Computational models have become essential in predicting medical device efficacy prior to clinical studies. To investigate the performance of a left-ventricular assist device (LVAD), a fully-coupled cardiac fluid-electromechanics finite element model was developed, incorporating electrical activation, passive and active myocardial mechanics, as well as blood hemodynamics solved simultaneously in an idealized biventricular geometry. Electrical activation was initiated using a simplified Purkinje network with one-way coupling to the surrounding myocardium. Phenomenological action potential and excitation-contraction equations were adapted to trigger myocardial contraction. Action potential propagation was formulated within a material frame to emulate gap junction-controlled propagation, such that the activation sequence was independent of myocardial deformation. Passive cardiac mechanics were governed by a transverse isotropic hyperelastic constitutive formulation. Blood velocity and pressure were determined by the incompressible Navier-Stokes formulations with a closed-loop Windkessel circuit governing the circulatory load. To investigate heart-LVAD interaction, we reduced the left ventricular (LV) contraction stress to mimic a failing heart, and inserted a LVAD cannula at the LV apex with continuous flow governing the outflow rate. A proportional controller was implemented to determine the pump motor voltage whilst maintaining pump motor speed. Following LVAD insertion, the model revealed a change in the LV pressure-volume loop shape from rectangular to triangular. At higher pump speeds, aortic ejection ceased and the LV decompressed to smaller end diastolic volumes. After multiple cycles, the LV cavity gradually collapsed along with a drop in pump motor current. The model was therefore able to predict ventricular collapse, indicating its utility for future development of control algorithms and pre-clinical testing of LVADs to avoid LV collapse in recipients. The optimal function of the heart relies on multiphysics phenomena, comprised of electrical activation, as well as myocardial contraction and blood hemodynamics. A deficiency in any of these physics will impact the others. As such, development of cardiac medical devices has to consider these multiphysics interactions during both pre-clinical and clinical studies. Multiphysics cardiac computational models can be a test-bed for early ideas to assess how therapeutic interventions could impact the diseased heart. This is especially beneficial to reduce the risk to patients with rare medical conditions, where the number of subjects that can be recruited for device clinical trials is limited. Results from computational simulations could be used to generate more plausible hypotheses for future testing; thus, helping to reduce development and technology transfer costs and potential clinical adverse reactions.Numerous studies have modeled the electromechanical interactions in the myocardium with varying levels of complexity  models of the ventricles are important in studying fluid loading on the myocardium as well as the heart's blood pumping efficiency. For the sake of simplicity, the contribution of electrical activation on wall mechanics was not considered in many ventricular FSI models. Instead, a temporal activation function was used to govern contraction uniformly across the entire left ventricle (LV) , classified as a Class III device by the United States' Food and Drug Administration (FDA). As it may pose a high risk to patients, the device needs to be rigorously tested prior to marketing approval. Whilst the LVAD works well for patients with the most severe heart disease, excessively high pump flow rate relative to ventricular filling rate can cause the ventricles to collapse. This phenomenon is known as ventricular suction and it represents a common complication  and (5) was chosen to vary linearly such that aepi is 0.12 and aendo is 0.07 to allow the myocardium to relax in the opposite direction of its activation  formulations Equations (1)–(5) were based on the phenomenological Nash and Panfilov  model, ma and k2 were adjusted to achieve a three-fold increase in conduction velocity and a 2-fold increase in AP upstroke velocity relative to their respective values in the myocardium .Membrane potential dynamics and AP propagation within the myocardium were governed by Equation (1), and within the Purkinje fibers by Equation (2), based on the modified cable equation developed by Alqahtani et al.  to take X operator in Equations (1) and (2) is defined within the material frame  such that σ, was set to exhibit an anisotropy ratio of 4:2:1 along the fiber, σ, was determined from conductivities defined in the local fiber (σf), sheet (σs) and normal-to-sheet (σn) directions according to:The ∇s et al. . The conf was set such that the epicardial activation time matched reported ex vivo experimental findings in human heart :The simplified form of active stress proposed by Nash and Panfilov , and furkTa controls the maximum magnitude of Ta. A, and B are fixed parameters introduced here to give units to the originally dimensionless formulation. The delay function, ϵ(Vm), was given by Göktepe and Kuhl , based on the formulation of Doll and Schweizerhof ψ=ψiI1 in Equation (10) denotes the first invariant of the isochoric elastic right Cauchy-Green tensor, C, whilst J in Equation (12) represents the determinant of the deformation gradient tensor, F. In order to ensure material stability, If4 was set to zero when If4 < 0, assuming that myocardial fibers do not contribute significantly to passive mechanics during compression :s−1 and β = 0.01 s.where z et al.  such thaTa was added to the 2nd Piola Kirchhoff stress, T, along the fiber, sheet, and normal-to-sheet directions, as shown in Equation (14) such that the magnitude of active stress in the sheet and normal-to-sheet directions was 40% of its value along the fiber direction  is the Green-Lagrange strain tensor whilst ψ is the myocardial strain-energy function.For simplicity, the ventricular base, including the valves, was assumed to be fixed to avoid translation and rotation. The endocardial boundary was also subjected to pressure and viscous stress from the blood fluid, discussed further below. The remaining myocardial boundaries were allowed to move freely.p, in the blood cavity were determined using the incompressible Newtonian Navier-Stokes formulation (Equation 15) and the continuity equation (Equation 16).The fluid velocity, x in Equations (15) and (16) describes the spatial derivative in the spatial frame such that ∇s, a sum of pressure and viscous stress, was applied along the normal of the endocardial surface (Equation 18):No-slip boundary conditions were applied at the ventricular base surface except at the inlets and outlets. FSI was implemented as a strongly coupled two-way framework where (1) the fluid velocity at the endocardium was set to be equal to the endocardial surface velocity (Equation 17), and (2) the fluid total stress, s to the solid formulation in Equation (13), a transformation from the spatial frame (i.e. the deformed frame where the Navier-Stokes equations are solved) to the material frame  is needed, as shown in Equation (19), where dv and dV are mesh element scale factors in the spatial and material frame respectively.To couple S denotes the traction from the fluid applied to the endocardial wall in material frame coordinates. The scale factors simply scale the local mesh element coordinates to the spatial and material frame coordinates.Here, meshψ, within the blood domain, Ωf:The fluid boundary deformation was governed by the myocardial structural deformation, Jmesh and I1,mesh, were given by Equations (21) and (22):The invariants, C1 and κmesh are artificial shear and bulk moduli, both set to a value of 1, whilst x represents the spatial frame coordinates and X the material frame coordinates. The non-linear stiffening parameter C2 was set to 0 for the healthy simulation and 100 for the LVAD simulations to cope with greater mesh deformation subsequent to cannula inclusion. Boundary conditions for Equation (20) were set to equal the displacement of the endocardial wall, where Qmi was applied to the mitral inlet, Qao to the aortic outlet, Qtri to the tricuspid inlet, and Qpa to the pulmonary artery outlet.The finite element model was coupled to a closed-loop Windkessel circulation Figures , based oThese laminar inflow and outflow boundary conditions were implemented by appending a fictitious tube to each boundary, and applying:x,t is the tangential derivative in spatial frame at the inlet or outlet surfaces. This equation essentially determines the velocity applied at the boundary, assuming the inlet and outlet boundaries are connected to a tube of Lext external length. A constraint, pext, was applied to ensure that the flow rate through the tube matched the assigned flow rate from the above. To ensure a laminar profile and prevent back-flow when the valve is closed, Lext needs to be sufficiently large to eliminate numerical backflow was set to 20 m in our model. It should be noted that the Lext does not represent an actual physical tube length, but a value chosen to satisfy the conditions.where ∇Vlv,∞ and Vrv,∞, respectively. The filling rate for each ventricle was calculated using Equation (24) based on a proportional controller:To begin, the Windkessel circuit was decoupled from the heart and the LV and RV chambers were inflated until they reached their end diastolic volumes, kflow is a fixed parameter set to 200,000 s−1.where Using the filled ventricle, the Windkessel circuit was connected and the model was solved using the initial Windkessel volumes listed in the A left ventricular assist device (LVAD) model was attached to the biventricular model via a cannula.kTa parameter of the standard model to half its value, in accordance with isometric twitch measurements in isolated epicardial tissue obtained from DCM patients , a common etiology of LVAD recipients. The geometry was modified by increasing the LV short axis and reducing its wall thickness as shown in Figures I, needed to maintain the pump motor speed, ω:Motor windings electrical equation, describing the current, Vpump, was controlled using a proportional controller, which maintains the pump speed at the desired speed, ωset:The voltage, Te, produced by the pump to the input flow rate, Q (L min−1):Electromagnetic torque transfer equation, relating the electromagnetic torque, P, produced by the pump to the input flow rate and pump speed:Pump hydraulic equation, relating the differential pressure, ΔPump inflow rate was obtained via the equation:The LVAD was represented by an ODE model that relates the hydraulic state experienced by the pump and pump motor current. The LVAD model, developed by Lim et al. , consistRin and Rout, were set to be proportional to the pump flow rate to account for the turbulence effect in the pump . This difference is the pressure head of the pump, ΔP, minus the pressure loss across the cannula. Equations (25)–(29) describe the relationships that determines the overall pump flow rate, Q, using this pressure difference.The LVAD equations were linked to the ventricular model via the numerical difference between LV pressure averaged at the cannula outlet boundary, E = 170 GPa, ν = 0.28, ρs = 2,329 kg m−3), was held fixed in space at its external parts outside the LV.A cylindrical generic pump cannula, 1.6 cm inner diameter and 0.2 cm wall thickness, was inserted at the LV apex of the biventricular geometry as shown in Figure Contact modeling between the LV endocardium and the cannula is necessary to prevent mesh collapse when surface contact occurs. In essence, the contact formulation applies loads that are inversely proportional to the distance between contacting surfaces when the surfaces move closer than a certain threshold. A wall distance formulation , was applied along the normal direction of the endocardial surface to counter its movement toward the contact destination surfaces:A boundary load