Blood flow in hemodialysis catheters: a numerical simulation and microscopic analysis

of in-vivo formed fibrin

Thabata Coaglio Lucas1*, Francesco Tessarolo2, 3, Victor Jakitsch1, Iole Caola4, Giuliano

Brunori5, Giandomenico Nollo2,3, Rudolf Huebner1

1Department of Mechanical Engineering, Universidade Federal de Minas Gerais, Antônio

Carlos Avenue 6627, 31270901 Belo Horizonte, Brazil.

2Department of Industrial Engineering, BIOtech Centre on Biomedical Technologies,

University of Trento, via delle Regole 101, 38123 Mattarello Trento, Italy

3Healthcare Research and Innovation Program (IRCS), Bruno Kessler Foundation, via

Sommarive 18, 38100, Povo, Trento, Italy

4Section of Electron Microscopy, Department of Medicine Laboratory, Azienda Provinciale

per i Servizi Sanitari di Trento, via Degasperi 79, 38123 Trento, Italy

5Department of Nephrology, S. Chiara Hospital, Largo Medaglie d’oro 7, 38123 Trento, Italy

Running Title: Blood flow in hemodialysis catheters

*Corresponding Author:

Thabata Coaglio Lucas, Bioengineering Laboratory, Department of Mechanical Engineering

Federal University of Minas Gerais. Antônio Carlos Avenue 6627, 31270901 Belo Horizonte

MG (Brazil).Tel: +55 31 34096677 / +5531 98280379 Fax: +55 31 344337 Email:

thabataclucas@yahoo.com.br

 

 

Abstract

Although catheters with side holes allow high flow rate during hemodialysis, they also

induce flow disturbances and create a critical hemodynamic environment that can favor fibrin

deposition and thrombus formation. This study compared the blood flow and analyzed the

influence of shear stress and the shear rate in fibrin deposition and thrombus conformation in

non-tunneled hemodialysis catheters with unobstructed side holes (unobstructed device) or

with some side holes obstructed by blood thrombi (obstructed device).

Computational fluid dynamics (CFD) was performed to simulate realistic blood flow under

laminar and turbulent conditions. The results from the numerical simulations were compared

to the fibrin distribution and thrombus architecture data obtained from scanning electron

microscopy (SEM) and two photons laser scanning microscopy (TPLSM) on human

thrombus formed in catheters removed from patients. CFD showed that regions of flow

eddies and separation were mainly found in the venous holes region. TPLSM

characterization of thrombi and fibrin structure in patients sample showed fibrin

conformations in accordance with simulated flux dynamics. Under laminar flow conditions,

the wall shear stress close to holes border increased from 87.3±0.2 Pa in the unobstructed

device to 176.2±0.5 Pa in the obstructed one. Under turbulent flow conditions, the shear

stress increased by 47% when comparing the obstructed to the unobstructed catheter. The

shear rates were generally higher than 5000s-1 and therefore sufficient to induce fibrin

deposition. This findings were supported by SEM data documenting a preferential fibrin

arrangement on side hole walls.

Keywords: central venous catheter, hemodyalisis, computational fluid dynamics, thrombus,

electron microscopy, two photon laser scanning microscopy

 

 

INTRODUCTION

Although arteriovenous fistulas are recommended as the preferred choice of vascular access

for hemodialysis patients, many patients dialyze using central venous catheters (CVCs)

(1,2). Catheters guide the venous blood towards the hemodialysis machine and return it to the

venous system in a safe and reliable way. Among different approaches the internal jugular

vein seems to have less adverse effects (2, 3, 4). The residence time for a non-tunneled

catheter in the vein should theoretically not exceed two weeks (3). However, in clinical

practice, catheters typically remain within a patient’s vein at least 2 months to ensure

sufficient lead-time for fistula maturation. During the residence time, blood thrombi and

fibrin sheet can develop, impairing lumen patency and catheter performance (4). Non-

tunneled catheters are generally designed with side holes facing different directions on the

