophy as well as virtual resection of obstructions to facilitate optimal surgical myectomy.The LV endocardium and epicardium were reconstructed in 3D from short-axis cine steady-state free precession (SSFP) cardiac magnetic resonance (CMR) images of five HCM patients. Pre-operative LV volumes were quantified and the location as well as extent of regional hypertrophy was visualized at end-diastole using color-maps of myocardial wall thickness. Next, virtual myectomy was performed to resect the LV in the end-diastolic phase, interactively using a pre-surgical planning software, QuantMD Surgery Explorer . Simulative of myectomy, a real-time, free-form surface deformation tool was used for endocardial point cloud morphing and surface fitting (by Poisson surface reconstruction) to eliminate obstructive hypertrophic muscle, at several surgeon-selected endocardial sites. The spatial influence of deformation at each site was dictated by a Gaussian function affecting normally-outward endocardial vertex movement, expressing spatial influence as a fraction of the Euclidean distance between any given endocardial vertex and a surgeon-specified site of myectomy and magnitude of myectomy at each site by the Gaussian's amplitude.The study cohort (n = 5) had a mean LV mass of 197.3 ± 3.51 g, mean end-diastolic volume (EDV) of 60.8 ± 21.49 mL and mean stroke volume of 47.4 ± 15.27 mL. After virtual myectomy, the mean EDV increased significantly to 85.6 ± 12.78 mL (p = 0.057) with a significant increase in the stroke volume to 71.54 ± 14.68 mL . Color-mapped endocardial surfaces were used to visualize the precise regional extent of virtual myectomy (in mm; see Figure Virtual surgical modeling was employed to leverage patient-specific CMR imaging to guide myectomy in five hypertrophic cardiomyopathy patients presenting with myocardial obstructions in the LV outflow tract, mid-ventricular or apical regions. Virtual myectomy was conducted at the end-diastolic cardiac phase, simulative of the likely shape of the LV in the cardioplegic state. Our methodology offered realistic & dexterous control of virtual resection and may help avoid excessive resection or damage to adjacent ventricular structures while optimizing the precise location and extent of hypertrophic myocardial resection.


Shock in newborn infants has unique etiopathologic origins that require careful assessment to direct specific interventions. Early diagnosis is key to successful management. Unlike adults and pediatric patients, shock in newborn infants is often recognized in the uncompensated phase by the presence of hypotension, which may be too late. The routine methods of evaluation used in the adult and pediatric population are often invasive and less feasible. We aim to discuss the pathophysiology in shock in newborn infants, including the transitional changes at birth and unique features that contribute to the challenges in early identification. Special emphasis has been placed on bedside focused echocardiography/focused cardiac ultrasound, which can be used as an additional tool for early, neonatologist driven, ongoing evaluation and management. An approach to goal oriented management of shock has been described and how bed side functional echocardiography can help in making a logical choice of intervention  in infants with shock. The key to the management of shock in the newborn period is early identification and determination of etiology to provide appropriate care. American College of Critical Care Medicine (ACCM) published clinical guidelines and practice parameters to promote “best practices” and to improve patient outcomes in pediatric and neonatal septic shock in 2002, with a subsequent update in 2007 . In addiDespite widespread dissemination of such guidelines, management of neonatal shock continues to rely on traditional methods of monitoring and management. We aim to discuss the unique pathophysiology associated with shock in newborns, with a focus on the very low birth weight (VLBW) infants, in addition to discussing newer modalities for hemodynamic monitoring, and the role of bedside functional echocardiography in management of neonatal shock.Shock is a pathophysiologic state characterized by an imbalance between oxygen delivery and oxygen demand in the tissues leading to tissue hypoxia. The initial compensated phase is characterized by neuroendocrine compensatory mechanisms with increased tissue oxygen extraction, leading to maintenance of blood pressure (BP) in the normal range. The blood flow and oxygen supply to vital organs are maintained at the expense of non-vital organs. The compensated phase may have additional signs such as tachycardia, prolonged capillary refill time (CRT) and decreased urine output. In adults, these features are commonly seen early in compensated shock. However, these features may be missed in neonates, in whom shock is mostly recognized in the uncompensated phase. This is in part due to the lack of data on normal BP ranges that would ensure adequate organ perfusion in the premature infant. The uncompensated phase of shock is characterized by a decrease in vital and non-vital organ perfusion, which may be evident by the development of lactic acidosis. This will eventually lead to cellular disruption with irreversible damage, clinically characterized by multiorgan failure and death.Myocardial dysfunction, abnormal peripheral vasoregulation and hypovolemia leading to decreased delivery of oxygen and nutrients to tissues are often the primary sources of neonatal shock. This is often complicated by relative adrenal insufficiency often seen in the premature infant. The causes and types of neonatal shock are described in Table The neonatal myocardium has fewer contractile elements compared with older children and adults . In partThis is distinct to the myocardial dysfunction beyond the transitional period when immature myocardium may have a lesser role to play. Hemodynamically significant PDA (hsPDA) is a common cause of hypotension in VLBW infants. The presence of an hsPDA with resultant decrease in diastolic BPs can also theoretically affect the perfusion of the myocardium, which primarily takes place during diastole. However, studies show no significant change in contractility with an hsPDA , 7. On tThe vascular smooth muscle tone and its complex regulation play a key role in pathogenesis of neonatal shock. A balance of the vasodilating and vasoconstricting forces regulates the tone. These factors may involve autocrine, endocrine, paracrine, and neuronal factors. Commonly described factors include vasopressin, nitric oxide, eicosanoids, catecholamines, and endothelin –12. A kein utero blood loss such as with maternal abruption, fetomaternal or fetoplacental hemorrhage or tight nuchal cord. Postnatally, acute blood loss may be associated with gut perforation following necrotizing enterocolitis, sub-galeal bleeding or intracranial hemorrhage. In addition to this, relative hypovolemia can be seen with capillary leak and vasodilatory shock in severe sepsis.Unlike in the pediatric or adult population, hypovolemia is not a very commonly encountered etiology of shock in the first few days of life. Causes of hypovolemia in newborns would include history of Pathophysiology of shock in newborns is unique since it is associated with physiologic transition from fetal circulation to neonatal circulation at birth. Suprasystemic pulmonary vascular resistance (PVR) in the prenatal period may remain elevated, especially in the presence of ongoing hypoxia and acidosis from sepsis, leading to persistent pulmonary hypertension of the newborn (PPHN). The latter contributes to right ventricular failure, and as such may need therapies directed to decrease right sided pressures. In addition to PPHN, newborn shock may be associated with closure of ductus arteriosus in a ductal dependent congenital heart lesion, as such requiring prostaglandin infusion to open and maintain patency of the ductus arteriosus (PDA).In addition, there is plenty of evidence suggesting low cortisol levels in sick term, late preterm, and preterm infants –19. BothAdequacy of the systemic and peripheral blood flow and thus oxygen delivery to the tissues can be measured by monitoring BP, CO, and/or systemic vascular resistance (SVR).via thermodilution or pulse induced contour cardiac output, pulmonary wedge pressure, or central venous pressure providing accurate assessment in adults or older children may be impractical in the premature infant. In addition to the difficulty associated with intracardiac shunt placement in VLBW infants, the dye dilution and thermodilution methods may not be accurate due to the presence of intracardiac and ductal shunts.Direct yet invasive measures of cardiovascular function such as CO measurement It is known that BP = CO × SVR. Both BP and CO can be measured. SVR is a derived value from the above equation. CO = heart rate (HR) × stroke volume (SV). SV depends on preload, myocardial contractility, and afterload conditions. The relationship between HR, cardiac filling, and CO has been shown in Figure Measurement of myocardial contractility using load dependent measures such as fractional shortening (FS) can be affected by the right ventricular dominance characteristic of fetal circulation. Appropriate assessment of myocardial activity requires measurement of load independent measures such as relation between velocity of circumferential fiber shortening and left ventricle (LV) wall stress indices .Blood pressure monitoring, preferably measured invasively can offer continuous real time assessment of the CO . HoweverArguably, flow is a better indicator of perfusion than the BP that drives the flow to the organs. However, flow measures such as LV and right ventricle (RV) output may not be accurately depictive of organ blood flow in VLBW infants due to the presence of above mentioned shunts in the transitional period. SVC flow may be used as a valid indicator of cerebral blood flow (CBF) –34. In fThe relation between mean BP and systemic blood flow is a complex one in a VLBW infant, especially in the first few days of life. Autoregulation ensures adequate perfusion to vital organs in states of hypoperfusion. However, cerebral autoregulation may be lacking in the VLBW transitionally at birth or during period of illness. 