Gcontact is the reciprocal of distance between surfaces, whilst Gcontact,lim is the set threshold (150 m−1). kcontact was set to 103 N mm−1, which was sufficient to prevent the wall contact. To enhance the blood domain mesh deformation with the added cannula, parameter C2 in Equation (20) was set to 100, introducing a non-linear stiffening to the mesh smoothing formulation.The dilated heart model was run for three cycles without the LVAD cannula added.The LVAD was connected to the dilated heart model at the low-end speed of 2,100 RPM and run for three stable cycles.The model was run for half a cycle (systolic) at 2,100 RPM, followed by a 300 ms duration where the pump speed was linearly ramped up to 150% of its original speed . This speed was held for three cycles, and the impact of this increased speed on the heart was observed. This was performed to determine whether aortic ejection ceased at high speed, as observed by Lim et al. in their simulation and animal experiments Lim et al. .Rmi by 25-fold and then running the model for three more cycles.To simulate suction or ventricular collapse, the LV preload was reduced significantly. Starting from the fourth cycle at 3,150 RPM pump speed, mitral stenosis was induced by increasing mitral resistance, The mesh was constructed using a combination of tetrahedral, pyramidal and prism elements. A fluid boundary layer of two elements was placed at the fluid domain side of the fluid-structure interface. The outlet and inlet surface meshes were refined by setting their edges to an average element size of 0.1 cm. The resulting normal mesh is shown in Figure CqCharacteristic quantities, Relative root mean square, Mesh convergence analysis was performed for the biventricular model using the following quantities:mn quantities represent the local electrical activation time, myocardial displacement and fluid pressure at the times of peak LV ejection and peak LV filling rate at a set of 40x40x40 material datapoints generated across the biventricular structure. The magnitudes of the first derivatives of Vm, p were extracted, and electrical activation time was determined by finding the time of maximum first derivative of Vm at each point.where the ζThe physics involved in the model are summarized in Figure S, applied as a boundary load for the myocardial mechanics. Deformation of the mesh in the blood domain was governed by the hyperelastic formulation (Equation 20), with boundary conditions given by the endocardial displacement (Equation 13) while the basal plane was held fixed.The fluid physics (Equations 15–16) was solved in the spatial frame within a moving mesh. The fluid-solid coupling occurs at the endocardial surfaces (Equations 17–19), with the fluid velocity set to the endocardial wall velocity and fluid total stress, R, variable as well as for all the solid mechanics formulations. All formulations were solved simultaneously and all variables were updated at each time step using a second order Backward Differentiation Formula (BDF) adaptive time-stepping method with an event detection algorithm implemented to ensure the BDF solver did not miss the stimulus events. Simulation results were output at 2 ms time steps.We utilized COMSOL Multiphysics 5.2a  for generating the geometry and mesh as well as solving all formulations. Fully-coupled PARDISO direct linear solver with an automatic damped Newton method were implemented, where all formulations were assembled in a single matrix and solved simultaneously together with preordering algorithm set to nested dissection multithreaded. Streamline diffusion and crosswind diffusion artificial stabilizations of the Navier-Stokes equations were applied, which enabled same-order velocity-pressure numerical integration and eased the requirement for a very fine fluid domain mesh. We adopted P1-P1 linear Lagrange elements for the Navier-Stokes equations and quadratic Lagrange elements for the membrane potential potential and recovery, Electric activation, myocardial motion, and blood velocity vector streamlines are shown in the snapshots of model behavior in Figure The model started from the end-diastolic volume (EDV) phase, where vortices from the previous filling phase still persist in both ventricles. The stimulus began at the basal entries of the Purkinje fibers. As the myocardium began to electrically activate, a torsional motion was noted. The blood flow vortices began to diminish and the flow was gradually directed toward the outlets until ejection was initiated. As the heart reached total electrical activation, blood flow streamlines were oriented toward the outlets in near straight lines. The apex also moved toward the base during contraction. The straight streamlines gradually diminished as the heart entered isovolumic relaxation.AP relaxation first occurred in the epicardium, whilst the endocardium repolarized last: the opposite of the activation direction. The filling was initiated even before the heart contraction fully relaxed. As the ventricles were being filled, blood vortices formed near the inlets. The vortices grew larger and moved apically. LV vortex shape was somewhat symmetrical whilst the RV displayed a more elongated and crescent-like shape due to the RV geometry. At the end of the cycle, the ventricles returned back to their starting EDV shape with larger vortex presents within the central region of the ventricular cavities. Smaller vortices can also be noted in the lateral regions nearing the myocardial wall, near the inlets.Following external stimulation of the basal entries of the Purkinje network, the AP propagated through the branching fibers toward the apex. The Purkinje fiber network took 34 ms to be wholly activated. As the AP traveled through the PMJ edges the myocardium was excited. The site of first myocardial activation occurred at the septal region 20 ms after the Purkinje base was stimulated.LV activation was initiated at three endocardial sites that merged into a single wavefront, which propagated radially toward the epicardium before finally reaching the base Figures . EpicardThe action potential and active stress profiles are displayed in Figures The fiber, sheet, and normal-to-sheet components of the 2nd Piola-Kirchoff stress and Green-Lagrange strain tensors were extracted from three material points, located at the epicardium of the LV free wall, septum and RV free wall. These points were taken at midway between the base and apex.To assess regional work, stress-strain loops were generated Figure  as in RuTorsional motion was assessed at three points in the myocardium: LV free wall, septum and RV free wall. These material points were positioned near the apical region at three quarters of the apicobasal distance. All points were placed at the epicardial surfaces. Torsion was mostly predominant in the LV free wall, with 16.5° counterclockwise motion when viewed upward from the apex, as shown in Figure Table −1, whilst the pulmonary velocity was 2.94 m s−1, as shown in Figure −1 during peak filling rate, whereas the tricuspid inlet velocity was 1.07 m s−1.The pressure-volume (PV) loop plots in Figure During peak ejection rate, a pressure gradient existed predominantly at the outlets Figure . The pref is the blood density, vf blood velocity, and dVfΩ the ventricular fluid cavity. The mean kinetic energy KE¯The average vorticity, The V Figure : (1) durCq, of the mechanical and fluid dynamics, as summarized in Table Mesh sensitivity analysis showed good convergence (< 5%) for global characteristic quantities, The model simulated the general impact of LVAD on the failing heart as well as, the impact of pump speed, reproducing ventricular collapse, when mitral stenosis was introduced.By reducing LV contractility and inducing heart dilation, the LV ejection fraction (EF) dropped to 16.5% and the LV end-diastolic volume (EDV) increased to 299 ml. EF was also reduced in the healthy RV to 40%. The LV pressure, along with aortic pressure, dropped to a peak systolic pressure of merely 89 mmHg, whilst the RV peak systolic pressure reached 32 mmHg.The qualitative behavior is shown in the snapshots of Figure Snapshots of the fluid streamlines in Figure Throughout the three cycles, the PV loops overlapped, indicating stability Figure . It can −1 from a baseline of 2.2 L min−1. Qualitatively, it can be seen that the pump motor current and flow rate waveforms coincided with the LV pressure waveform due to direct correlation with ΔP in Equations (25)–(28).The LVAD model also provides measures of pump electrical current, pump flow rate and speed. With the proportional controller applied, the pump speed was nearly constant with very insignificant (< 1%) drops during systolic contraction. The pump motor current showed a bump to 0.57 A from a baseline of 0.35 A, coinciding with increased pump flow rate to a peak of 6.2 L minA sudden increase of pump speed to 3,150 RPM immediately resulted in an increased aortic pressure, causing a cessation of aortic ejection in the subsequent cycles Figure . Both pe−1 to a new baseline of 7 L min−1. During systole, the pump flow rate peaked at ≈9 L min−1 as opposed to 6.2 L min−1 at the lower, 2,100 RPM, speed. Consequently, the pump current also elevated to a baseline of 1.2 A with a maximum current of 1.5 A.As shown in Figure With the cessation of aortic ejection, the blood flow streamlines during systole were directed toward the cannula at a higher velocity, as shown in the snapshots in Figure As the cycles went on, the LV pressure-volume loops gradually became more triangular Figure . The LV The stress-strain loops measured at the epicardial sites along the mid-apicobasal distance, LV and RV free walls and septum, suggest a potential change in regional work in the ventricles. The fiber components of the stress-strain loop showed a gradual shift to the left as the speed increased, especially for the LV free wall Figure . On the In the first cycle with mitral stenosis, the systolic measures did not appear to be affected by the stenosis. During diastole, the introduction of mitral stenosis severely impeded the inflow into the LV, resulting in an immediate imbalance between the pump flow rate and input flow rate into the pump. As such, the LV collapsed as the cycles continued.The major impact of mitral stenosis appeared from the second cycle onwards, in which the ejected blood was not replenished in the LV for the subsequent cycles, resulting in a lower LV EDV in subsequent cycles. Consequently, the LV pressure generated from the second cycle onwards was significantly lower (52.9 mmHg) than prior to mitral stenosis (86.6 mmHg). The LV pressure dropped even further in the third cycle to 30 mmHg, nearly similar to the RV pressure, as shown in Figure The collapsing LV also resulted in the LV PV loop to be smaller than the RV Figure . The LV The snapshots in Figure −1 and 1.3 A, respectively. The waveforms also became less pulsatile with the onset of ventricular collapse.The reduction in LV pressure also manifested in a reduced peak pump flow rate and pump current Figure . NeverthIn this study, we developed a cardiac multiphysics modeling framework for biventricular electrical, mechanical and fluid dynamics physics in a healthy heart. The modeling framework was then applied to simulate a failing heart with LVAD support. This framework can be applied in future pre-clinical investigations of device performance under various pathological conditions.The simulated LV activation pattern is nearly radial and the RV activation pattern is nearly tangential ending at the basal region, whilst the interventricular septum is predominantly activated in a left to right sequence. The first epicardial breakthrough occurred in the RV due to its thinner wall structure. As such, the model's electrical activation reproduced normal human electrical activation sequence  than typical values reported for healthy humans (< 30 mmHg), even though the outlet boundary area is within the healthy human range  device by the US FDA.kTa, the standard model was modified to simulate typical characteristics of LVAD recipients. The