catheter wall. These side holes may affect the flow pattern, shear stress distribution, flow

separation and recirculation, thus providing a critical hemodynamic environment. Blood

velocity and pressure fluctuations close to the catheter surface could cause a transition from a

laminar to a turbulent flow regime and promote the onset of the blood clotting process with

thrombus formation in the catheter holes. Microscopic structure and composition of venous

thrombi have been previously reported (5,6), and we recently showed that blood thrombi

formed in a hemodialysis CVC present a complex structure and are mainly composed of

fibrin and erythrocytes (7). A fibrin sheath commonly develops within 24 hours after CVC

insertion at the region where the catheter contacts the vessel wall, and may completely

enclose the proximal portion of the CVC in five to seven days (8,9). Fibrin conformation is

remodeled by shear stress and mechanical stimuli, which alter the structural integrity of the

thrombus. Hemodynamic forces are modulators of fibrin structure and result in the

development and progression of the thrombus.

 

 

Numerical models based on computational fluid dynamics (CFD) allow the investigation of

hemodynamic factors that lead to thrombus formation and progression. Given the

microscopic three dimensional architecture of catheter thrombi, scanning electron

microscopy (SEM) has been reported as a powerful technique for characterizing the surface

structure (10-12). SEM has been applied for understanding the extent of fibrin network

formation and the organization of fibrin fibers over the thrombus and in the CVC holes (10).

The same technique can provide information on the alignment of fibrin fibers in proximity to

the CVC holes and tip, reflecting blood flow pattern in these crucial catheter regions. Two

photon laser scanning microscopy (TPLSM), an optical technique based on fluorescence

induced by two photon absorption, has recently been proposed as an additional tool for

characterizing blood thrombi in a fully hydrated state without the need to section the sample

(7,8). Recently we reported the potential of TPLSM in imaging the fibrin structure of thrombi

in CVCs removed from patients (6). All these methods allow a thorough investigation of the

local blood flow and dynamic conditions in relation to fibrin deposition and blood thrombus

formation resulting from the CVC presence in the vein.

To provide further insights about the role of blood flow and catheter design in modulating

thrombus formation, this work compared the blood flow in non-tunneled hemodialysis

catheters with and without obstruction of some side holes and analyzed the influence of shear

stress and the shear rate in fibrin conformation. The results obtained by CFD were discussed

in the light of the fibrin patterns obtained by SEM and TPLSM on a set of non-tunneled

hemodialysis CVCs removed from patients.

 

 

MATERIALS AND METHODS

Three-dimensional geometry and computational mesh

A realistic CVC model, similar to the clinical device MedCOMP/HEMO-CATH®

(Harleysville, PA), was designed in SolidWorks™ (Solidworks Inc.,Concord, MA). The CVC

design was characterized by a dual lumen system, arranged in a single shaft 2 mm in

diameter and 20 cm in length. Fig. 1a shows the realistic dual-lumen CVC that contains one

outflow lumen with three venous outlet holes and a second inflow lumen with four arterial

inlet holes. The internal cross sectional areas of both the arterial and venous lumens were

2.93 mm2. Catheter holes were labeled as showed in Fig.1. In the model of the unobstructed

CVC, each hole was completely pervious. In the model for the obstructed CVC, venous holes

#2 and #3 and arterial hole #6 were completely obstructed by a thrombus-type occlusion (red

areas in Fig. 1b). In the obstructed CVC model, also the venous and arterial lumens were

partially obstructed, reducing the original cross sectional area to 2.04 and 1.33mm2,

respectively. According to this configuration, the measurements of flow parameters were

obtained for holes #1, #4, #5, #7, and #8. A realistic model of the internal jugular right and

left veins and superior vena cava up to the right atrium were obtained by segmentation of CT

images from the Visible Human Project® (U.S National Library of Medicine, Bethesda, MD)

(Fig. 2). Two hundred axial anatomical images (2048 x 1216 pixels) at 1 mm intervals were

used to create the jugular veins geometry. The vein model was imported into SolidWorks™

and bounded by interpolating polynomial curves. A mesh independency study was carried

out performing simulations with coarse to fine mesh sizes until results were no longer

influenced. The results exhibited a maximum of 3% difference. The mesh resolution was

maximized near the catheter and vein walls and consisted of approximately 1.3x106 nodes,

5.0x106 tetrahedral elements and 6.3x102 pyramidal elements. A boundary layer consisting

of eight rows was established, with an expansion factor of 1.2 and a total depth of 1 mm.