30 mm Hg has been proposed as the cutoff for mean BP below which cerebral perfusion may not be adequate; however, this relation may not be entirely accurate given differing findings from various studies –39.Other indirect clinical measures of cardiovascular function include CRT, urine output, HR, and presence of lactic acidosis. A combination of such measures rather than individual assessments may offer increased specificity for detecting low flow states .Central venous pressure approximates right atrial pressure and can give valuable information regarding the preloading conditions and in assessing response to volume in critically ill patients. Normal numbers have been described for term and preterm infants, with special emphasis on the trend pattern. However, feasibility in the newborn infants remains questionable owing to the invasive nature of the procedure –43.2) is considered as the balance between oxygen demand and delivery and has been used as a determinant for tissue hypoxia. It is one of the key targets (in addition to the determinants for preload and contractility) for goal directed management in severe sepsis and septic shock (2) in preterm infants have also been described, but widespread use is limited by the invasive nature of the procedure along with the effect of persisting fetal shunts in the newborn period  oxygen difference is the difference between the oxygen content of the arterial and venous blood. The normal numbers would be less than 5 ml/100 ml of blood or 25%. In low output states, A-V extraction increases, decreasing the mixed venous saturation, and hence increasing the difference. In distributive shock, there is decrease in the oxygen extraction, leading to higher mixed venous saturation and hence a narrowed difference. This measure offers an excellent estimate of tissue oxygen delivery, however, limited again by the invasiveness of the procedure.Electrical cardiometry  allows for assessment of CO by measuring the changes in thoracic electrical bioimpedance caused by the cardiac cycle. It is non-invasive, easy to apply, offers continuous assessment and has recently been validated against invasive methods of CO measurements in hemodynamically stable newborns. However, more data are needed for validation in neonates with hemodynamic compromise before its widespread clinical applicability , 49.In a2) and cerebral fractional tissue oxygen extraction. Reference values with comparison with other modalities have now been published, but greater clinical applicability rests with trend monitoring rather than absolute numbers  has been evaluated as a tool to assess cerebral and peripheral oxygenation and oxygen extraction. The benefit of this modality is the availability of continuous measurement. Among other measures, NIRS can provide values for cerebral oxygenation  has been described for continuous assessment of capillary oxygen saturation in various organs , 56.UnliThe plethysmographic signal of pulse oximeter can be used to calculate ratio of the pulsatile and non-pulsatile components, described as Perfusion Index. This has been recently studied and is found to be reasonably predictive of low flow states, including patent ductus arteriosus –59. RefeFunctional cardiac MRI has been recently evaluated as an additional feasible tool to evaluate cardiac hemodynamics, especially PDA. At this time it is deemed to be an insightful research tool while awaiting more studies , 61. Rol2, signs of decreased peripheral perfusion such as acidosis, elevated lactate. In spite of being a late finding, hypotension is the most commonly used determinant of decreased perfusion in NICU given the ease of monitoring. The clinician must keep a keen eye on the other signs of symptoms discussed earlier before the “ischemic threshold” is reached for low BP. The other traditionally used clinical parameters such as clinical assessment, HR, CRT, urine output, and serum lactate are also proxy indirect markers of cardiovascular well-being. Early bedside focused echocardiography, described by some as focused cardiac ultrasound, can help in early identification of underlying pathophysiology and targeting specific therapy. The aim of focused echocardiography is not to rule out congenital heart defect but to gain physiological information which can help in delivering goal oriented time specific intervention. Interventions should be based after carefully considering the underlying pathophysiology  has been well studied in adult population. From a single center study (2001), Rivers et al. reported that EGDT provides significant benefits with respect to outcome in patients with severe sepsis and septic shock . FollowiIn contrast with adults, EGDT has not been well studied in neonates and children. The ACCM guidelines (2007) recommended its use in neonatal shock; however, there is not widespread use of EGDT in management of neonatal shock. This is partly due to the fact that septic shock accounts for only a small percentage of shock in the NICU and partly due to the lack of non-invasive measures for hemodynamic monitoring. There is limited evidence on use of central venous oxygenation, cardiac index, non-invasive CO monitoring, and assessing wedge pressure in neonates, and some the parameters are being evaluated in the research studies. However, at this stage, they have limited role in the clinical practice. Modern assessment modalities such as functional echocardiography and NIRS offer non-invasive methods of hemodynamic assessment. Functional echocardiography in particular can be a huge asset to the neonatologist in the initial stabilization and subsequent monitoring in intensive care unit. Various echocardiographic measures are discussed further in the next section.Functional echocardiography refers to bedside point of care echocardiography that can provide real time hemodynamic information by assessing cardiac function, loading conditions (preload and afterload) and CO Table . It is nAn initial comprehensive echocardiographic study can aid in ruling out a congenital heart lesion, particularly pulmonary atresia and coarctation of the aorta. In addition to ruling out CHD, this should also include assessment of the PDA with its effects on cardiac hemodynamics. The size of the duct (>1.5 mm at the point of maximum constriction is considered significant), direction of shunting, left atrial to aortic root ratio , left pulmonary artery , and pattern of diastolic flow in the post ductal descending aorta can be used to determine hemodynamic significance of a PDA. Following this, the next step in management is identification of underlying pathophysiology and categorization of shock as distributive, hypovolemic, obstructive, cardiogenic, or dissociative , and the right heart. Qualitative assessment includes “eyeballing” the heart in apical four chamber view Figure  while quSimpson’s biplane method can be used to assess the left ventricular end-diastolic area and volume. Neonates may have physiologic right ventricle (RV) dominance, but in general, the right ventricular dimensions are smaller than those of the LV. A dilated right atrium (RA) may indicate volume overloading of the right side of the heart and in the presence of bowing of intra-atrial septum toward the left atrium may indicate elevated right atrial pressure and hence pulmonary hypertension. By contrast, the triad of a “kissing” small LV cavity, RV size, and a normal or small RA is strongly associated with hypovolemia.Dmax) and minimum (Dmin) IVC diameter from the subcostal view also predicts fluid responsiveness  during inspiration and expiration has been reported to have high predictive value with a sensitivity and specificity exceeding 90% , 85. Simsiveness . IVC dissiveness .The qualitative measures include “eyeballing” of the contractility from the apical four chamber view, PLAX, parasternal short axis view (PSAX), or subcostal view Figure . This maThe above measures may be affected by load conditions. Appropriate assessment of myocardial activity requires measurement of load independent measures such as relation between velocity of circumferential fiber shortening and LV wall stress indices .Newer more accurate yet less feasible modalities at this time include speckle tracking, strain rate, and 3-D imaging.Persistent pulmonary hypertension of the newborn is a common condition in the NICU. A detailed assessment of RV function is out of scope of this review article and readers may refer to guidelines for assessment of RV function in neonates which have been published in American Journal of Echocardiography . AssessmThis can be assessed through estimation of LV and RV outputs (RVOs) and systemic blood flow measures. RVO and LV output (LVO) can be easily measured using VTI proximal to the pulmonary valve and the AV valve, respectively. The LVO is an estimation of systemic blood flow, and RVO is an estimation of systemic venous return in the absence of cardiac shunts. However, both LV and RVOs may be affected by the presence of fetal shunts. In spite of limitations such as significant intra observer variability, effect of shunts and errors with high angle of insonation, biventricular output measures are commonly used due to reliability with experienced echocardiographer. These values can be trended to follow the impact of interventions in real time. SVC flow may be unaffected by such fetal shunting and has been discussed as a surrogate for CBF but with conflicting reports about association with impaired neurodevelopmental outcomes , 34, 36.One of the challenges with using diameters of vessels in newborns is the standardization of the size based on the infant. A ratio between vessels accounts for this. As an assessment of preload to the heart, the IVC is a useful marker for adequate fluid balance. Coupled with the descending aorta in cross section, the IVC/Ao ratio is a useful age/size adjusted marker for assessing a low volume status. Chen et al. demonstrated that an IVC:Ao ratio 0.8 was associated with dehydration (86% sensitivity) in pediatric patients with gastroenteritis. Further work needs to be done in the premature newborn population but the IVC/Ao ratio may have promise as an objective measure of fluid status .Shock in the newborn period is associated with unique pathophysiologic states that need careful assessment and individualized approach for management. Early recognition of shock and its underlying pathophysiology is critical in instituting early target specific intervention, which may improve outcomes in patients with neonatal shock. A focused bedside functional echocardiography can provide vital anatomic and physiologic information to such management. Widespread use is limited because of its lack of availability, structured training programs for neonatologists and data on clinical outcomes. This modality should be further explored to generate data for therapeutic end points that can be used to standardize and protocolize the management of neonatal shock. We recommend that focused echocardiography in neonatal shock should be regarded as an extension of clinical examination and other traditionally used clinical parameters.YS and FV reviewed the literature and prepared the manuscript. YS provided images and tables for the article. AK edited the article and helped in finalizing the manuscript.The authors declare that the research was conducted in the absence of any commercial or financial relationships that could be construed as a potential conflict of interest.


Reduced lung function is common and associated with increased cardiovascular morbidity and mortality, even in asymptomatic individuals without diagnosed respiratory disease. Previous studies have identified relationships between lung function and cardiovascular structure in individuals with pulmonary disease, but the relationships in those free from diagnosed cardiorespiratory disease have not been fully explored.UK Biobank is a prospective cohort study of community participants in the United Kingdom. Individuals self-reported demographics and co-morbidities, and a subset underwent cardiovascular magnetic resonance (CMR) imaging and spirometry. CMR images were analysed to derive ventricular volumes and mass. The relationships between CMR-derived measures and spirometry and age were modelled with multivariable linear regression, taking account of the effects of possible confounders.1 and FVC were associated with smaller left ventricular end-diastolic (−5.21ml per standard deviation (SD) change in FEV1, −5.69ml per SD change in FVC), end-systolic  and stroke volumes ; right ventricular end-diastolic , end-systolic  and stroke volumes ; and with lower left ventricular mass . Changes of comparable magnitude and direction were observed per decade increase in age.Data were available for 4,975 individuals, and after exclusion of those with pre-existing cardiorespiratory disease and unacceptable spirometry, 1,406 were included in the analyses. In fully-adjusted multivariable linear models lower FEV1 and FVC are associated with smaller ventricular volumes and reduced ventricular mass. The changes seen per standard deviation change in FEV1 and FVC are comparable to one decade of ageing.This study shows that reduced FEV Respiratory disease is common and under-diagnosed, and a significant and growing cause of death and disability. Much of1) had an all-cause mortality almost twice that of the highest quintile. The risks were similar in asymptomatic  individuals.Lung function is independently associated with cardiovascular morbidity and mortality,5, both 1 to forced vital capacity (FVC) ratio less than 0.7, the generally accepted cut-off for obstructive lung disease[The prevalence of subclinical impaired lung function is great: in the National Health and Nutrition Examination Survey (NHANES) and NHANES III cohort studies of community volunteers in the United States around 1g disease. Almost 1 to FVC ratio. Barr et al.[1 to FVC ratio and left ventricular end-diastolic and stroke volumes. These analyses were conducted in an all-comers population, many of whom had existing cardiovascular and respiratory disease. Furthermore, the effects of FEV1 and FVC per se were not studied. Given the prevalence of subclinical changes in lung function, together with the associations with adverse outcomes, we set out to explore the relationships between lung function and cardiovascular structure and function in a population free from diagnosed cardiovascular and respiratory disease.Previous work has explored some of the relationships between lung disease and cardiovascular structure. Using data from the MESA cohort, Grau and colleagues demonstrrr et al. continueth June 2011 (Ref 11/NW/0382). None of the authors had direct contact with the study participants. This report is a cross-sectional analysis of the subset of participants who took part in the imaging pilot programme.UK Biobank is a large prospective cohort study of approximately 500,000 unselected community volunteers aged 40 to 69 at the time of enrolment, living in the United Kingdom. The design and conduct of the study have been described in detail previously. This stDemographics and doctor-diagnosed co-morbidities were self-reported by electronic questionnaire and interview with a healthcare professional. Data collected during the imaging visit were used in the analyses, except where unavailable, in which case data from the enrolment visit were used. If a participant did not answer a question regarding a comorbid diagnosis, or did not know, it was assumed they did not have the condition. Physical measurements  and smoking status were captured exclusively at the time of imaging.Smokers were defined as individuals who smoke, or used to smoke, on all or most days. Those who smoked occasionally were deemed to be never smokers. For current or previous smokers, pack year history was calculated as the product of the number of packs of cigarettes smoked per day and the difference between age started smoking and age stopped smoking (or age at imaging for current smokers).A subgroup of participants is undergoing cardiovascular magnetic resonance imaging (CMR). The CMR acquisition protocol and post-processing have been described previously. In brie1), forced vital capacity (FVC), and an automated assessment of measurement quality were recorded.Spirometry without bronchodilator administration was performed at the time of imaging according to a standard protocol using a Vitalograph Pneumotrac 6800 spirometer . Each pa1 to FVC ratio less than 0.7.Spirometry blows were excluded from the analysis if the automated quality assessment was anything other than ‘acceptable’. Participants were excluded if they had fewer than two acceptable blows, if the coefficient of variation for the two or three acceptable blows exceeded 5%, or if the difference between the best and second best acceptable blow exceeded 150ml, in accordance with established guidelines. To inve1 and FVC were approximately normally distributed and the assumptions for linear