diseased heart model's LV EF (16.5%), LV internal diameter (7.4 cm) and LV EDV (299 ml) are within the ranges observed in severe dilated heart failure patients , which linearly relates the maximum pressure that can be generated with the LV volume. During LV collapse following mitral occlusion, the LV is held at an extremely low volume, so, according to the ESPVR it is expected that the generated LV pressure is low   should be implanted in a patient to ensure optimal management.the impact of cannula shape and positioning on LVAD performance. Non-optimal cannula placement can reduce pump performance and has been suggested to induce thrombosis  rule, which relaxes the clinical efficacy requirement due to the lack of test subjects .In terms of electrophysiology, the model presently does not include several mechanofeedback components such as the stretch-activated currents and stretch effect on the cellular membrane's caveolae. Stretch-activated channels have been proposed to be capable of triggering electrical activation via mechanical stretch alone Sachs, . HoweverThe use of phenomenological AP and active stress formulations provided considerable simplification, since it reduced the number of variables required to be solved relative to biophysically-based models. This also enabled the use of larger time step steps, for more efficient computation. Mechanical dependencies of the contraction strength may be necessary for specific future studies to simulate the Frank-Starling mechanism, which was not considered in the present model. The mechanism enables the heart to pump at a greater ejection fraction should it be filled with a larger amount of blood (Guyton and Hall, Passive myocardial mechanical properties were simply assumed to be a transverse isotropic hyperelastic, whilst in reality, the myocardium exhibits orthotrophic properties (Dokos et al., Viscoelasticity is a common property of myocardial tissue, contributed by the collagenous structure and intercellular fluids, helping to damp out vibration, due to the blood flow (Cansiz et al., Active atrial contribution to the filling phase was not considered in the present study. The atrial contraction is expected to produce an additional peak in the filling flow rate waveform, which forms the “A” wave commonly observed in Doppler imaging of the flow waveforms (Sohn et al., Valve structure will influence the intra-ventricular blood flow profile. Furthermore, the angled position of the valves may also ease flow resistance during filling and ejection. Among LVAD recipients, tricuspid regurgitation has also been noted, in particular in those who developed RV failure, since the RV may not be able to cope with the restored systemic venous return (Hayek et al., k2, resulting in increased conduction velocity, will require finer meshing to achieve a converged conduction velocity and to prevent numerical convergence failure at large element size. Therefore, future mesh settings need to be tailored to the requirements of the study.The model mesh we have used was selected as a compromise between global convergence and computational cost. Fluid dynamics, especially during the diastolic phase, can involve more complex flow patterns, due to vortex formation and the presence of low velocity regions; these may require a finer mesh to be accurately modeled. With the present mesh setting, a converged solution was obtained for blood flow systolic characteristics, but the convergence slightly worsens during the isovolumic phases and diastolic filling. Nevertheless, qualitative and global measures such as vortex shape and flow waveforms can be sufficiently obtained with present mesh settings. In addition, changes to action potential parameters such as σ and The cardiac multiphysics framework developed was applied in a biventricular structure with an idealized Purkinje fiber network utilizing a modified form of the standard cable equation, taking into account the change in Purkinje fiber radius. The Windkessel circulation was expanded by including a closed-loop circulation which linked both ventricular chambers. The model managed to simulate healthy LV and RV electrical activation sequences, mechanical behavior as well as global hemodynamics. This highlights the model's capability for future biventricular modeling work. The model was also tested to simulate LVAD support and impact of the pump on ventricular function. The model was shown to be capable of predicting changes in pump variables following changes in the state of the heart, possibly aiding pump controller design in future. The fluid-electromechanics model enables study of the link between electrical activation, myocardial mechanics and blood hemodynamics characteristics, allowing holistic simulations of new therapeutic approaches in their ideation stage, aiding to speed up technology transfer to pre-clinical and clinical trial stages.AAB carried out the computational modeling. AAB, AA, and SD developed the framework for the standard model. AAB and AA drafted the manuscript. AAB, AA, MS, NL, and SD developed the framework for LVAD simulations, conceived of and coordinated the study, reviewed the manuscript and gave final approval for publication.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.


Coronary angiography excluded stenotic lesions. A wrap-around left anterior descending (LAD) with myocardial bridging in the mid-segment was observed Video 2: IVUS demonstrating systolic compression of the mid-LAD with a minimum lumen area oscillating from 3.06 mmVideo 3: Left ventricular (LV) angiography revealing typical mid-apical ballooning with hypercontractility of the basal segmentsVideo 4: Follow-up Echocardiography illustrating restoration of left ventricular function


Background: The possibility of non-invasively assessing load-independent parameters characterizing cardiac function is of high clinical value. Typically, these parameters are assessed during resting conditions. However, for diagnostic purposes, the parameter behavior across a physiologically relevant range of heart rate and loads is more relevant than the isolated measurements performed at rest. This study sought to evaluate changes in non-invasive estimations of load-independent parameters of left-ventricular contraction and relaxation patterns at rest and during dobutamine stress.Methods: We applied a previously developed approach that combines non-invasive measurements with a physiologically-based, reduced-order model of the cardiovascular system to provide subject-specific estimates of parameters characterizing left ventricular function. In this model, the contractile state of the heart at each time point along the cardiac cycle is modeled using a time-varying elastance curve. Non-invasive data, including four-dimensional magnetic resonance imaging (4D Flow MRI) measurements, were acquired in nine subjects without a known heart disease at rest and during dobutamine stress. For each of the study subjects, we constructed two personalized models corresponding to the resting and the stress state.Results: Applying the modeling framework, we identified significant increases in the left ventricular contraction rate constant [from 1.5 ± 0.3 to 2 ± 0.5 (p = 0.038)] and relaxation constant [from 37.2 ± 6.9 to 46.1 ± 12 (p = 0.028)]. In addition, we found a significant decrease in the elastance diastolic time constant from 0.4 ± 0.04 s to 0.3 ± 0.03 s (p = 0.008).Conclusions: The integrated image-modeling approach allows the assessment of cardiovascular function given as model-based parameters. The agreement between the estimated parameter values and previously reported effects of dobutamine demonstrates the potential of the approach to assess advanced metrics of pathophysiology that are otherwise difficult to obtain non-invasively in clinical practice. Quantification of left ventricular (LV) contractility is critical to the detection and management of many cardiovascular diseases, including ischemic heart disease, valvular disorders, and heart failure. The gold standard for deriving load-independent parameters characterizing LV contractility and diastolic relaxation relies on specialized and rarely performed direct invasive measurements of LV pressures  data .The subjects underwent an MRI protocol consisting of two acquisitions, one acquisition at rest and a second acquisition during the infusion of dobutamine. The dobutamine infusion protocol was initiated after the acquisition of 4D Flow MRI data performed at rest. Dobutamine was administered intravenously starting at a dose between 5 and 10 μg/kg/min, and the dose was increased or decreased in intervals of ~2 min depending on the observed effect on the subject's heart rate and arterial pressure. The target heart rate was 50% higher than the subject's rest heart rate, in order to achieve heart rates in the range of those from previous studies on stress testing with cardiac magnetic resonance (CMR) . During the examination, we acquired morphological cine bSSFP images and 4D Flow MRI data.2 and the LAx images 0.83 × 0.83 mm2.Morphological bSSFP images included three- and four-chamber long-axis (LAx) images and a stack of SAx images. During the scan performed after the infusion of dobutamine, only the three-chamber image and the SAx stack were acquired. The images were acquired during end-expiratory breath holds using the following settings: echo time 1.4 ms, repetition time 2.8 ms, flip angle 45° and slice thickness 8 mm. The bSSFP images were reconstructed into 30 time frames. The SAx images were reconstructed with a pixel size of 0.9 × 0.9 mm3 and the temporal resolution 41.6 ms. This resulted in a scan time of ~7–8 and 10–15 min excluding and including navigator efficiency, respectively, for a heart rate of 60 bpm. Following acquisition, the data were retrospectively reconstructed into 40 time frames and corrected for concomitant gradient fields on the scanner. Subsequently, the data were transferred to an offline station for post-processing and analysis. Post-processing was performed using in-house developed software written in Matlab . This included correction for background errors using a weighted second-order polynomial fit to static tissue in the thorax  is the chamber pressure and V(t) is the volume. Vo is the unstressed volume, defined as the volume-axis intercept of the end-diastolic relationship with the volume axis.where The shape of the elastance curve was approximated by a Double-Hill function . For the left ventricle, Emax corresponds to the end-systolic elastance and Emin characterizes the passive diastolic elastance of the chamber. A representative cardiac elastance curve is illustrated in Figure S, αD, RC, and RR determine the shape of the curve within each cardiac period. The rate of chamber contraction, reflected as the slope of the ascending part of the curve, is represented by the parameter RC. Similarly, RR corresponds to the rate of chamber relaxation . The systolic and diastolic time constants  are defined as a function of T and the dimensionless shape parameters αS and αD (τS = αS · T and τD = αD · T). αS and αD determine the length of systole and diastole relative to each cardiac cycle each cardiac cycle, thereby controlling the time of end-systole (tmax). An increase in αD and/or a decrease in αs yield an increase in tmax , and vice versa. In the following, we refer to αS and αD as systolic and diastolic time constant shape parameters, respectively. The scaling factor α ensures a maximal elastance of Emax.where The pressure gradient across the aortic valve was modeled using the formulation of the net pressure gradient introduced by Garcia et al. , which tR and Rv), a compliance (C), and an inductance (L). In each segment, the resistance R and the inductance L represent frictional losses and mass flow inertia, respectively, while the combination of compliance C and resistance Rv represent the viscoelastic properties of the vessel wall . Details on the simulation settings for solving the model equations are described in Casas et al.  was derived as the average cardiac cycle duration