 

 

Boundary conditions and numerical simulation

A laminar and turbulent transient flow, k–ω Shear Stress Transport (SST) model, were

implemented in ANSYS CFX® 12.1 (ANSYS-Fluent Inc., Lebanon, NH, USA). The SST

model was designed to give highly accurate predictions of the onset and the amount of flow

separation under adverse pressure gradients by the inclusion of transport effects into the

formulation of the eddy-viscosity. Therefore, this model was substantially more accurate for

determine the wall shear stress in the internal hole wall of the catheter. The time dependent

and incompressible Navier-Stokes and Carreau-Yassuda (C-Y) model equations were used

for the mathematical description (9, 10). The following set of parameters was used for blood

analog fluid: µ∞= 3.45 x 10-3 Pa·S, µ0=56 x10-3 Pa .S, a=1.25, n=0.22, and λ=1.19 s (9) .

Where µ∞ is the viscosity at infinite shear rates, µ0 is the zero shear rate viscosity, λ is the

relaxation time in seconds, a and n parameters can be varied to match the shear thinning

behavior of blood. Blood viscosity in humans is non-Newtonian. With increasing shear rate,

the viscosity decrease. In this work the values of µ∞ and µ0 were obtained from optimum

hematocrit in the physiological plasma viscosity. The viscosity increases exponentially with

increasing hematocrit. The C-Y model is widely accepted and used to correct the variations

in blood viscosity caused by a variation in hematocrit. Blood density was set equal to 1050

kg/m3. The values for the turbulent kinetic energy, specific rate of turbulence, eddy

frequency, and turbulent viscosity were modeled according to Abraham et al. (10). The

duration of a cardiac cycle was assumed to be 0.8 s, yielding a heart rate of 75 beats per

minute. A second-order backward Euler method was used for the time integration. The

doubling time step was performed, and no obvious differences in the results of the mass flow

rate in the veins were observed. A constant time step was employed, where ∆t = 0.0005 s

with 1600 total time steps per cardiac cycle. The waveforms of the jugular blood inlet

 

 

velocity and atrium pressure are presented in Fig. 2 and were approximated from the

experimental curves found in the literature and intravenous color Doppler lines (11). The (A)

wave occurs when the atrium contracts, increasing atrial pressure. In the same time, the blood

is propelled in a retrograde direction toward the veins. When the tricuspid valve closes, the

systole wave (S) occurs. The transitional (V) wave corresponds to atrial overfilling against a

closed tricuspid valve, anticipating the opening of valve in the diastole (D). The inlet and

outlet waveforms designed for this study showed a correlation between flow velocity in the

internal jugular veins and the pressure in the right atrium. The model was initially configured

with flow data from a steady-state simulation with an inlet velocity of 0.18 m/s for the right

internal jugular vein and 0.20 m/s for the left internal jugular vein. The outlet pressure of the

vein was 533.28 Pa. Mass flow in the CVC venous lumen was set to 0.0052 kg s-1. The outlet

for the CVC arterial lumen had a pressure of 33330.60 Pa. Parameters were set according to

National Kidney Foundation guidelines for hemodialysis catheters (3). The results obtained

for the steady-flow calculations were used as the initial values for unsteady flow, and five

cycles of calculations were conducted for checking variations in the parameters. A no-slip

boundary condition was imposed along the vessel walls. The continuous variables were

expressed as means ± SDs.