regression were satisfied. FEV1 and FVC were standardised to the mean .The CMR-derived parameters and FEV1 were modelled with multivariable linear regression, using age, sex, ethnicity, height, weight, systolic blood pressure, resting heart rate, Townsend deprivation index , education level , regular alcohol consumption (defined as three or more occasions per week), smoking history (pack years) and any diagnosis of hypertension or diabetes as co-variates. Height and weight were included as covariates in the regression models, rather than indexing the dependent variables to body surface area, since the use of ratios in regression analysis is liable to spurious results and misinterpretation[The relationships between nine CMR-derived parameters and both FVC and FEVpretation,16. The 1 and FVC were used as co-variates. Interaction terms were used to explore any stratification by sex in the relationships between the CMR-derived parameters and the primary exposure variable  in each regression model.To place the effects of spirometry on CMR-derived parameters in context, the relationships between the CMR-derived parameters and age were modelled in a similar fashion, except that both FEVEach regression analysis was performed on a complete-case basis without imputation of missing data. Nine of thirteen covariates had no missing data, and the covariate with most missing data  lacked fewer than one percent of observations. The outcome variables (CMR-derived parameters) had fewer than two percent missing observations, which were clustered in 22 participants who were missing all the outcome variables and another one who was missing right ventricular parameters.1 or FVC, or per decade change in age. P values were calculated using Student’s t test or ANOVA for continuous variables and Chi-squared test for categorical variables. Statistical analyses were performed using R version 3.3.2[Regression coefficients are presented as the change in the CMR-derived parameter per standard deviation change in FEVion 3.3.2.1,221 individuals were excluded on account of pre-existing cardiorespiratory disease, the definition of which is provided in 1, are described in 1 were older, more likely to be female, shorter, lighter, had lower diastolic blood pressure and higher resting heart rates, and were more likely to have obstructive spirometry.1,406 participants were included in the primary analysis. On account of missing data, the number of observations in the regression models was 1,366 or 1,367 depending on the CMR-derived parameter being studied. Baseline characteristics, stratified by tertile of FEV1 and FVC were associated with smaller left ventricular (LV) end-diastolic volume, LV end-systolic volume, LV stroke volume, right ventricular (RV) end-diastolic volume, RV end-systolic volume, RV stroke volume, and LV mass , LV end-systolic volume , RV end-diastolic volume , RV end-systolic volume , and RV stroke volume  . Values In a sensitivity analysis employing a broader definition of acceptable spirometry (at least two ‘acceptable’ blows with no restriction on the range or coefficient of variation of these blows) 2,070 individuals met the criteria for inclusion. There were minimal changes to the effect sizes in the multivariable linear regression, and the 95% confidence intervals were narrower on account of the larger sample size . There wThe key finding of our study is that in a large cohort of individuals without prior diagnosis of cardiorespiratory disease, lower lung function is associated with smaller left and right ventricular end-systolic, end-diastolic, and stroke volumes; and with lower left ventricular mass. These relationships are independent of other variables known to affect cardiac structure. There was no association with ejection fraction.1 to FVC ratio was associated with LV end-diastolic volume and LV stroke volume[1 and FVC per se were not studied. Using echocardiography, Watz and colleagues[1. Our study extends these previous findings to a cohort without diagnosed heart or lung disease, and demonstrates consistent relationships between spirometry and multiple measures of left and right ventricular structure and function. The effect sizes seen in this study (2.29ml to 5.84ml per standard deviation change in FEV1 or FVC) are comparable to those seen in MESA[The results of our study confirm and expand on previous work. The Multi Ethnic Study of Atherosclerosis (MESA) lung substudy has previously identified relationships between emphysema severity, quantified by thoracic computed tomography, and right and left ventricular volumes and mass,10,18, ake volume althougholleagues showed tn in MESA. Furthern in MESA.Age is a significant risk factor for cardiovascular morbidity and mortality,22, and 1 and FVC is comparable to that seen with ageing. Notably, the changes in CMR-derived parameters per standard deviation of FEV1 and FVC are approximately the same as those seen with one decade of ageing. This suggests lower lung function is associated with a ‘premature ageing’ effect on the ventricle. This provides potential insight into the mechanisms responsible for adverse cardiovascular outcomes in those with deranged lung function, and highlights the potential importance of measuring lung function as a marker of cardiovascular ‘ageing’ and risk.In the current study, the variation in ventricular volumes and mass seen with lower FEV1, FVC, or age, despite the significant influence of these factors on other measures of ventricular function. Similar results were found in a MESA analysis of individuals with emphysema[Our analysis did not identify any relationship between ejection fraction and FEVemphysema. Our resThis study has several limitations. The design of UK Biobank renders it liable to selection bias, and it is likely that the population recruited to the study is not completely representative of the population as a whole. The self-reporting of co-morbidities and lifestyle factors such as smoking is liable to ascertainment bias. The relatively low prevalence of smoking observed in our study population may be explained, at least in part, by these two factors.1 and FVC were represented, and thus the ability to examine relationships between these parameters and cardiac structure was preserved. Furthermore, the sensitivity analysis revealed that exclusion of these individuals did not materially affect the results, and suggests that no significant selection bias was introduced by the exclusion of individuals with unacceptable spirometry. The inclusion only of individuals with reproducible, high quality spirometry increases the confidence in the relationships identified by this study.Many participants were excluded from the analysis as their spirometry did not meet conventional criteria for reproducibility and validity. It is possible that the variability in spirometry arose from its acquisition in the non-specialist UK Biobank assessment centre, rather than in a dedicated pulmonary function laboratory. Furthermore, the study protocol limited participants to producing three blows, rather than repeating the measurement until valid and reproducible results were achieved, as is commonly done in clinical practice. Nonetheless, even after individuals with unacceptable variation had been excluded the full range of FEVThis study raises a number of questions worthy of further analysis, including the long term cardiovascular outcomes of individuals with subclinical changes in lung function, the changes in cardiovascular phenotype over time in those with deranged spirometry, and evaluation of ventricular fibrosis in participants with impaired lung function through T1 mapping. These will be amenable to investigation as longitudinal follow-up and parametric mapping data become available from UK Biobank.1 and FVC are associated with smaller left and right ventricular volumes, and lower left ventricular mass. The changes in ventricular structure per standard deviation fall in FVC and FEV1 are similar to those seen with a one decade increase in age, and may shed light on the mechanisms underlying increased cardiovascular risk in those with subclinical changes in lung function, as well as the importance of lung function as a risk factor for cardiovascular disease.In a large cohort of patients without known cardiorespiratory disease, lower FEVS1 File(DOCX)Click here for additional data file.S2 File(DOCX)Click here for additional data file.S1 Fig1 prior to the standardisation of the lung function. R2 is the coefficient of explained variance, calculated as the square of the Pearson correlation coefficient between the CMR-derived parameter and FEV1 on a complete pairs basis.Each panel shows the association between one CMR-derived parameter and FEV(TIF)Click here for additional data file.S2 Fig2 is the coefficient of explained variance, calculated as the square of the Pearson correlation coefficient between the CMR-derived parameter and FVC on a complete pairs basis.Each panel shows the association between one CMR-derived parameter and FVC prior to standardisation of the lung function. R(TIF)Click here for additional data file.