throughout the 4D Flow MRI acquisition. Estimation of the left ventricular elastance requires computing the left ventricular end-systolic volume (LVESV). This volume was calculated by manual delineation of the endocardial border, including the papillary muscles, in the short-axis stack at the time of end systole. In estimating the EOA of the aortic valve, the stroke volume (SV) was computed by integrating the volumetric flow curve at the aortic valve over the entire cardiac cycle. The velocity-time integral of the transvalvular flow (VTI) was calculated using the instantaneous velocity at the center of the valve. The cross-sectional area of the aorta was measured at the time of peak systole, distal to the coronary artery ostia.Transmitral (MV) and transaortic (AV) valvular flows were extracted using a semi-automatic retrospective valve tracking approach  was computed as the mean arterial blood pressure (MAP) divided by the cardiac output (CO), as described elsewhere , and the density of blood (ρ) are also defined as parameters. Twenty three parameters were adjusted to be subject-specific, while the remaining 17 were defined based on literature values from previous studies  .Statistical analysis was performed using SPSS Statistics v.23.0 . We performed non-parametric Wilcoxon signed-ranked tests to assess the differences between variables at rest and under dobutamine stress. A All subjects completed the study according to the protocol and no complications were seen in relation to the infusion of dobutamine. Across all subjects, the peak dose infused during the scan time was 15 ± 6 μg/kg/min (range 10–25 μg/kg/min). The maximal peak dose of 25 μg/kg/min was reached in two subjects. Table p = 0.008) and 5.6 ± 1 to 9.2 ± 1.7 l/min (p = 0.008), respectively  and cardiac output (CO) induced by the infusion of dobutamine are shown in Figure Figure p = 0.011). Dobutamine administration also increased the late (A) peak flow from 164.6 ± 45.5 to 266.4 ± 94.5 mL/s (p = 0.008). Furthermore, we observed that the E and A waves of the mitral flow waveform progressively merged with increasing heart rates, as illustrated in Supplementary Figure p = 0.14). Supplementary Figure p = 0.008), as did the peak flow at the ascending aorta .Figure Emax, LV) from 1.8 ± 0.5 to 4.7 ± 3.6 mmHg/mL (p = 0.013), as seen in Figure Emin, LV) did not change , extracted from the estimated time-varying elastance curve. Among all subjects, the decrease in this percentage was between 1 and 10%, with larger decreases in the diastolic contribution associated with higher heart rates. Shortening of diastole in relation to the cardiac cycle length was reflected in the elastance timing parameters by a significant increase in the diastolic time constant shape parameter, αD, LV, from 0.4 ± 0.05 to 0.5 ± 0.05 . Analysis of model-based parameters characterizing the vascular system revealed a non-significant increase in the compliance of the ascending aorta, Caa, from 0.11 ± 0.06 to 0.16 ± 0.1 mL/mmHg (p = 0.11).The increase in heart rate due to dobutamine infusion resulted in shortening of both the systolic and diastolic phases, with a moderately greater reduction in the duration of diastole in relation to systole. The percentage of diastolic duration relative to the cardiac cycle length was calculated based on the time of end-systole (We evaluated a previously developed imaging-modeling approach for quantification of load-independent parameters of LV contractility and relaxation. An advantage of this approach is that it allows the estimation of physiological parameters, which typically require invasive measurements, in a non-invasive fashion. The evaluation was performed on a cohort of nine subjects without a known heart disease, both at rest, and under stress conditions induced by the infusion of dobutamine. Our results indicate that the approach is suited for both description of the experimental data and determination of parameters characterizing LV systolic and diastolic function across a physiologically relevant range of heart rates and loading conditions. This demonstrates that our approach has the ability to describe operating conditions for which it was not originally developed and serves to validate the realism of the model. This is therefore a key stepping stone toward eventual clinical implementation.Several studies have investigated the effects of clinically induced stress on cardiovascular function based on non-invasive indices derived from Doppler ultrasound or 2D PC-MRI measurements, such as left ventricular volumes and mitral inflow patterns. In line with our results, these studies have reported a decrease in left ventricular ESV (Steding-Ehrenborg et al., The model-based parameters revealed subject-specific increases in contractility as indicated by the increase in rate of left ventricular contraction under stress conditions, as well as an enhancement of diastolic function by faster early relaxation. The latter was demonstrated by an increase in the rate of left ventricular relaxation. These results are consistent with previous studies reporting an increase in these parameters based on invasive measurements (Karliner et al., In the present study, we estimated the maximal elastance of the left ventricle using a single-beat approach, assuming a volume intercept of 10 mmHg as defined in previous studies (Mynard et al., Our results showed no significant differences between passive diastolic compliance at rest and during stress. Schmitt et al.  reportedThe current study has several limitations besides the ones mentioned in our previous publication (Casas et al., The study only included a limited number of subjects. This limits the statistical power of the study and the extrapolation of our results. This should therefore be taken into consideration when interpreting the statistical significance of the results. Nevertheless, the study group displays a broad range of physiologically relevant loads and heart rates, and the changes in parameters were significant despite the low total study population. Additional studies including a larger number of subjects should be conducted to confirm the clinical relevance of the statistical significances reported in the current work.In addition, we only evaluated the changes in parameters with stress in subjects without known cardiovascular disease. The analysis of these parameters in subjects with a spectrum of diseases, such as ischemic heart disease, valvular disorders, and heart failure is subject to further studies. Those studies should focus on assessing the added value of these parameters over conventional image-derived parameters in diagnosing and optimizing the treatment of the disease.The integrated imaging-modeling approach used in this study provided subject-specific parameter values that behaved appropriately, according to previously reported effects of dobutamine on cardiovascular function, across a physiologically relevant range of loads and heart rates. With further validation, the approach could be used to assess these parameters, which are difficult to obtain non-invasively in clinical practice, thereby complementing conventional non-invasive indices used in the assessment of ventricular function. Future validation studies should preferably involve comparison to invasive measurements and application on larger populations, including both healthy subjects, and patients representing a spectrum of cardiovascular diseases. The personalized assessment may have the potential to generate advanced metrics of cardiovascular physiology and pathophysiology that could extend beyond conventional techniques for both diagnosis and optimization of a personalized medical regimen.The datasets for this manuscript are not publicly available due to restrictions in the ethical approval. Requests to access the datasets should be directed to the corresponding author.BC, AB, MK, C-JC, and TE participated in the conception and design of the study. C-JC performed the recruitment of patients and carried out the data acquisition. BC and FV were involved in the analysis of the data. BC, GC, and TE participated in the implementation of the method. BC performed the computational analysis and drafted the manuscript. BC, GC, AB, MK, C-JC, and TE interpreted the results. All authors edited and revised the manuscript. All authors read and approved the final version of 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.


Under conscious conditions, there was a significant decrease in the gain  as well as the lower plateau of the baroreflex in HF compared to control animals. HF was also associated with significantly increased respiratory rate  and incidence of apneas , compared to control sheep. The microembolization model of heart failure is associated with an increase in left ventricular end-diastolic pressure, impaired cardiac function, and altered baroreflex control of the heart. These findings suggest this chronic model of HF is appropriate to use for investigating interventions aimed at improving neural control in HF.Testing new therapies in heart failure (HF) requires a chronic stable model of HF in large animals. Microembolization of the coronary arteries has been used to model HF previously; however, neural control has not been previously explored in this model. Thus the aim of this study was to further characterize neural control in this model of HF. HF was induced by infusion of microspheres  into the proximal left coronary artery or left descending coronary arteries, with three sequential embolizations over 3 weeks. Twelve to 14 weeks after the final embolization, and when ejection fraction had decreased below 45%, animals were instrumented to record blood pressure and heart rate. Baroreflex control of heart rate was investigated in conscious animals. Additionally, pressure-volume loops were constructed under anesthesia. Embolization-induced HF was associated with a decrease in mean arterial pressure (67 ± 2 vs. 85 ± 4 mmHg,  Heart failure with reduced ejection fraction (HFrEF) is characterized by progressive dysfunction of left ventricular muscle, myocyte remodeling, and activation of autonomic and hormonal systems . DespiteNumerous studies have been conducted on small animal models of HF and these have been instrumental in understanding the pathogenesis of HF. However, there are significant differences between small animal models of HF and human HF beyond scale. These include differences in baseline heart rate, oxygen consumption, and excitation-contraction coupling in cardiac tissue . TherefoThe primary methods to induce HFrEF in large animals include high rate pacing, cardiotoxin infusion, and coronary ligation. The pacing-induced HF model has been used in sheep , pigs T, and dogIntracoronary or intravenous infusion of the cardiotoxin doxorubicin in dogs  and sheeFinally, coronary ligation has been used to induce HF in large animals . InfarctRecently, the microembolization model has been used in sheep to develop a chronic model of HF over a period of 3 months. The embolizations have been conducted using injection of either microspheres  or gelfoad libitum.Experiments were conducted on conscious, adult female Romney ewes weighing 50–80 kg, housed in individual crates, and acclimatized to laboratory conditions  and human contact before any experiments. All experiments and surgical procedures were approved by the Animal Ethics Committee of the University of Auckland. The sheep were fed 2 kg/day (Country harvest pellets) and had access to water 2 mixture. Sheep were given antibiotic injections  at the start of the surgery. Additionally, to provide analgesia, sheep were given Ketofen 10%  at the start of surgery.Nine sheep  underwent three weekly sequential embolizations of either the proximal left coronary artery or left descending coronary arteries. Six control sheep  without microembolizations were also used in the study. Anesthesia was induced with 2% Diprivan (Propofol)  and maintained with a 2% isoflurane-air-OThe methods employed have been described before . Once anEchocardiogram recordings were obtained and analyzed