Characterization of thrombus and fibrin structure in CVC removed from patients

Samples of real CVC devices with blood thrombi obstructing some of the side holes were

obtained from a collection of 23 non-tunneled CVC (MedCOMP/HEMO-CATH®,

Harleysville, PA; (median insertion time 17 (5-150) days) gathered during a recent

observational clinical trial performed from August 2011 to July 2012 at the Nephrology

Department of the Trento Hospital in Italy. The study was approved by the Ethics Committee

Concerning Research Involving Human Subjects and Clinical Trials (Process 04/2010) of

Trento, Italy. Samples were treated as previously reported (6). Briefly, CVCs were rinsed in

 

 

sterile saline and fixed in 4% buffered formaldehyde immediately after removal from the

patient. Segments of the CVC containing a venous catheter hole were considered for analysis

by SEM and TPLSM. SEM analysis was conducted after sample dehydration in ascending

idro-alcoholic solutions, drying in a laminar flow cabinet overnight and sputter-coating with

gold. Sample imaging was performed with a XL30 ESEM FEG scanning electron

microscope (FEI, Philips, Nederland). Samples for TPLSM were washed for 2 min in

deionized water and stained by immersion in 1mM Rhodamine 6G (Sigma-Aldrich) for 20

min in the dark. Sample imaging was performed with a two photon laser scanning

microscope (Ultima IV, Prairie Technologies). An ultra-short pulsed laser (Mai Tai Deep See

HP, Spectra-Physics) served as the light source. The excitation wavelength was set to 800nm,

and the fluorescence was collected in the red band (607+/-20nm). The images were acquired

with a 100x objective lens (NA 1.0, water immersion, Olympus LUM Plan FI) and a dwell

time of 50 µs.

RESULTS

Velocity vectors and streamlines

Representative streamlines for the turbulent flow into the vein in the unobstructed and

obstructed catheter shaft at t=0.4s (corresponding to the systole) are shown in Fig. 3. The

flow pattern was non-uniform and irregular in both the patent and occluded catheter models.

A recirculation zone and stagnation point, due to the outlet flow from the venous hole, can be

observed in the enlarged view of Fig. 3. However, recirculation eddies with reverse

streamlines and changes in direction overtime were markedly more evident in the obstructed

catheter model. The microscopic characterization of the corresponding area by TPLSM and

SEM showed fibrin fibers were not aligned in a parallel fashion but were frequently arranged

 

 

in circles following flow disturbances (Figs. 4a and 5a). Fibrin fibers bifurcate and exhibit a

change in alignment and arrangement. Twisted fibers were also observed.

Wall Shear stress and shear strain rate

The shear stress in the hole wall #4 under turbulent flows (t=0.4s) is shown in Figs. 6a and

6b for the unobstructed and obstructed catheter models respectively. Figs 6a and 6b show the

internal hole wall that was rotated to enable the viewing of peaks in the shear stress that were

high in the wall inside the side holes from obstructed catheters. Maximal shear stress values

reach 1000 Pa between the edge of the bottom of the hole wall and the wall of the catheter

shaft. The average strain rate value in the two catheter model differed of about 35% in the

laminar flow and 38% in the turbulent flow conditions. Arterial hole # 8 was subjected to the

highest wall shear stress and shear strain since the blood velocity was maximal. Wall shear

stress and shear strain rates for the other pervious holes are summarized in Table 1. The SEM

micrograph of the internal venous hole wall (Figs. 4b and 4c) shoved a preferential

deposition of fibrin in accordance to the areas with highest shear stress indicated in Fig. 6b.

Figs. 6c and 6d show the distribution of the shear strain rate in the external surface of the

catheter around venous hole #4 in both the unobstructed and obstructed catheter. Simulations

showed the presence of catheter areas subjected to strain rate higher that 5000 s-1 and 10000

s-1 for the unobstructed and for the obstructed catheter model respectively.