T1) and extracellular volume (ECV) offers a means of identifying pathological changes in myocardial tissue, including diffuse changes that may be invisible to existing T1-weighted methods. This technique has recently shown strong clinical utility for pathologies such as Anderson-Fabry disease and amyloidosis and has generated clinical interest as a possible means of detecting small changes in diffuse fibrosis; however, scatter in T1 and ECV estimates offers challenges for detecting these changes, and bias limits comparisons between sites and vendors. There are several technical and physiological pitfalls that influence the accuracy (bias) and precision (repeatability) of T1 and ECV mapping methods. The goal of this review is to describe the most significant of these, and detail current solutions, in order to aid scientists and clinicians to maximise the utility of T1 mapping in their clinical or research setting. A detailed summary of technical and physiological factors, issues relating to contrast agents, and specific disease-related issues is provided, along with some considerations on the future directions of the field.Mapping of the longitudinal relaxation time ( T1, and extracellular volume (ECV) in the human heart has recently shot to prominence on the merits of the modified Look–Locker inversion recovery (MOLLI) imaging sequence and related techniques [T1-weighted techniques such as late gadolinium enhancement (LGE) imaging. For focal fibrosis, LGE provides excellent delineation of lesions with some means of quantifying their volume; however, LGE does not give a T1 estimate and may not be able to identify widely distributed or diffuse myocardial diseases. For example, when LGE is applied in diffuse fibrosis, the myocardium can appear isointense and indistinguishable from normal myocardium, as multiple uncalibrated factors affect the image brightness. These are clear limitations of LGE techniques, and in such situations quantitative T1 and ECV mapping is advocated [T1 mapping is now widespread, with most MRI manufacturers offering T1 mapping solutions. However, great care must be taken when applying these methods clinically, given the need for protocol optimisation and locally-derived normal ranges.Mapping of the longitudinal relaxation time, dvocated , 3. Use T1 mapping has been the focus of several studies to date [T1 and ECV mapping that affect accuracy and precision, and this review aims to give a comprehensive description of these, along with potential solutions, with the intent of aiding physicists and clinicians to maximise the clinical utility of T1 mapping. Indications will also be given as to what is reasonably achievable with myocardial T1 mapping in specific clinical applications, without a full clinical review, for which the reader is directed to Haaf et al. [T1 mapping methods, the available pulse sequences, and the history of the technique will be discussed briefly, but again readers are directed to more-detailed reviews for full technical information: for example, by Kellman et al. [T1 mapping protocol [The accuracy and precision of myocardial  to date –7, and hf et al. , Taylor f et al. , Kammerlf et al. , and Schf et al. , amongstn et al.  and Higgn et al. . Finallyprotocol . A folloT1 and ECV estimation: first, technical and physiological pitfalls; second, issues relating to contrast agents; and third, specific disease-related issues. It will conclude with future directions of the field and some summary recommendations, including guidance on how T1 estimation accuracy might be traded for increased clinical utility. Within each section, pitfalls will be introduced, their mechanisms described, and suggestions offered for how to mitigate or eliminate them, along with possible future solutions, where available. For ease of reference, Table T1 and ECV mapping accuracy and precision.This review will focus on sources of bias and variability in myocardial T1 mapping methods currently rely on preparing the longitudinal magnetisation using inversion or saturation radiofrequency (RF) pulses, applied to the whole imaging volume. Many pulse sequences for T1 mapping can be grouped according to the magnetisation preparation: inversion recovery sequences, including MOLLI [T1 mapping (SMART1Map) [T1 (SAP-T1) [T1-weighted image acquisitions are performed during a breath-hold, over several cardiac cycles. The aim is to sample the T1 recovery over a range of delay times, and pixel-by-pixel curve-fitting is used to estimate T1 values. This produces a pixelwise T1 map, often with other output maps as quality indicators; see Kellman et al. for a flowchart illustrating the pipeline of T1 and ECV map generation [All routinely available ng MOLLI  and shorng MOLLI ; saturatng MOLLI , saturatRT1Map) , and shoSAP-T1) ; and mixSAP-T1) . In pracneration . Most MRneration , 22.T1 in the heart: the inversion-recovery cine sequence, also known as the Look–Locker cine (LL-cine) technique [T1 curve-fitting. This approach uses spoiled gradient recalled echo (GRE) cine, avoiding the factors affecting single-shot balanced steady-state free-precession (bSSFP). Note that many studies that have used the LL-cine method have repeated it at multiple washout times, as this improves ECV accuracy [Non-mapping approaches can also estimate echnique , 24, relT1 mapping methods and schemes being addressed where appropriate. A common nomenclature for T1 mapping schemes will also be adopted [Given its popularity, the original MOLLI sequence will be considered the default method in this review, with other  adopted . This noT1 estimation: namely, for sampling of the recovering longitudinal magnetisation with a series of small-flip-angle excitation pulses [T1, known as T1*, when curve fitting:S(t) is the signal at time TI after inversion. The “true” T1 is usually longer than T1* and can be calculated using the so-called Look–Locker correction:The original 3b(3b)3b(3b)5b MOLLI sequence  and subsn pulses , 26. TheT2 and, weakly, to magnetisation transfer (MT), unlike spoiled GRE; (2) a relatively large excitation flip angle, 35° at 1.5T; and (3) an intermittent sampling scheme, governed by heart rate for original MOLLI, to maximise the spread of inversion recovery times. Factors (1) and (2) lead to progressive saturation of the longitudinal magnetisation, causing a negative T1 bias even after Look–Locker correction. Furthermore, Look–Locker sets are separated by pause intervals, aiming to allow sufficient recovery of the magnetisation prior to the next inversion pulse. If any of these pause intervals are too short for the application, be it native or post-contrast, this can lead to additional bias in T1 estimates. This can be particularly problematic for fast heart rates, as well as longer T1 values, such as in native myocardium at 3T.The Look–Locker correction assumes continual repetition of a small flip angle spoiled GRE readout. MOLLI-based techniques violate the Look–Locker assumptions by using: (1) a bSSFP readout, which is sensitive to T1 mapping and a 4s(1s)2s(1s)1s scheme for post-contrast acquisitions [Bias resulting from limited magnetisation recovery can be mitigated by using an optimised MOLLI acquisition scheme: by extending the inter-inversion pause in beats or by specifying pauses in seconds rather than beats . The 201isitions .T1 recovery. The saturation recovery is still affected during each SASHA single-shot readout, but this produces negligible bias in T1 estimates obtained from curve fitting of the independent saturation-recovery images [Alternatively, saturation recovery methods such as SASHA  sample ty images , 27.T1 mapping with single-shot imaging, where in-plane spatial resolution is necessarily somewhat coarse. Furthermore, the endo- and epi-cardial borders of the myocardium are often oblique to the imaged slice, especially for planes far from the mid-ventricle or in abnormal ventricles. Partial volume can lead to substantial errors, which motivates careful ROI delineation on T1 maps [Partial volume of different tissues within a voxel is prevalent in T1 maps .T1, especially when the tissues included in a voxel have strongly different T1 values. Prominent effects occur at the endocardial border where the difference in T1 between blood and myocardium leads to overestimation of native T1 estimation in subendocardial voxels. Partial volume also occurs between myocardium and other tissues, most often fat, whose chemical shift causes variable bias in the pixel T1 [Partial volume causes bias in the apparent ixel T1 ; this haT1 mapping sequence would theoretically reduce partial volume effects, but demands longer single-shot imaging duration if no trade-offs are made, increasing the risk of cardiac-motion blurring. Parallel imaging and partial Fourier in the phase-encode direction are commonly employed to allay this problem. It is essential to avoid partial volume when drawing regions of interest (ROIs), which are often limited to mid-wall when possible [Finer in-plane resolution of the