Supporting InformationWestein et al. 10.1073/pnas.1209905110SI MethodsMaterials. Alexa Fluor (AF)-568–labeled monoclonal antibody(mAb) against mouse P-selectin (α-CD62P) was from EmfretAnalytics. Carboxyfluorescein diacetate succinimidyl ester (CFSE),Fluo-4 acetoxymethyl ester, 3,30-dihexyloxacarbocyanine iodide(DiOC6), AF-546–labeled fibrinogen, FITC-labeled goat anti-mouse IgG, and AF-488 phalloidin were from Invitrogen. Plasmapurified vWFwas fromUniversityMedical CenterUtrecht (Utrecht,The Netherlands). Polydimethylsiloxane (PDMS) (Sylgard) wasfrom Dow Corning. Mouse anti-vWF monoclonal antibody wasfrom Santa Cruz.

Animals. Wild-type C57BL/6 mice of either sex were obtainedfrom Charles River. All experiments were approved by theMaastricht University animal experimental and care committees.

In Vivo Stenosis Model. C57BL/6 mice were anesthetized by s.c.injection of ketamine and xylazine (0.1 and 0.02mg/g body weight);body temperature was held at 37 °C. The right or left carotid arterywas carefully dissected free from surrounding tissue, and the an-imal was injected i.v. with CFSE-labeled platelets from a donormouse, as described (1). Labeled platelets represented 10–15% ofthe total circulating platelet pool. Mild vessel damage was in-duced by topical application of 12.5% (wt/vol) FeCl3 for 30 s.Local stenosis was achieved by manual indentation of the carotidwall with a 27-gauge needle controlled by a manual microma-nipulator. P-selectin expression was detected with AF-568–con-jugated anti-P-selectin mAb (injected at 2.25 μg/g); fibrin(ogen)with AF-546–labeled fibrinogen (injected at 3.4 μg/g). Compoundswere injected i.v. at 10 min before the experiment.Multiphoton fluorescence microscopy was used because of its

high penetration depth due to the high excitation wavelength.Thrombus composition was monitored in real time (50 frames/s),using a Leica SP5 multiphoton system integrated with a resonant-scanning module (8,000 line scans/s). Multiphoton fluorescencewas excited at 800 nm with a 120 ± 20-fs–pulsed Ti:sapphire laser.Samples were observed with a HCX APO L 20× water-immersionobjective (numerical aperture, 1.0), field of view of 435 × 435 μm,and captured at 50 Hz as 512 × 512 pixels. Fluorescence fromCFSE and AF dyes was detected by two direct photomultipliertube detectors at 525 ± 25 nm (CFSE) and 585 ± 20 nm (AF-546and -568). Postprocessing of multiphoton images was done withLeica LAS AF suite software (median filter with 1 pixel radiusand contrast enhancement).

Microfluidic Stenosis-Containing Chips. The newly designed PDMSmicrofluidic chips contained flow microchannels with dimensionsof 52 μm high, 300 μm wide, and 17 mm long. The microchannelhad a stenotic region producing 20%, 40%, 60%, or 80% oc-clusion, due to a half-circular insertion of 600–1,000 μm in di-ameter. Chips were manufactured by soft lithography, using a 52-μm-high SU-8 master mold that was fabricated by standardphotolithography methods. Briefly, a quartz/chromium mask wasdeveloped containing the above-mentioned geometries and wasused with UV illumination to transfer the patterns onto a wafercontaining a thin spin coated photoresist layer (UV-sensitivepolymer, also termed SU-8 master). The UV-sensitive photoresistis first subjected to soft baking for hardening by solvent degassingand then brought in close contact with the mask using a maskaligner. UV exposure through regions of the mask without achrome layer leads to the local activation of the photoresist(soluble). Unexposed areas stay insoluble, allowing the formation

of an inverted image of the mask onto the wafer. This SU-8master wafer can then be used for PDMS replication. PDMS(Sylgard 184) prepolymer was mixed with curing agent at a ratioof 10:1 (wt/wt), and thoroughly degassed. The PDMS materialwas poured onto the master mold and cured at 60 °C overnight.PDMS chips were peeled off, and 1-mm inlets and outlets wereopened using a blunt needle. The chips with microchannels weresealed with precoated glass coverslips by either hydrophobic in-teraction or plasma bonding (40 s at 100 W, 400 mTorr; HarrickPlasma), and blocked with 1% (wt/vol) BSA.Wall shear rates of the input flow for these microchannels were

calculated from volumetric flow rates by using the equation, γ= 6Q/wh2. Herein, γ is the wall shear rate (in seconds−1), Q is the vol-umetric flow rate (in milliliters per second), w is the channel width(in centimeters), and h is the channel height (in centimeters).