DISCUSSION

The computational parameters provided a flow pattern that was in agreement with the

microscopic arrangement of the fibrin revealed by SEM and TPLSM on in-vivo formed

thrombi. The overall flow patterns within the vein and around the venous side holes

contained regions of flow eddies and flow separation in both the obstructed and unobstructed

catheter modeleds. Previous studies reported that platelet aggregation preferentially occurs at

 

 

regions of low shear stress, i.e. in flow eddies and in flow separation regions (2,12-13). In the

flow eddies zone, the wall shear stress has been reported to remain at a low level (0.0-1.0 Pa)

throughout the cardiac cycle due to the reversed flow conditions (14-15). The velocity

vectors showed that flow eddies were not present in front of the arterial holes, but were seen

in the downstream of the venous hole. Differently, flow separation and stagnation regions

were observed close to the catheter wall, between the arterial holes. These findings suggested

that the flow disturbance in the venous hole regions can occur more frequently than in

arterial hole regions, thus creating conditions for accelerating the development of blood

thrombi (16). In addition, the blood flux among the venous holes is separated into

downstream high velocity jets exiting from the three venous holes and low velocity

circulation along the catheter surface wall, as shown in the enlarged view in Fig. 3. These

conditions are a plausible explanation for the arrangement of the fibrin plaque on the hole

wall, which can be associated with the flow separation region between the three venous

holes. In these areas the fibrin attaches to the catheter wall and propagates along the direction

of flow, generating a lateral aggregation close to the hole border and eventually forming a

fibrin plaque that can obstruct the hole. To prevent catheter clotting, heparin typically is

placed into the catheter lumens at the end of dialysis as a catheter lock. In spite of the fact

that multiple side holes provide good blood flow, they also may allow the catheter lock to

seep out, thus increasing the risk of clotting. Seepage of the catheter locking solution

increases the risk of clotting and thrombus formation between the last of the side holes and

the catheter tip.

Real-time TPLSM imaging of the thrombogenesis process in mice subjected to laser injury of

the arterial intima showed that also the neo-formed thrombus perturbs the blood flow (7).

RBCs and platelets were observed to move slowly in the regions of vortices and were

captured and incorporated into the growing thrombus (7). Although we were not able to

 

 

follow thrombus formation in patients in real time, TPLSM imaging of thrombi formed on

CVCs removed from patients showed fibrin plaques that develop in a corresponding region

containing flow eddies. Furthermore, the presence of circles and twisted layers in the fibrin

plaque we observed by both SEM and TPLSM techniques, suggested the presence of eddy

flow close to the hole rim (2). In this region, the fibers bifurcate, and alterations in the

arrangement of the fibrin in the thrombi were observed also by TPLSM as reported in Fig.

5a. The comparison of flow parameters between unobstructed and obstructed device, showed

that the shear stress value in the internal wall increased by 50% in venous hole #4 under a

laminar regime and by 47% under turbulent flow conditions. It has been reported that higher

shear stress can induce conformational changes in fibrinogen adsorbed onto a polymer, thus

increasing the binding surface area between the fibrinogen and platelet receptors (17-18). At

an elevated shear strain rate, typically in excess of 10,000 s-1, binding between the

glycoprotein receptor of platelets and vWf is known to be unique, with the capability to

recruit fast flowing platelets at a high shear rate in the absence of any other platelet activators

(19). Mareels et al. (20) proposed a model of CVC for numerical study with completely open

either close side holes or reduced the side hole diameter in size. When the diameter of the

side holes was halved the shear rate increased near the side hole walls and in the lumen wall

near the side holes. Conversely, in this study, the side holes that were not closed, in the

obstructed catheter showed an increased of the shear rate in the internal wall of the side hole.