before embolizations and 3 months after the first embolization procedure. This was also done for the group of control sheep. The echocardiogram, using a Hewlett Packard Sonos 1,000, was performed while the sheep were conscious. In the long-axis M-wave echocardiography, diastole, systole, fractional shortening, and ejection fraction parameters were obtained and calculated for the left ventricle.After 3 months, once sheep were deemed to have sufficient left ventricle dysfunction (ejection fraction <45%), the animals were instrumented to measure mean arterial pressure (MAP), heart rate (HR), and diaphragmatic electromyography (dEMG) as an index of respiration. The instrumentation procedure was also conducted in a group of control sheep. The procedure for electrodes placement for dEMG has been described previously . Two strAll recordings were done at least 3 days after instrumentation surgery. Blood pressure was obtained from a pressure probe unit . Heart rate (HR) was calculated from blood pressure channel. In conscious, standing sheep, HR and MAP were obtained from a 2-h recording and averages were obtained for each animal. To determine the adrenergic effects on hemodynamic parameters and heart rate, β-adrenergic receptor blockade  was infused (30 mg bolus followed by 0.5 mg/kg/h infusion for 90 min).In two groups of six conscious sheep, after a 5-min baseline recording of mean arterial pressure and heart rate, baroreflex curves were generated by measuring the responses of heart rate to increasing doses of phenylephrine hydrochloride  and sodium nitroprusside . For analysis, the baseline blood pressures were sorted from the lowest to the highest pressures and put into bins of 3 mmHg change each. The mean systolic blood pressure of each bin was plotted against the mean HR.Venous blood samples (10 ml) were collected into an EDTA  tube. Plasma was rapidly separated with a centrifuge at 4°C at 3,000 rpm, within 5 min of blood collection, and snap-frozen at −80°C. The assays for brain natriuretic peptide (BNP) , epinephTo assess diaphragmatic EMG parameters, resting breathing rate and apnea periods were measured in control and HF animals. Breathing rate average was obtained from dEMG activity in a 12-h baseline period in each animal. Apnea was defined as cessation of diaphragmatic activity. To be considered significant, apnea events had to persist for a minimum of 4 s. To quantify apnea incidence, we calculated an apnea index to indicate the number of apnea periods (>4 s) occurring in a 12-h period.via a guide cannula, along the longitudinal axis. The catheter was connected to a Millar  and AD Instruments pressure-volume processing unit and signals were acquired using LabChart, a data acquisition and data analysis software. Volume correction was done through an estimation, using a hypertonic saline solution  infusion. A cardiac output transonic flow probe was also inserted around the aorta. Briefly, a thoracotomy was performed and fourth rib removed to access the heart. A flow probe was implanted on the aorta  and connected to LabChart to measure cardiac output while performing pressure-volume loops.The acute pressure-volume loop experiments were conducted at the end of the protocol under anesthesia. To determine the left ventricle pressure-volume relationship, a conductance catheter was placed into the left ventricle through the left carotid artery. This method has been described previously in detail by Pressure-volume loops were analyzed offline using LabChart (AD Instruments). After volume correction, 10 cardiac cycles at baseline were analyzed. Also, cardiac output recordings from the animal were used for alpha calibrations. Stroke volume, ejection fraction, left ventricular end-diastolic and end-systolic pressure and volume, maximum rate of pressure generation (dP/dt max), peak rate of pressure decline (dP/dt min), and maximum conduction velocity (dV/dt max) were measured.At the end of these experiments, the sheep were euthanized with an overdose of sodium pentobarbitone  . Once respiration and cardiac function had ceased, cardiac tissue was collected for histological analysis.At the end of the terminal experiment, the hearts were collected and a portion of the left ventricle wall  was fixed in neutral buffered formalin . Tissue blocks were then transferred to 30% sucrose solution. Heart blocks were then cut at 40-μm sections using a cryostat. To measure collagen deposition in the left ventricle wall of the sheep hearts, 12 sections (six control and six HF sheep hearts) were stained with Masson’s Trichrome stain (MTS) solution . Cardiac muscle fibers stain red and collagen stains blue. Sections were subjected to two washes and then three changes of 100% ethanol and finally one wash of xylene solution before cover slipping.t-tests (two groups – control vs. heart failure). The effects of the individual microembolization procedures on the baseline levels of ECG (before and after each microembolization) and ejection fraction and fractional shortening (before and 12–14 weeks after microembolizations) were determined using paired Student’s t-tests . A one-way ANOVA test was used for changes in resting heart rate plotted in 3 mmHg change bins. Data were analyzed using the statistical package SigmaStat . Data were considered significant if p < 0.05.All data are expressed as mean ± SEM, except where indicated. The effects of the microembolizations on the baseline levels of MAP, HR, ejection fraction, fractional shortening, respiratory rate, apnea incidence, heart and body weight, collagen fibers deposition, pressure-volume relationship parameters, and the arterial baroreceptor relationships of HR and sysBP were determined using unpaired Student’s p < 0.05; p < 0.05). These differences in embolization procedures 2 and 3 compared to embolization 1 were also observed in the T-wave amplitude.Embolization of the coronary artery at the first time point resulted in a significant change in the lower S-T segment height and also the T-wave amplitude (p < 0.001) and fractional shortening (p < 0.0001) in HF sheep, after 12–14 weeks (p < 0.01), and an increase in heart rate  compared to the control animals. Plasma levels of norepinephrine and brain natriuretic peptide were also significantly higher (p < 0.05) (p < 0.01)  and there was no significant change in MAP between the groups after propranolol infusion.The resting levels of hemodynamic variables in the control and HF sheep are shown in 14 weeks . HF was  < 0.05) . The res < 0.01) . The HF p < 0.05), as well as the upper  and lower  plateaus of the HR baroreflex curve ; however, the difference between the groups remained intact. Interestingly, there was no change in baseline cardiac output between the normal and the HF sheep. The LV end-diastolic pressures and volumes in the HF sheep were significantly higher, similar to end-systolic volume . The wet heart weight of the sheep in HF was also significantly greater than that of the control sheep  there is a depression in the gain of the baroreflex in this model of HF; (2) diaphragmatic electromyography showed increased respiratory rate in the HF animals compared to the control animals, suggesting increased inspiratory effort; (3) in addition to an increase in respiratory rate, there was an increased incidence of apneas in this animal model. Taken together, our findings suggest that this model of HFrEF replicates the neural and respiratory instabilities seen in human HF.One of the main clinical signs of a myocardial infarct is changes in the ECG, namely changes in the S-T segment and the T-wave. For the S-T segment, in most cases, patients that have had a suspected myocardial infarct present with an elevation in this segment . In contSeveral other studies, in patients, have shown collagen deposition in the myocardium, suggestive of ventricular remodeling . The groOne of the ways to assess the degree of HF is the construction of pressure-volume loops. One of the advantages of constructing pressure-volume loops is the measurement of parameters independent of pre- and after load . A hallmIn the present study, we show significant increases in plasma levels of brain natriuretic peptide (BNP) and norepinephrine . These iWe report a significant decrease in the MAP of the animals with HF. One mechanism that serves to maintain BP is the arterial baroreflex and previous studies have reported a dampening of the sensitivity of the arterial baroreflex control of heart rate in HF  as well In addition to changes in hemodynamic parameters, we also observed an increase in the breathing rate and an increase in the incidence of apneas. Chronic hyperventilation has been reported in HF patients  and pulmWe report successful induction of heart failure in sheep using repeated percutaneous injection of microspheres into the coronary arteries. This approach targets the coronary circulation and thereby circumvents potential systemic effects if a toxin is administered intravenously. Since the HF is induced by a durable ischemic insult, interventions with novel pacing paradigms can be undertaken in pacing naïve hearts, avoiding the potential confounder of pacing-induced changes. Practically, repeated access to the arterial circulation is an important consideration. Previous studies have constructed carotid artery loops to allow repeated access. We percutaneously accessed the femoral arteries and following embolization, digitally compressed the puncture site for hemostasis without the use of a closure device, returning the animal to the crate/pen without any indwelling cannulae retained for future access. This has a number of advantages over other methods of surgical approach to perform the microembolization. Cardiac catheterization using these techniques requires access to fluoroscopy to guide cannulation and confirm successful and stable engagement in the coronary vessels.There are a few limitations to this approach as well. One important limitation of this approach is the diffuse pattern of blockage and fibrosis that is observed. As such, it is difficult to quantify the region of infarcted tissue adequately as would be possible with coronary ligation. We utilized ejection fraction and fractional shortening from the echocardiographic data to quantify the decline in heart function. In addition, three out of the nine sheep did not survive the first embolization as they developed pulmonary congestion that could be resolved using diuretics. In addition to these three animals, two of the animals studied also needed administration of diuretics after the embolization procedure, and in these two cases, the congestion was resolved and the animals continued in the protocol. While the presence of congestion makes this model clinically relevant, this does necessitate increased animal numbers to complete a cohort of animals for a study.In conclusion, our study indicates that chronic heart failure can be successfully induced in sheep using repeated injection of microspheres into the coronary arteries. After 2 months, the animals develop clinical signs of heart failure. There is marked increase in heart weight with histological evidence of ventricular fibrosis. There is a decrease in mean arterial pressure and an increase in incidence of apneas. Furthermore, pressure-volume loops show altered left ventricle dynamics in the heart failure sheep and the baroreflex challenge under conscious conditions showed a significant decrease in the gain. We conclude that this is a good model of HF to test changes in either neural control or respiratory function after interventions.The data gathered in this study are available upon request to the corresponding author.The animal study was reviewed and approved by The University of Auckland Animal Ethics Committee.YA, NL, MP, IL, and RR collected the ECG data for this study. All other data were collected by RR, MP, and YA. RR supervised the study. YA completed the data analysis and wrote initial manuscript. All authors contributed in revising the manuscript toward the final version.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.