possible , 30. CarT1 mapping aims to eliminate blood partial volume for improved T1 estimation accuracy [“Black-blood” accuracy , wherebyaccuracy , and hasaccuracy . Howeverk-space image acquisition over multiple cycles is required, and is often combined with some undersampling strategy [T1.To achieve finer spatial resolution, and thus reduce the impact of partial volume in source images on pixelwise mapping, segmented strategy , 35. SegT1 mapping methods promise to reduce intershot-motion-related errors [Finally, free-breathing d errors –39. HoweT1 mapping usually assume exact inversion or saturation of the longitudinal magnetisation, or fit an extra parameter instead, reducing precision. Another approach uses prior knowledge of an inversion factor, which can be estimated from fully relaxed reference images to correct the estimated T1 [The fitting models used in ated T1 –42.B1) fields to achieve their prescribed flip-angle, so are sensitive to B1 inhomogeneity, which can be substantial across the heart, particularly at 3T [B0 inhomogeneity. Therefore, adiabatic inversion pulses or composite saturation pulses are widely used for mapping to reduce sensitivity to both B0 and B1 inhomogeneity [T2 decay during their execution and introduce sensitivity to off-resonance phase accumulation. Kellman, Herzka, and Hansen suggest the use of a relatively short tan/tanh adiabatic pulse for optimal inversion efficiency [T1 estimation, as opposed to fitting of saturation efficiency as a third parameter; and (2), reliable removal of any history effect from previous cycles by nulling the longitudinal magnetisation prior to each independent shot.Conventional RF pulses require accurate RF transmit B1) fielB0 and B1 errors, for mapping it is vital to optimise both B0 and B1 over the relevant volume. See Fig. B0 and B1 maps, in vivo. The advent of RF (B1) shimming hardware with optimised volume calibration methods has enabled substantial improvements in B1 uniformity over the heart [Although the aforementioned RF pulse designs are more tolerant to he heart  even withe heart .Fig. 2InT1 mapping methods use bSSFP to sample the recovering longitudinal magnetisation, as this method offers a higher signal-to-noise ratio (SNR) than spoiled GRE. However, bSSFP is sensitive to T2 and off-resonance, where both sensitivities are modulated by excitation flip-angle, RF pulse repetition time (TR), and magnetisation transfer [T1 when the bSSFP flip angle is higher or the TR is longer, and native myocardial T1 values are more affected than the shorter post-contrast T1 values. Off-resonance is larger at 3T, mitigated by use of a lower flip angle [T1 error through use of a shorter TR is largely limited by patient peripheral nerve stimulation [Most transfer . These fip angle , but redmulation .k-space trajectory [Another consideration with bSSFP is the transient period before the steady-state is established, which is characterised by oscillatory magnetisation, causing image artefacts. The intensity of the oscillations depends on both the bSSFP catalysation used to stabilise the signal, and the ajectory .T1 mapping, some single-shot bSSFP images at short inversion-recovery times are required for optimal curve-fitting, although for SASHA a later start is preferred, as earlier readouts have low SNR [For  low SNR . Short dT1 mapping’s limited time window before image data must be acquired, especially for centre-out k-space ordering, which is discussed later. Catalysation details may be concealed from the scanner’s user interface, and may change through software upgrades without warning. Schemes include half-alpha [T1 estimation accuracy and precision, with bias errors of the order of 5% or more for some approaches [Several catalysation or “priming” sequences are used in bSSFP to accelerate the signal’s approach to the steady state, and these are particularly important in lf-alpha , and linlf-alpha  3b(3b)5b MOLLI, reported mean (standard deviation) native T1 values of 950 (21) ms at 1.5T and 1052 (23) ms at 3T, and mean (standard deviation) ECVs of 0.25 (0.04) at 1.5T and 0.26 (0.04) at 3T [T1 at 3T can lead to insufficient longitudinal recovery between Look–Locker sets, causing T1 underestimation; furthermore, reduced myocardial T2 at 3T relative to 1.5T introduces additional negative bias due to the T2/T1 weighting of bSSFP and signal decay during preparation pulses [A large multicentre study of native 4) at 3T . The incB0 and B1 inhomogeneity at 3T can be mitigated using appropriate B0 and RF shimming, respectively. Shallower excitation flip angles can also allay these effects, if the increased SNR at 3T is traded off [Increased aded off .T1 mapping requires breath-holding to minimise respiratory motion while source images are acquired. Original MOLLI used a breath-hold duration of approximately 17 cardiac cycles [T1 estimation accuracy and precision. Often, misregistration is not readily apparent in calculated T1 maps, unless a confidence-map is provided alongside or overlaid on the T1 map. See Fig. Currently, c cycles . A shortc cycles , but it T1 mapping source images should be examined carefully for displacements, even if motion-corrected images are also available. Mislocated tissue in or through the selected slice, for any of the shots, may preclude correct mapping of localised disease, such as myocarditis. If significant displacements are found, nonrigid registration may register most myocardial pixels [T1 mapping is challenging due to the large image-contrast variations between source images, and can be unreliable when tissues are imaged near the null point of longitudinal magnetisation recovery. While there are strategies for dealing with this issue [T1 map. This issue is another reason why midwall, septum-only ROIs tend to be more reliable.Post-acquisition quality control has some aspects in common for respiratory motion and cardiac misgating or arrhythmia: l pixels , 65, butT1 mapping acquisitions have been reported [Free-breathing reported , 36–39, reported , the impT1 map calculation.Source images for mapping must be acquired in the same phase of the cardiac cycle to ensure registration for pixelwise Some variation in R–R interval is normal, and MOLLI-variant sequences record the real-time R–R increments to the inversion-recovery time during each Look–Locker set. Arrhythmia may be tolerable provided the trigger to the pulse sequence is followed by a reasonably normal ventricular contraction and diastolic pause (diastasis). If arrhythmia interferes with diastolic timing, it is feasible to acquire the single-shot images in end-systole , 67. TheT1 has been reported for mapping at end systole versus end diastole [T1-relaxation process is sensitive to the myocardial relaxation or contraction state at the time of the image.A slightly shorter native diastole , 67; howdiastole . The carT1 mapping. Again, an elastic image registration algorithm may be able to account for this, but source images should be checked rigorously. If mis-registration goes uncorrected, distortion of fitted T1 recovery curves is likely, particularly in the subendocardium. For this reason, some types of cardiac arrhythmia can be a major problem, but novel methods promise robust performance in such conditions [Similar to respiratory motion, cardiac mistriggering causes mis-registration of source images for nditions .The single-shot imaging duration should not exceed the length of the cardiac pause, be it diastolic or systolic. Tong et al. empirically estimated that the shot duration should not exceed 150 ms for minimal cardiac motion artefacts . It is rT1 mapping acquisition is that the longitudinal magnetisation may not have recovered before the first inversion [The image-readout duration can also be reduced using parallel imaging methods that acquire coil profiles separately: namely, before the scan or immediately after the last single-shot image. Note a related pitfall with coil profiles or any other prescan applied immediately prior to a nversion , though T1 is important for the calculation of ECV. Complications in Look–Locker correction arise for blood that is at least partially replaced by fresh wash-in in the image slice for each cardiac cycle. Completely “fresh” blood, while still acted upon by the initial non-selective inversion, has not experienced previous shots of the current Look–Locker set since inversion and, therefore, does not require Look–Locker correction. However, even a normal heart ejects only around 55–75% of left ventricular blood per cycle [T1 mapping acquisition the arriving blood may have experienced a distorted magnetisation preparation at some upstream location, despite application of optimised preparation pulses. This can occur in short bore scanner systems [Measurement of blood er cycle , so the T1 in the myocardium and blood pool, it is also important to consider the MT phenomenon, which has been shown to influence the accuracy of T1 mapping [T1 mapping, the bound pool is mostly unaffected by the inversion