Although in this study the diameter was not reduced, the shear rate was particularly high in

the internal wall of the side hole that was not closed in the obstructed catheter. We found a

shear rate higher than 10,000 s-1 around the venous hole #4 in the obstructed catheter model

and around the arterial hole #8 under both turbulent and laminar flow conditions. This factor

is a plausible explanation for the establishment of additional bonds, leading to elevated fibrin

development and subsequent thrombus formation in these holes. In contrast, at shear rates

 

 

<1000 s-1, platelet aggregation is mediated by interactions between fibrinogen and the

receptor integrin on the platelet membrane (19). This process is independent of vWf binding,

although an in vitro study has demonstrated that, at a low shear rate (250 s-1), complementary

platelet aggregation mediated by vWf can also occur (12). In the literature, high values of

shear stress (10.0-800.0Pa) are known to cause platelet activation in brief exposures of a few

milliseconds (14). In this study, a part form the arterial hole #5, all other inlet and outlet

holes were subjected to non-physiological high shear stress. Side holes are therefore critical

regions for accelerated platelet activation due to changes in blood flow parameters that are

crucial for fibrin propagation. Moreover, at the threshold shear stress for venous homeostasis

(1.4 to 6.0 Pa), platelet activation and aggregation can be enhanced by an increased shear

stress exposure-time (14). Althought short exposure (0.1s in each 90s) to a shear stress of 6.0

Pa did not induced significant changes in platelet activation (21), there were indications that

longer exposure (30 minutes) might lead to a considerable amount of thrombin generation

and accelerated platelet production (21). Moreover, the thrombus can be initiated under low

shear stress (<0.41Pa) and low shear strain rate (<54 s-1) conditions when the blood flow is in

continuous contact with implantable polymeric devices (22). In this study, we found that high

shear stress is predominant in the side holes and can be a mechanical precursor for the

formation of hemodialysis catheter thrombi. The regions of disturbed flow marked by flow

separation and flow eddies zones with low values of shear stress could contribute to an

increased residence time of blood components, which may subsequently result in the

stimulus of platelet aggregation by prolonged contact with the polymeric CVC (1,14). More

complex mechanisms of activation were also proposed for the formation of thrombi in

presence of mechanical heart valve. Platelets that were exposed for short periods of time to

high shear stress (200-240Pa) did not recover their normal quiescent activity state under

subsequent conditions of low shear stress (17). This sensitization behavior was proposed for

 

 

understanding the high risk of thromboembolic complications in patients with mechanical

heart valves (17). These findings also provide insights into the complex dynamic flow when

a CVC remains in the vein, which includes combined zones of very high shear stress and

flow eddies. Some limitations of the study should be considered. First, the arterial holes were

not assessed by SEM and TPLM. Preliminary data on a sample subset did not show

significant differences in composition between thrombi formed on venous or arterial CVC

holes. However, future studies should evaluate the difference between the venous and arterial

holes with respect to the fibrin conformation and arrangement. Second, the corpuscular

microstructure of blood and the possible fluid-structure interactions (cell-to-cell collisions

and cell-surface adhesion) were not taken into account in the CFD simulation and can

influence the flow pattern and the flow dynamic. However, the hypothesis that the fibrin

conformation could be associated with local hemodynamic conditions was supported by

SEM and TPLM, giving a first local validation to the flow pattern found by the simulation

model.

To our knowledge, this was the first study to investigate the dynamics of blood flow in the

side holes of CVCs and to compare results from computational fluid dynamics to fibrin

conformation in thrombi formed on hemodialysis catheter retrieved from patients.

The methodology applied in this study can provide insights for improvements in the design

and performance of existing and new designs for hemodialysis CVC.

ACKNOWLEDGMENTS

The authors thank CAPES /Process BEX: 1014/11-0 for the financial support.

Clinical specimen were collected in the research project “RuBiCat - Role of the Biofilm and

device modification for the Prevention of Catheter related infections”. The project was co-

 

 

funded by the Autonomous Province of Trento, Grant 2008, and Fondazione Cassa di

Risparmio di Trento e Rovereto, Young investigator Grant 2009.

 

 

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Table 1: Shear stress in the internal wall of the side holes and shear strain rate close to the hole

rim.