Optimal right ventricular lead placement remains controversial. Large studies investigating the safety and long-term prognosis of apical and septal right ventricular lead placement have been lacking.CArdiac REsynchronization in Heart Failure trial. Secondary endpoints were all-cause mortality, new heart failure and their composite.Consecutive patients undergoing pacemaker insertion for high-degree atrioventricular block at Edinburgh Heart Centre were investigated. Periprocedural 30-day complications were defined (infection/bleeding/pneumothorax/tamponade/lead displacement). Long-term clinical outcomes were obtained from the General Register of Scotland and electronic medical records. The primary endpoint was a composite of all-cause mortality, new heart failure, hospitalisation for a major cardiovascular event, as per the 820 patients were included, 204 (25%) paced from the septum and 616 (75%) from the apex. All baseline variables were similar with the exception of age . Procedure duration , complication rates (18 (8.8) vs 46 (7.5)%, p=0.5) and postimplant QRS duration (152 (23) vs 154 (27) ms, p=0.4) were similar. After 1041 days (IQR 564), 278 patients met the primary endpoint, with no difference between the septal and apical groups in unadjusted (HR 0.86 (95% CIs 0.64 to 1.15)) or multivariable analysis correcting for age, gender and comorbidity (HR 0.97 (95% CI 0.72 to 1.30)). Similarly, no differences were observed in the secondary endpoints.This large real-world cohort of patients undergoing right ventricular lead placement in the septum or apex demonstrated no difference in procedural complications nor long-term clinical outcomes. Both pacing strategies appear reasonable in routine practice. There remains a lack of consensus on optimal right ventricular lead placement.Periprocedural complication rates and long-term clinical outcomes for septal and apical right ventricular lead placement are similar.Fluoroscopy-guided right ventricular septal lead placement does not consistently result in a narrow paced QRS duration.Patients in whom narrow paced QRS duration is achieved have a favourable long-term clinical outcome.Reduction of QRS duration is crucial and should be prioritised when placing right ventricular leads.Pacing strategies for routine clinical practice that result in more consistent reduction of QRS should be the focus of future research in this field.Cardiac pacing devices are central to modern cardiology with over 500 000 devices implanted in the annually across Europe.Conventionally, right ventricular leads are placed at the apex, but increasing evidence suggests this strategy may have deleterious effects on cardiac function by producing an iatrogenic left bundle branch block (LBBB) pattern on the ECG and dyssynchronous ventricular contraction.Alternative pacing strategies achieving more physiological depolarisation might improve ventricular synchrony and protect against these detrimental effects. These include minimal ventricular pacing algorithms, upgrade to cardiac resynchronisation therapy and His-bundle pacing.In the present study, we aimed to investigate the procedural safety and long-term clinical outcomes of a large real-world cohort of patients with higher degree atrioventricular block non-selectively assigned to pacing operators with preference for either septal or apical right ventricular pacing strategies.Consecutive patients undergoing pacemaker device implantation from 16 April 2010 to 29 September 2016 at the Edinburgh Heart Centre were included in the study. Over the study period, there were five operators with two favouring septal right ventricular lead placement and three favouring apical lead placement. All septal lead placements were achieved with active fixation leads using stylets fashioned by the operators to facilitate septal positioning. No preshaped stylets or steerable sheaths were used. The final septal positions were confirmed using fluoroscopy (posteroarterior (PA) and left anterior oblique (LAO) 40 projections) and assessment of the current of injury. In comparison, most apical lead placements were achieved with active fixation leads  using PA and right anterior oblique fluoroscopy projections. Patients listed for permanent pacemakers were placed in a central pool and then non-selectively allocated to 1 of 6 weekly lists at our institution. Inclusion criteria were defined to select patients undergoing new right ventricular lead implantation (either new device implants or ventricular lead repositioning) for high-degree atrioventricular block (second/Mobitz type 2 or third-degree atrioventricular block). Exclusion criteria were pacemaker implantation for isolated sick sinus syndrome or other conditions associated with low ventricular pacing burdens and patients undergoing cardiac defibrillator or resynchronisation therapy system implantation. Patients in sinus rhythm undergoing single-lead device implantation were excluded in an attempt to exclude frail and unstable patients with an inherently poor prognosis who may be more likely to have an apical lead placed . Clinica2), liver failure, metastatic cancer, stroke or transient ischaemic attack, diabetes mellitus, hypertension (persistent blood pressure 140/90 mm Hg or higher), ischaemic heart disease (symptomatic of angina pectoris or imaging diagnosis), previous coronary arterial bypass grafting, valvular heart disease (murmur or echocardiographic diagnosis), previous valvular surgery and interventricular conduction delays  blinded to clinical outcomes. Data on preimplantation medication and comorbid conditions were recorded including atrial arrhythmias, congestive heart failure (symptomatic or echocardiographic ejection fraction <50%), chronic kidney disease . Our composite primary endpoint was prespecified and based on the Statistical analysis was carried out in R  and GraphPad Prism . Continuous variables were assessed using parametric and non-parametric tests as appropriate, while categorical variables were compared using the Fisher’s exact test. Cox regression modelling was performed to define the HRs for all primary and secondary endpoints after adjusting for age, sex and comorbidities . A post hoc Cox regression analysis with selected relevant clinical endpoints was also carried out. Finally, we performed a subgroup analysis to investigate the breadth of the paced QRS duration in patients with a septal lead (narrow complex paced QRS <130 ms vs broad complex paced QRS ≥130 ms) and the effect of paced QRS duration on the primary endpoint.In total, 820 eligible patients were identified with 616 (75%) in the apical and 204 (25%) in the septal right ventricular pacing groups. These two groups were well balanced for all baseline clinical variables including gender, comorbidities, medication and routine baseline laboratory tests. The sole exception was age with patients having a septal lead being on average 3.7 years younger than those with apical leads .There was no difference in procedure times (defined as time from skin preparation to skin closure) between apical and septal groups  or fluoroscopic screening times . Pneumothorax, infection, bleeding and haematoma were the most frequent complications with no observed differences between the two groups in their incidence  or in thAfter a median of 1041 (IQR 564) days follow-up, 278 patients met the primary endpoint . There wIn a subgroup (n=343) where paced postimplant electrocardiograms were available, there were no differences in paced QRS duration between the septal and apical groups (mean 152±23 (n=87) ms vs 154±27 (n=256) ms, respectively, p=0.4). Across the cohort as a whole, 16% of patients  demonstrated a narrow paced QRS (<130 ms) (excluding fusion and pseudofusion beats), while 84% (n=289) had a broad paced QRS (≥130 ms). Interestingly, patients in whom a narrow paced QRS (<130 ms) was achieved, irrespective of pacing site, demonstrated improved clinical outcomes compared with patients with a broad QRS . Kaplan-Meier curve analysis confirmed this finding with improved time-to-event outcomes in the narrow complex group p=0.01; .In this large-scale retrospective analysis, where patients underwent right ventricular septal or apical pacing, we have demonstrated that right ventricular septal pacing is a safe alternative to the conventional apical approach with no increase in periprocedural complication rates or procedural times. However, septal pacing did not result in improved long-term clinical outcomes, perhaps because it did not consistently result in a narrow complex paced QRS. Both right ventricular septal and apical pacing are therefore reasonable strategies with further work required to develop simple approaches that more consistently reduce the paced duration of the QRS.While not a prospective randomised trial, the present study provides analysis of real-world consecutive patient data at a tertiary referral centre, where patients are non-selectively assigned from a central pool to pacing lists performed by operators with a preference for either apical of septal lead placement. As a consequence, the right ventricular septal and apical pacing cohorts were well balanced across every baseline clinical characteristics that were collected with the sole exception of age. Interestingly, the septal cohort was on average 3.7 years younger, perhaps reflecting the tendency to be less aggressive in striving for a septal pacing site in elderly patients. We therefore provide a robust and large real-world dataset with which to investigate the procedural safety and long-term clinical outcomes of septal versus apical pacing.Septal lead placement poses technical challenges,Our second aim was to investigate whether septal pacing is associated with a favourable long-term prognosis. Despite a high event rate, no difference was observed between the right ventricular septal and apical pacing groups in the primary composite clinical endpoint, at median follow-up of 1041 days. Similarly, no difference was observed for any of the secondary clinical endpoints in multivariable analysis. What is the explanation for this result given the well-established literature outlining the adverse outcomes associated with LBBB and right ventricular apical pacing? Selection bias has potential to favour apical lead placement given that operators may be more likely to attempt septal lead implantation in patients with known left ventricular impairment. However, the high degree of similarity in comorbidity and medication between our patient groups suggests this was not the case. It is the analysis of the postimplantation ECGs that appears to provide the answer. In this large subgroup (n=343), no difference was observed in the paced QRS duration between the septal and apical pacing groups. Indeed, only 16% of the septal group achieved a narrow paced QRS (<130 ms), similar to the proportion in the apical cohort. Septal pacing therefore failed to reduce ventricular dyssynchrony compared with apical pacing, thereby explaining the similar outcomes. One potential explanation for this failure is that fluoroscopy is insufficient to guide accurate positioning of the pacing lead on to the septum. Recent studies have confirmed that only a minority of right ventricular leads are placed in the true septum using conventional fluoroscopic views and some have advocated the use of individualised LAO projection and echocardiography to confirm septal lead positioning.Interestingly, the patients in whom a narrow QRS paced rhythm (<130 ms) was achieved, irrespective of pacing site, did demonstrate a lower risk of future adverse events compared with those with a broad QRS≥130 ms . This is consistent with the hypothesis that pacemaker-related LBBB is associated with an adverse prognosis and supports strategies that aim to achieve narrow complex pacing. Successful reduction in QRS duration can be achieved with cardiac resynchronisation therapy, but this is time consuming. His bundle pacing is an alternative strategy with potential to produce the same QRS duration and mechanical function as atrial pacing.Echocardiography and pacing burden data were not available for analysis. This study was a retrospective analysis at a single centre and so is inherently limited by unforeseen counfounding factors and lack of randomisation. However, the system of non-selective patient allocation to pacing operators resulted in the balancing of all clinical variables examined with the exception of age. Similar balancing might therefore be expected among the unknown confounders.In this large real-world observational analysis, procedural safety and long-term clinical outcomes were similar between patients undergoing apical and septal right ventricular pacing. While septal pacing appears a safe and feasible strategy, further work is required to develop methods that can more consistently deliver narrow QRS right ventricular pacing.