pulse, so exchange during the long inversion-recovery delays distorts the shape of the T1 recovery curve and introduces T1 underestimation. The extent of this effect varies between tissues, and is substantially smaller in blood than in myocardium [When estimating  mapping , 47. Exc mapping . For invocardium , 74.T1 estimation precision. It can also be mitigated in inversion-recovery T1 mapping by use of lower-flip-angle excitation pulses and a longer imaging TR. Alternatively, the MT effect could be exploited in native T1 mapping for greater disease discrimination in pathologies such as myocardial infarction (MI), ischaemia, and iron overload—all of which have all demonstrated MT.The MT effect can be allayed by using SASHA with a three-parameter curve-fit, at the expense of reduced T1 and ECV mapping, increasing bias, scatter, or both. Cardiac and respiratory motion are particularly problematic, and strict quality control routines may help correct these where possible. At 3T, inhomogeneous B0 and RF transmit fields also become prominent sources of error, which may be mitigated with appropriate B0 shimming and RF transmit calibration. However, despite these prominent pitfalls, no issue dominates, and with considerable expertise, care, and attention to multiple factors, users can mitigate many sources of error, with the level of optimisation depending on their specific clinical or research applications.Each technical and physiological challenge listed here may cause errors in T1 mapping are subject to their own issues and controversies for deriving estimates of myocardial ECV. This section discusses the various assumptions made about GBCAs in T1 and ECV mapping.Gadolinium-based contrast agents (GBCAs) in myocardial r1, in myocardium and blood pool:Several basic assumptions regarding GBCA estimation of ECV are stated here first, with further details later. Strictly, we assume that the contrast agent has identical relaxivities, re units , 76. Theλ [If we assume that cient, λ , 77.T1 range late after the myocardial first-pass.Secondly, we assume that the GBCA does not enter myocytes or blood cells, and that it is instead in dynamic equilibrium of water relaxation inside those cells, because of fast-exchange of water through cell walls, as follows:T2), and is assumed to be highly permeable to interstitial fluid, including the GBCA. If the GBCA did not enter the collagen volume, but achieved relaxation equilibrium with it by fast exchange, it would manifest as an abnormally low ECV, resembling myocyte hypertrophy.Both amyloid deposition and collagen accumulation in fibrosis increase the interstitial space and break up myocyte packing. Collagen itself is of negligible non-permeated or “dark” volume , gadobenate dimeglumine (Gd-BOPTA), and gadobutrol; these are known under the trade names “Magnevist”, “Multihance”, and “Gadovist”, respectively. Each agent has differing relaxivities and binding properties, which can lead to differences in the estimated ECV; further complications arise due to relaxivity variations with different field strengths.Several GBCAs are available for T1 values around 15 ms lower than Gd-BOPTA, with no statistically significant difference seen in blood pool [R1 of GBCAs in human blood plasma for different contrast agent types and at different field strengths [The Gd-BOPTA agent’s aromatic ring enables weak plasma protein binding, leading to a lower molecular tumbling rate and thus a longer rotational MR correlation time and higher relaxivity in blood plasma and myocardial interstitial fluid compared to Gd-DTPA and gadobutrol. Furthermore, a lower dose of Gd-BOPTA has been shown to have similar diagnostic efficacy to a higher dose of Gd-DTPA in LGE imaging of MI . Kawel eood pool . This retrengths , 82, whiT1 mapping. If investigators plan to use one of these agents, they should do so consistently, and report this clearly in any inter-site comparisons.The relatively greater presence of albumin in blood versus myocardium means the distribution of protein-bound contrast agent between these pools is likely to be different than for non-protein-bound equivalents, altering the ratio of the change in relaxation rate of myocardium and blood and thus altering ECV. Given that this distorts one of the assumptions of in gadolinium-based ECV estimation, use of a protein-bound contrast agent will slightly modify partition coefficient estimation by T1 mapping acquisitions can be acquired during GBCA washout for improved accuracy in estimating the gadolinium blood-myocardium partition coefficient [R1 versus blood R1. The partition coefficient has been shown to deviate from this model in the early washout phase [R1 values greater than 4 s−1; these data are often excluded from ECV calculations to avoid bias .The method of contrast administration also introduces variability to ECV calculation. Under most conditions, the two-compartment steady-state assumption, stated in Eq. , holds tfficient , 88, 89,ut phase , 91, cauT1-mapping situations, it appears that the fast-exchange assumption is sound.Equations  and 8)  8) are dT1 mapping [T1 and blood-analysed haematocrit. Support for this approach is spreading [The blood for haematocrit assessment should be taken contemporaneously with  mapping , to avoi mapping ; it is cpreading , 94; howpreading , 96.T1 scans, will improve ECV accuracy and precision, but such a protocol is difficult to fit into busy clinical schedules.Although there are multiple issues with contrast agent types and field-strengths, none of these appear to dominate. There perhaps remains a dilemma between optimal ECV assessment and clinical feasibility: for example, the use of multiple acquisitions during GBCA washout, along with multiple averages for native The aim of this section is to highlight the impact of the above topics in specific clinical applications, with examples and possible solutions.T1 related to sex and age, though there is currently no consensus on whether these also influence ECV [T1 and ECV might increase or decrease with age, but the debate over these issues is outside the scope of this review. We should, however, point out that these changes are small, and demonstrable only over large groups, with a similar scatter to T1 and ECV estimates in diffuse fibrosis, which are discussed later.There appear to be subtle differences in native myocardial ence ECV , 97–99. T1 measurements in the basal, mid, and apical regions of the left ventricle in healthy volunteers [T1 values in the lateral wall versus the inter-ventricular septum [B0 distortion, and lower coil sensitivity all reduce accuracy and precision in the lateral wall. ECV, on the other hand, does not differ significantly between the lateral wall and septum [T1 by off-resonance is likely the main source. Figure T1 throughout the heart; lower T1 values are seen in inferolateral segments, which typically demonstrate off-resonance due to interfaces with the lung and the posterior vein of the left ventricle [Regarding myocardial region, there appears to be no statistically significant difference between native lunteers . Severalr septum , 97, 101entricle , 103.FigT1 and ECV mapping has attracted a lot of interest for both clinical and research applications due to its potential for accurate and precise tissue characterisation on a pixel-by-pixel basis. While T1 mapping still shows promise for aiding precision medicine and influencing management of individual patients, recent work has been more pragmatic, with applications focusing on large patient populations and specific conditions.Myocardial T1 and ECV globally, with minimal overlap with healthy ranges [T1 mapping can also be helpful in highlighting abnormal GBCA washout kinetics, which show a specific pattern in amyloid patients [T1 mapping to deliver strong diagnostic and prognostic utility.Amyloid is a relatively rare multi-system condition where deposition of misfolded fibrillary protein in tissues and organs can cause expansion of the myocardial extracellular space and impairment of cardiac function . This sty ranges , making y ranges . Post-copatients , and a hpatients . With laT1 compared to healthy volunteers, and thus represents a strong application of native T1 mapping [T1 in AFD likely does not result directly from the short T1 of fat [T1 increases sometimes seen in fatty infiltration of chronic MI, which is discussed later.Another rare multi-system syndrome, Anderson-Fabry disease (AFD) is characterised by intracellular accumulation of glycosphingolipids. This leads to progressive cardiac, renal, and cerebrovascular disease , and thu mapping –111, aga1 of fat , contrasT1 mapping [T1 maps due to increased interstitial fluid content; indeed, native T1 mapping has higher diagnostic accuracy for identifying oedema than T2-weighted imaging in Takotsubo [T1 is strongly increased, this is often sharply localised in myocardium, and thus operators should take care to localise to the relevant myocardial region. Furthermore, understanding of