LaminarFlow TurbulentFlow

Wall shear stress (Pa)

(mean ±SD)

Shear strain rate (s-1)

(mean±SD)

Wall shear stress (Pa)

(mean ±SD)

Shear strain rate (s-1)

(mean±SD)

Unobstructed

catheter

Obstructed

catheter

Unobstructed

catheter

Obstructed

catheter

Unobstructed

catheter

Obstructed

catheter

Unobstructed

catheter

Obstructed

catheter

Venous hole #4 87.3±0.2 176.2±0.5 8047±18 12400±7 130.4±0.2 244.1±0.1 6449±18 10497±5

Arterial hole #8 278±3 284±4 16000±150 6000±150 295±3 299±3 11103±130 11090±120

Arterial hole #7 70.9±1.5 72±2 5302±60 5795±78 64±1 66.7±1.5 3477±54 3794±59

Arterial hole #5 2.5±0.3 7.3±0.3 391±39 1014±32 2.0±0.3 5.5±0.3 272±33 632±40

Note: Spatial distribution were calculated at t=0.4s. Wall Shear stress and shear strain rate

values were averaged on the entire internal wall of the side hole as seen in Fig 6a and b.

 

 

Figure 1: The realistic model of an unobstructed (left) and obstructed (right) hemodialysis

CVC used in the CFD simulations. The numbers indicate the labels as referred in the text:

venous hole tip (#1), three venous side holes (#2, #3, #4), four arterial side holes (#5, #6, #7,

#8). The red areas in the obstructed CVC indicate the shape of the simulated CVC thrombi,

obstructing arterial hole (#6) and venous holes (#2 and #3).

 

 

Figure 2: Design and boundary conditions of the CFD model. Three-dimensional geometry

of the jugular veins (ligh grey) and the intravenous portion of the hemodialysis catheter (dark

grey). Arrows indicate blood inlets (where velocity data presented in the upper graph were

applied) and the blood outlet (where pressure data presented in the lower graph were applied

as a boundary condition). The temporal evolution of velocity and pressure during a whole

cardiac cycle are reported indicating diastole (D), atrial contraction (A), systole (S), atrial

filling (V).

 

 

Figure 3: Streamline profiles in the vein and within CVC holes for the unobstructed (left)

and obstructed catheter (right). The central part of each image show a comprehensive view of

the vein and the proximal portion of the catheter, where side holes are located. Insets present

a closer view of venous (V) and arterial (A) side hole.

 

 

Figure 4: Scanning electron microscopy images of fibrin formed in-vivo on CVC removed

from patients. (a): fibrin plaque on the catheter surface around a venous side hole. The

heterogeneous direction of fibrin strands, with branching points are related to the presence of

high share strain rate as shown in Fig. 6d. (b): Unobstructed venous side hole (corresponding

to position #4) with initial deposition of fibrin layer. Fibrin was found preferentially on the

hole wall, coherently to the area with highest shear stress indicated in Fig. 6b. (c): Closed up

view of the initial fibrin layers formed on the wall of the side hole. Original magnification

500x (a), 29x (b) and 500x (c).

 

 

Figure 5: Two Photons Laser Scanning Microscopy of fibrin plaque formed in-vivo on CVC

removed from patients. a): fibrin plaque on the catheter surface around a venous side hole, in

the corresponding area where share strain rate was high, as shown in Fig. 6d. The dashed line

indicate approximately the position of the hole rim. Fibrin plaque is thicker and more

structures close to hole. Only sparse fibrin network is present in the areas subjected to low

shear strain rate (lower part of the micrograph). Round shaped structures close to hole rim are

white blood cells. b): Fibrin plaque microstructure in an area subjected to high share strain

rate. Fibrin strands and bundles have anisotropic directions and many branching point are

present (arrows) forming circles of fibers. Bar is 25 µm.

 

 

Figure 6: Distribution of shear stress and strain at CVC surface of venous hole #4. Wall

shear stress on the hole profile under turbulent flow for the unobstructed (a) and obstructed

(b) catheter scenario. Shear strain rate around the hole rim for the unobstructed (c) and

obstructed (d) catheter.