Weaning post-cardiac surgery patients from mechanical ventilation (MV) poses a big challenge to these patients. Optimized left ventricular-arterial coupling (VAC) may be crucial for reducing the MV duration of these patients. However, there is no research exploring the relationship between VAC and the duration of MV. We performed this study to investigate the relationship between left ventricular-arterial coupling (VAC) and prolonged mechanical ventilation (MV) in severe post-cardiac surgery patients.This was a single-center retrospective study of 56 severe post-cardiac surgery patients from January 2015 to December 2017 at the Department of Critical Care Medicine of Peking Union Medical College Hospital. Patients were divided into two groups according to the duration of MV . Hemodynamics and tissue perfusion data were collected or calculated at admission (T0) and 48 h after admission (T1) to the ICU.P = 0.018), lactate , and Ea/Ees . Increased Ea/Ees  and lactate at T1  were independently associated with prolonged MV. There was a significant relationship between Ea/EesT1 and the duration of MV . The area under the receiver operating characteristic (AUC) of the left VAC for predicting prolonged MV was 0.801, and the cutoff value for Ea/Ees was 5.12, with 65.0% sensitivity and 90.0% specificity.In terms of hemodynamic and tissue perfusion data, there were no differences between the two groups at admission (T0). Compared with the non-prolonged MV group after 48 h in the ICU (T1), the prolonged MV group had significantly higher values for heart rate (108 ± 13 vs 97 ± 12, Left ventricular-arterial coupling was associated with prolonged mechanical ventilation in severe post-cardiac surgery patients. The assessment and optimization of left VAC might be helpful in reducing duration of MV in these patients. Prolonged mechanical ventilation (PMV) is required by 2.4–9.9% of post-cardiac surgery patients and is associated with increased medical resource use and mortality –4. CardiSeveral investigators have found that low cardiac output (CO) was associated with PMV in post-cardiac surgery patients , 8. HoweThe LV pumps blood to the arterial vessel, which is the output of energy . ImportaAs we all know, left ventricular-arterial coupling (VAC) was used to assess cardiac work efficiency . There aFor a given beat, the pressure–volume area  for unstable hemodynamics from January 2015 to December 2017 at the Department of Critical Care Medicine of Peking Union Medical College Hospital. These patients had a preoperative class III or IV New York Heart Association (NYHA) classification. The exclusion criteria include 1. Discharged within 48 h; 2. Discharged against medical advice or died in the first 6 days; 3. Acquired pneumonia or serious cerebrovascular accidents in the first 6 days of ICU admission.A standardized weaning protocol has been applied for all included patients. We divided patients into two groups according to the duration of MV . Medical records were reviewed to obtain information about gender, age, concomitant diseases, types of surgery, the 1st day SOFA, the duration of extracorporeal circulation (ECC), vasopressor and inotrope treatment, and the duration of MV and ICU stay. We also collected hemodynamics and tissue perfusion data at two time points  from the Critical Care Monitor System of the Department of Critical Care Medicine at Peking Union Medical College Hospital, which recorded real-time clinical data from bedside equipment. This system is maintained by Donghua software cooperation through DtHealth system. We also collected the PaO2/FiO2 (P/F) data on the first and second day in the ICU.Severe post-cardiac surgery patients defined as after conventional treatments, patients still have the following problems: a. lactate clearance less than 30% in the first 24 h;b. high dose of norepinephrine, more than 1μg/kg/min; c. undefined shock type, requiring more accurate monitoring techniques, like PiCCO.PMV group: prolonged MV group, defined as MV > 6 days; non-PMV group: non-prolonged MV group, defined as MV ≤6 days.End-systolic aortic pressure can be estimated as 0.9 times the peak brachial systolic pressure. In this study, we calculated Ea as 0.9*SBP divided by SV (stroke volume), which can be expressed as Ea = 0.9 SBP/SV.The simplified formula for calculating Ees was Ees = SBP/(GEDV/4-SV). Technically, Ees = ESP/(ESV-V0). ESP is LV end-systolic pressure and ESV is LV end-systolic volume. V0 represents a purely theoretical LV volume at zero intracavitary pressure . We consA descriptive analysis was performed. Continuous variables are expressed as the mean ± standard deviation, and categorical variables are expressed as absolute values and percentages. For continuous variables, the data were analyzed using Student’s t-test, the Mann-Whitney U test or the Kruskal-Wallis test depending on data distribution and the number of variables. Categorical variables were analyzed using the chi-square test or Fisher’s test.p value < 0.2). General demographics were also used in the model. Discrimination of values was performed using the receiver operating characteristic (ROC) analysis. All comparisons were two-tailed, and P < 0.05 was required to exclude the null hypothesis. The statistical analysis was performed using IBM SPSS Statistics, Version 20.0 .Variables were introduced into a multivariable binary logistic regression model if they were significantly associated with prolonged MV in the univariate analysis . No significant differences were found in the use of inotropes between the two groups during the three periods. The demographics and clinical characteristics of all the patients are presented in Table There were no significant differences in gender, age, types of surgery (CABG (coronary artery bypass grafting), pericardiectomy, valve surgery, ventricular septal defect repair, atrial neoplasm resection, aorta replacement) or concomitant diseases  between the two groups. The PMV group had a significantly higher SOFA score than the non-PMV group , SBPT0 (systolic blood pressure), CVPT0 , P/FT0 (PaO2/FiO2), lactateT0, CIT0 (cardiac index), SVIT0 (stroke volume index), SVRIT0 (systemic vascular resistance index), Ea,T0, Ees,T0 or Ea/Ees,T0 between the two groups (Table T1 (P = 0.018), lactateT1 (P < 0.001), and Ea/Ees,T1  and lower values for Ees,T1 (P = 0.022), SVIT1 (P = 0.049), SBPT1 (P = 0.014).There were no differences in HRP/FT0 PaO/FiO2, laT1and the duration of MV . There were no correlations between SVIT1 or CIT1 and the duration of MV.There was a significant correlation between Ea/Ees, MV Fig. ; the corT1 , Ea/Ees,T1  and lactateT1  were independent predictors of PMV in severe post-cardiac surgery patients , 0.801 (95% CI:0.664–0.938), 0.816 (95% CI:0.680–0.952), respectively , 5.12  and 1.60 , respectively, based on the maximum Youden index.ROC curves were drawn to compare the predictive values of three different variables for PMV Fig. . The AUCThe main findings of this study are as follows: (a) VAC is associated with the duration of MV, and a worse VAC indicates a longer duration of MV; (b) VAC is an independent risk factor for PMV. A cut-off value of VAC > 5.12 predicts PMV with 65% sensitivity and 90% specificity.Prolonged MV is associated with many complications , 18, sucWhen weaning begins, the cardiac preload and afterload of the LV increase dramatically, which requires the LV to increase its SW. If a patient’s heart has a greater SW reserve before weaning, which indicates better VAC, it will be much easier to achieve adequate SW and wean the patient earlier. In several previous studies, SW before weaning was considered a predictive factor for PMV . Actuallbefore-VACafter), which can reflect the cardiac compensatory function capability. Kensuke [,T1  was a risk factor for prolonged MV. Hence, the assessment and optimization of the left VAC might be helpful during ventilator weaning in severe post-cardiac surgery patients. Further studies are required to validate VAC-directed weaning from ventilation.Several studies , 21 haveVasodilators and inotropes are always used in severe post-cardiac surgery patients to assist in weaning from MV, which also optimizes VAC. Vasodilators, which decrease Ea by reducing the afterload, and inotropes, which augment Ees by strengthening contractility , are two0 out of the calculation formula and using GEDV/4 as LVEDV might be an arbitrary decision. However, we had several reasons for making this decision: (a) we considered the heart as a whole, because PMV is not only a problem of poor left ventricular function but also a thing of right ventricle [Several limitations should be acknowledged. First, leaving Ventricle . It mighentricle , and (c)entricle . Most imentricle –32. We aentricle , 34. FifCompared to CO, VAC is better in predicting PMV in severe post-cardiac surgery patients, and it’s also an independent risk factor for PMV, which means poorer VAC indicates higher possibility of PMV. The assessment and optimization of left VAC might be helpful when weaning severe post-cardiac surgery patients from MV.