localised technical pitfalls is also valuable: such as off-resonance, whose impact is further modulated by B1 changes over the heart. It should be noted that published studies have typically excluded patients with major epicardial coronary disease or past MI. For T1 mapping to become clinically meaningful in myocarditis and Takotsubo, it should not only positively confirm the diagnosis, but also rule out acute MI.Myocarditis and Takotsubo cardiomyopathy are also potential clinical applications for  mapping , 113, beakotsubo , and myoakotsubo . In thisT1 and ECV currently demonstrate limited sensitivity, but may ultimately be of clinical utility if their precision is improved, or if confounds are addressed.In some pathologies, T1 on native T1 maps [T1 mapping pitfalls in acute MI, additional to that of local disease. Firstly, microvascular obstruction, or the “no reflow” phenomenon [Like myocarditis and Takotsubo, acute MI also presents with oedema, which leads to elevated local T1 maps –118. Theenomenon , causes enomenon .T2, estimated T1, and true T1 [T1 mapping, and thus CMR studies should be timed appropriately after reperfusion therapy [Secondly, myocardial haemorrhage often occurs concomitantly with microvascular obstruction , and is true T1 . Clearly therapy .T1 mapping, chronic MI presents the challenge of lipomatous metaplasia, which affects around 24–47% of MI patients [T1 mapping.For patients , 123. ThT1 recovery curve. Attempting to fit these data with a monoexponential curve typically leads to lower myocardial T1 estimates due to inclusion of short T1 lipids, which have T1 values around 370 ms at 1.5T and 450 ms at 3T [T1 mapping the mono- versus bi-exponential “model mismatch” issue is usually not the dominant factor, because most T1 mapping uses bSSFP, in which fat and water are typically out of phase due to the frequency offset of fat [T1 occurs; higher fat fractions lead to undefined T1 estimates; and fat-like T1 values are seen only at the highest fat fractions. In chronic MI, fat fractions typically do not exceed 35% [T1 estimation bias is expected, assuming fat and water are out of phase. This reduces the specificity of native T1 mapping in chronic MI, because similar, but genuine, changes occur in oedema or inflammation. It is currently unclear what effect fatty changes have on ECV. See Fig. T1 map acquired in a chronic MI patient with lipomatous metaplasia.Ignoring bSSFP characteristics, a mixture of in-phase lipid and water signals within a voxel produces a biexponential ms at 3T . Howevert of fat . Complext of fat , 126, anceed 35% , and thuT1 mapping might only reliably be reduced by replacing bSSFP with spoiled GRE imaging, in which the fat phase-difference depends purely on the TE. However, the reduced SNR of spoiled GRE relative to bSSFP would have to be considered. Alternatively, quantitative myocardial fat-fraction mapping could be applied [T1 mapping in skeletal muscle [The complex impact of fat partial volume in  applied , as recel muscle  and the l muscle .T1 mapping [Several pathologies cause diffuse fibrosis of the myocardium, such as hypertrophic and dilated cardiomyopathies (HCM and DCM), atrial fibrillation, aortic stenosis, heart failure with reduced or preserved ejection fractions (HFrEF and HFpEF), congenital heart disease, hypertension , and dia mapping .T1 values are increased in diffuse fibrosis, but not to as great an extent as in oedema or amyloidosis, and ECV is also slightly elevated. In general, scatter in estimated T1 values seems to be of a similar order to that seen in diffuse-fibrosis changes, and this has delayed the wider uptake of T1 mapping. Several investigators have shown the usefulness of ECV as a prognostic indicator [T1 [In general, native ndicator , 134, bundicator , and evetor [T1  Fig. 8)T1 valueT1 and ECV values measured in controls versus those measured in patients with likely diffuse fibrosis. This is partly due to the small changes seen in diffuse fibrosis, particularly early in the disease when reversing it would be of great clinical benefit before irreversible damage occurs to the myocardium. It is also related to the many sources of dispersion in T1 parameter estimation. As yet, it appears that there is no one factor that can be adjusted to achieve adequate T1 estimation precision for detecting early diffuse fibrosis. Many recently reported clinical studies naturally retained older T1 mapping protocols, due to the constraints of their study length, or follow-up periods in the “prognosis” papers. Despite these problems on the individual level, which may be overcome by stricter control of errors, T1 mapping of diffuse fibrosis offers concrete benefits in large-scale studies: offering a means of testing treatment effects and characterising differences on a population level [Clearly there is substantial overlap between native on level , 137.T1 and ECV are important clinical measures that allow us to characterise the myocardium in a fashion complementary to LGE. The substantial overlap of these measures between patients and controls in some cardiac conditions offers challenges to the clinical use of T1 mapping, at least with current methodology. However, in specific conditions, such as amyloid and Anderson-Fabry disease, T1 mapping has an important diagnostic and prognostic role, and is being widely adopted into routine clinical use.Both native T1 and ECV mapping have been discussed throughout this review. We will now highlight several promising areas of development that may determine the future directions of the field.Potential future solutions for errors in T1 mapping would appear to offer major gains with regards to T1 mapping accuracy and precision, as well as enabling scanning of patients with compromised breath-holding. There are already several publications demonstrating the benefits of free-breathing T1 mapping [Free-breathing  mapping –39; howeT1 mapping have incorporated simulations to improve the accuracy of bSSFP MOLLI [T1 mapping [Certain recent implementations of myocardial FP MOLLI , 139, to mapping , and to  mapping . As yet, mapping .T1, T2, and proton density mapping of the myocardium is also possible [T1, T2, and proton density maps [T1 mapping and requires long computation times for pattern matching.Simultaneous possible –144, witpossible –144. Magity maps , as wellity maps ; howeverT1 estimate in a local normal range may necessitate high precision, but not high accuracy. Indeed, accuracy may be sacrificed deliberately to yield T1 estimates that are better able to discriminate normal tissue from pathology. For example, increasing the MOLLI excitation flip angle to 50° increases sensitivity to off-resonance, which reduces T1 accuracy, but also increases MT effects, which differ between tissue types. This appears effective in detecting diffuse fibrosis in the septum [T1 is less relevant if local normal ranges are used, as they should be for any T1 mapping implementation, under current guidelines [Placing a particular e septum , 147. Thidelines .T1 and ECV as clinical biomarkers using normal ranges that are transferable between sites and vendors [Conversely, accuracy is important for establishing myocardial  vendors .T1 and ECV mapping, and it should be clear from these considerations that users should take great care when deviating from manufacturers’ advised T1 mapping protocols. While improving the apparent quality of maps and source images, users may inadvertently reduce the precision of T1 estimates, damaging clinical utility. Whatever T1 mapping setup is used, it is essential that its performance is characterised in local normal ranges, and that it is applied only for those clinical questions that its precision can support. Ongoing quality control and reassessment is also required to ensure a local normal range remains valid; the reader is referred to the upcoming SCMR and EACVI consensus for recommendations in this regard.This review has discussed reasons for inaccuracy and imprecision in T1 and ECV mapping methods show utility in groupwise comparisons through to individual clinical tests. For some conditions, like diffuse fibrosis, mapping methods serve as weakly prognostic biomarkers that might be beneficial in combination with other diagnostic information about an individual patient. In other, albeit quite rare, conditions native T1 and ECV mapping can provide strong diagnostic data.In the research and clinical applications described here, current native Fundamentally, cardiac mapping methods have not changed radically since Messroghli et al. first introduced MOLLI ; howeverT1 estimates is due to physiological differences in true T1, or how much might be eliminated if the potentially correctable issues discussed here could be addressed robustly. If these pitfalls can be accounted for simply, quickly, and reliably, without need for specialist attention, T1 and ECV mapping may ultimately support more widespread clinical applications.It remains unclear how much of the scatter observed in