Evidence available suggests that periprocedural bradycardia negates the benefit of primary percutaneous coronary intervention (PPCI) and worsens the prognosis of patients with acute ST-elevation myocardial infarction (STEMI).  To investigate the risk factors of periprocedural bradycardia during PPCI in patients with acute STEMI.  We enrolled 2,536 acute STEMI patients who had PPCI from November 2007 to June 2018 in Beijing Anzhen Hospital, Capital Medical University. We divided all patients into two groups according to periprocedural bradycardia  during PPCI: periprocedural bradycardia group (434 cases) and control group (2102 cases). We compared demographic, clinical, and angiographic characteristics of the two groups. We analyzed the risk factors of periprocedural bradycardia. P < 0.05). The area under the receiver operating characteristic curve was 0.8390.  The incident rate was 17.1% (434/2536). Logistic regression analysis showed that the differences between the two groups in no-reflow, the culprit vessel was LAD, using thrombus aspiration devices during operation, gender, completely block of culprit vessel, and intraoperative hypotension were statistically significant . No-reflow, the culprit vessel was not LAD, using thrombus aspiration devices during operation, gender, completely block of culprit vessel, and intraoperative hypotension may be independent risk factors for predicting periprocedural bradycardia during PPCI in patients with acute STEMI. We registered this study with WHO International Clinical Trials Registry Platform (ICTRP) (registration number:  All types of conduction disturbances can occur in the context of an acute myocardial infarction (AMI) . HemodynThe derivation cohort was 2,536 patients with acute STEMI presenting within 12 hours from the symptom onset who were consecutively treated with PPCI in Emergency and Critical Care Center, Beijing Anzhen Hospital, Capital Medical University, from November 2007 to June 2018.The inclusion criteria are as follows: (1) patients with acute STEMI presenting within 12 hours from the symptom onset who were treated with PPCI. (2) Age above 18 years and less than 80 years old male and nonpregnant women. The term AMI should be used when there is acute myocardial injury with clinical evidence of acute myocardial ischemia and with detection of a rise and/or fall of cTnI values with at least one value above the 99th percentile upper reference limit and at least one of the following: (1) symptoms of myocardial ischemia and (2) new ischemic ECG changes [The exclusion criteria are as follows: (1) patients who received thrombolysis; (2) patients who received bivalirudin; (3) cardiac rupture, ventricular septal perforation, and other mechanical complications; (4) pulmonary embolism, aortic dissection, and acute cerebrovascular disease.Prior to emergency angiography, all patients received 300 mg of aspirin, 300 to 600 mg of clopidogrel, or 180 mg of ticagrelor and unfractionated heparin .Periprocedural bradycardia means preoperative heart rate was ≥50 times/min, and intraoperative heart rate was <50 times/min persistent or transient . PreoperWe selected 10 predictor variables for inclusion in our prediction rule from the larger set according to clinical relevance and the results of baseline descriptive statistics in our cohort of patients treated with PPCI. 10 potential candidate variables were age, gender, hypertension history, diabetes history, coronary artery disease history, culprit vessel site, completely block of culprit vessel, no-reflow, intraoperative hypotension (preoperative blood pressure was ≥90/60 mmHg and intraoperative blood pressure was <90/60 mmHg persistent or transient), and using thrombus aspiration devices during operation.Angiographic criteria were used for the diagnosis of no-reflow . No-refln (%). We kept all continuous data as continuous and retained on the original scale. We used univariable and multivariable logistic regression models to identify the correlates of periprocedural bradycardia during PPCI. We entered all variables of We presented data as mean ± SD or Discrimination was the ability of the model to differentiate between patients who do and do not experience periprocedural bradycardia during the study period. This measure was quantified by calculating the area under the receiver operating characteristic (ROC) curve (AUC).P value <0.05 was considered statistically significant.We performed statistical analyses with STATA version 15.1 . All tests were two-sided, and a During PPCI procedure, 434 patients had periprocedural bradycardia , and 2102 patients had no periprocedural bradycardia (control group). The results are shown in We used univariable and multivariable logistic regression to identify predictors of periprocedural bradycardia during PPCI. We identified 8 variables , the culprit vessel was right coronary artery (RCA), completely block of culprit vessel, using thrombus aspiration devices during operation, no-reflow, and intraoperative hypotension) as predictors of periprocedural bradycardia in univariable analysis. After application of the backward variable selection method, 6 variables  remained as significant independent predictors of periprocedural bradycardia during PPCI. The results are shown in Tables P=1/(1 + exp(−(−1.937454 + −0.6133868 ∗ G + 0.8055676 ∗ CBCV + 0.6712826 ∗ CNR + 0.556152 ∗ TA + −1.932014 ∗ LAD + 2.308423 ∗ IH))). LAD = the culprit vessel was LAD, CNR = no-reflow, G = gender, CADH = coronary artery disease history, CBCV = completely block of culprit vessel, IH = intraoperative hypotension, TA = using thrombus aspiration devices during operation, 0 = No, and 1 = Yes; women are coded as 0 and men as 1.According to the above risk factors, we can calculate the predicted probability of periprocedural bradycardia using the following formula: We drew the receiver operating characteristic (ROC) curve . The areAll types of conduction disturbances can occur in the context of an AMI, and these are influenced by multiple mechanisms (often concomitant) including ischemia, extent and location of myocardial infarction, reperfusion, and autonomic effects affecting electrical conduction or the sinus or atrioventricular node . We inveThe culprit vessel is RCA or left circumflex coronary artery (LCX) is more likely to evoke periprocedural bradycardia. On the one hand, RCA or LCX supply blood to the sinus node and atrioventricular node. Ischemia reperfusion can induce the apoptosis of sinoatrial node cells , 6, so bIntraoperative hypotension is often accompanied with periprocedural bradycardia. Acute inferior myocardial infarction often induces transient sinus bradycardia through vagal enhancement, known as Bezold–Jarisch reflex, which is explained by preferential distribution of the vagal nerve in the inferior wall . The BezPeriprocedural bradycardia is associated with epicardial reperfusion but may also be a sign of no-reflow on coronary angiography. Myocardial reperfusion can evoke excitation of cardiac vagal nerve endings , which cUsing thrombus aspiration devices during operation is more likely to evoke periprocedural bradycardia. Periprocedural bradycardia associated with mechanical thrombectomy using the AngioJet device is a well-described phenomenon in the coronary vessels , 18. GlySome patients were enrolled >10 years ago; thus, their treatment may not conform to current standards and techniques.No-reflow, the culprit vessel was not LAD, using thrombus aspiration devices during operation, gender, completely block of culprit vessel, and intraoperative hypotension may be independent risk factors for predicting periprocedural bradycardia during PPCI in patients with acute STEMI.


To report first clinical experience on three cases of congenital complete heart block and the use of a pacemaker system with a maximum lower rate interval of 95 beats per minute.We retrospectively analyzed three patients treated with a pacemaker system with a maximum lower rate interval of 95 beats per minute suffering from congenital complete heart block. We report a follow up period of 2.9 years, focusing on the patients’ growth, development, and adverse events, as well as pacemaker function.In all three patients pacemaker function was impeccable, including minute ventilation sensor rate adaption. All patients showed limited growths as expected, adequate development, good feeding tolerability and circadiane heart rate adaption. One patient experienced skin traction and revision. All patients showed high aortic velocity time integral values after birth.The use of a pacemaker system with a maximum lower rate interval of 95 beats per minute in infants suffering from congenital complete heart block and showing high aortic VTI values seems to be feasible and to result in limited growths but adequate development.The online version of this article (10.1186/s12887-019-1494-7) contains supplementary material, which is available to authorized users. Indications for pacemaker therapy in the neonatal population are rare and include e.g. congenital complete heart block (CCHB) . In the We hypothesize that an LRI of 630 ms allows for adequate growths and development in newborns and infants suffering from CCHB, as cardiac output adaption took place prenatally.We performed a systematic literature research on the use of the Kora pacemaker system in neonates using PUBMED, and MEDLINE databases. To the best of our knowledge, we are the first to report on 3 infants in whom a single chamber pacemaker system with an LRI of 630 ms was implanted. We report a follow up period of 2.9 years, focusing on the patients’ growth, feeding behavior and development, as well as pacemaker function, parameters, heart rate modulation and function of minute ventilation sensors. Furthermore adverse events and tolerability of the pacemaker system were reviewed. The following conditions were considered adverse events: infection of the implantation site, skin traction, suture dehiscence, lead breakage/dislocation, development of scar tissue resulting in elevation of threshold levels. For assessment of growth World Health Organization growth charts were used. Pacemaker parameters and minute ventilation sensor function and activity were analyzed using follow up protocols saved in our database. Protocols were assessed regarding programmed mode, upper rate interval, intrinsic rhythm, impedance, threshold, amplitude, impulse duration, circadian heart rate distribution and rate adaption during crying and feeding. Furthermore we assessed echocardiography records regarding aortic velocity time integral as an estimator on cardiac output.Demographics on included patients are given in Table Except for skin traction, which resulted in suture dehiscence and pacemaker revision in one newborn, no adverse events occurred. All devices were programmed in VVIR mode at the lower rate interval of 630 ms. Initial programming specifics are given in Table We report on the first clinical experience with a single chamber pacemaker system limited to an LRI of 630 ms in neonates. Due to the limited LRI in this single chamber system, concern has been raised whether it would suffice for neonates to contain adequate cardiac minute volume and growth. As we detected high values of aortic VTI (cardiac output marker) in immediate postnatal echocardiography studies, we assumed that cardiac output compensation was sufficient to use a maximum LRI of 95 bpm. Cardiac output compensation seems to have evolved prenatally due to low fetal heart rates. Along follow up VTI values remained constant. We hoped that the smallness of the device would aid in regard of a better implantation outcome and tolerability concerning skin traction and suture dehiscence. Furthermore we figured an MV sensor for rate response being a good choice in infants. Despite using such a small system suture dehiscence accounted for a complication rate of 33%. The function of MV sensors in children with pacemakers has been shown by Cabrera et al. [A limited lower rate interval of 630 ms seems to suffice adequate development along with limited growth in infants suffering from CCHB, given a sufficient adaption in neonatal stoke volume. Size of the pacemaker system, however, seems crucial in regard of complications like skin-traction and surgical revision. Never the less, children remain poor cousins in cardiac device therapy. Small sized pacemaker systems meeting pediatric needs in terms of rate limitation and implemented sensors are warranted, especially for the neonatal population. However, such devices are still lacking.Additional file 1:Figure S1. Growth centile curves of patient 1. Legend: x-axis shows age in months, y-axis shows body weight in kilogram on the right lower side and body lengths in cm on the left and right upper side. (JPG 930 kb)Additional file 2:Figure S2. Growth centile curves of patient 2. Legend: x-axis shows age in months, y-axis shows body weight in kilogram on the right lower side and body lengths in cm on the left and right upper side. (JPG 931 kb)Additional file 3:Figure S3. Growth centile curves of patient 3. Legend: x-axis shows age in months, y-axis shows body weight in kilogram on the right lower side and body lengths in cm on the left and right upper side. (JPG 935 kb)Additional file 4:Figure S4. Denver Developmental Screening Test results of patient 1. (PDF 198 kb)Additional file 5:Figure S5. Denver Developmental Screening Test results of patient 2. (PDF 202 kb)Additional file 6:Figure S6. Denver Developmental Screening Test results of patient 3. (PDF 198 kb)