Endothelial Cell Culture in Microchannels.Human umbilical vascularendothelial cells (HUVECs) (passages 3–6) were cultured inEGM-2 SingleQuots medium (Lonza) and grown to 80–90%confluency before trypsinization. Trypsinized cells were seeded(20 × 106 cells/mL) into microfluidic stenosis channels precoatedwith 2 mg/mL fibronectin (kind gift from Sanquin), and un-attached cells were flushed out after 30-min incubation at 37 °C.The HUVECs were cultured inside the channels overnight at37 °C and 5% CO2 atmosphere. Human blood or EBM-2 basalmedia was perfused through HUVEC-covered channels, as in-dicated. Staining for release of vWF was performed by addingmouse anti-vWF monoclonal antibody (1:100) and FITC-labeledgoat anti-mouse IgG (1:100) to perfusion medium. Controlswere performed by staining with secondary antibody only.Release and surface expression of vWF were visualized in

real time by fluorescence microscopy (Leica DMI5000M withDFC310Fx camera). For actin staining, HUVECs inside the chipwere fixed by perfusion with 3.7% (wt/vol) paraformaldehyde inPBS and incubating for 15 min at room temperature. Then, thecells were stained with 1 μg/mL AF-488–labeled phalloidin inphysiological salt solution with 0.1% (wt/vol) Triton X-100 for 30min at 37 °C. Staining patterns were analyzed with a Zeiss LSM5confocal microscope.

In Vitro Flow Perfusions. Time-lapse fluorescence images duringwhole-blood flow throughmicrofluidic chambers were obtained intile-scan mode, using a Zeiss LSM Life-7 confocal microscopesystem. Rises in cytosolic Ca2+ were monitored during flow inadherent platelets, preloaded with 4 μM Fluo-4 acetoxymethylester, as previously described (2). Imaging was at 2 frames/s.Numbers of platelets with increased Ca2+ were determinedby fluorescence thresholding and time-averaging over 30 s. Tomeasure stenosis-specific increases in platelet-bound vWF, pla-telets were allowed to adhere to a continuous coated surface of100 μg/mL fibrinogen under static conditions for 1 min. Platelet-depleted blood was prepared by centrifugation of citrate anti-coagulated whole blood at 300 × g for 10 min and replacing thePRP fraction with plasma. Platelet-depleted blood was perfusedthrough the stenosis channels at 1,000 s−1 input wall shear or at5,700 s−1 and 8,000 s−1 through straight channel for 4 min fol-lowed by fixation with 4% (wt/vol) paraformaldehyde. Rates of5,700 s−1 and 8,000 s−1 were chosen because these are the wallshear rates present at the stenosis outlet and apex, respectively.Platelets and vWF were stained with 0.5 μg/mL DiOC6 and 100μg/mL rabbit-anti-human vWF IgG plus 2 μg/mL TRITC-con-jugated goat anti-rabbit IgG, respectively.

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Platelet Velocities. Flow velocities of single free-flowing platelets,labeled with DiOC6, were obtained with fast line scanning con-focal microscopy by measuring travel trajectories of plateletsflowing 2 μm above the channel bottom. Due to the high velocity,each platelet appeared as a streak in single line scans of 96 μs/line.The length of this streak was corrected for the platelet diameter(2 μm), and converted to a velocity value by dividing by 96 μs. Foreach platelet, an x,y position was determined. Velocity data andx,y positions were combined in a color-coded velocity map.

Computational Fluid Dynamics and Calculations. Fluid dynamics weresimulated with a Comsol Multiphysics 4.2 laminar flow module.The simulations considered the microfluidic chamber dimensions(channel height, 52 μm; width, 300 μm); and stenotic occlusionsranging from 80% to 20% with diameters of 600 and 1,000 μm.A laminar flow state was described by solving Navier–Stokesequations. The inlet velocity was set to 5,555, 9,259, and 18,518μm/s, resulting in wall shear rates of 600, 1,000, and 2,000 s−1,respectively. A mesh comprising 652,000 uniformly-sized tri-angular elements was applied to each model. For all simulations,no-slip boundary conditions and Newtonian fluids were assumed.Fluid traveling through the stenosis inlet experiences an in-

crease in forward velocity due to the constant volumetric flow rateand the decrease in cross-sectional area. This increase in forwardvelocity leads to flow elongation, or flow extension. Basically, thismeans that the fluid experiences strain in the flow direction. Thestrain rate or elongation rate _E is given by the change in velocityover the distance, or dv

dx. Because the geometry of the channel iswell defined, the elongation rate can be calculated for everyposition in the stenosis in the following way. x is the positionalong the channel direction, y is the width of the channel, and r isthe radius of the stenosis.The geometry of the channel region from nonstenosis to

maximal stenosis follows from the definition of a circle with radiusr in an x–y coordinate system as follows:

x2 + y2 = r2

or

yðxÞ =ffiffiffiffiffiffiffiffiffiffiffiffiffir2 − x2

p:

To match the exact geometry of the channel, this general formneeds two additional parameters: a, the length of the stenosis(shifting the circle to the right in the coordinate system); and b,the remaining width of the channel at maximal stenosis (shiftingthe circle upward in the coordinate system). Now, the start of the

stenosis is x = 0 and the side of the channel is y = 0. Fora schematic visualization of these parameters, see Fig. S5B:

yðxÞ = ðr+ bÞ−ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffir2 − ðx− aÞ2

q:

Flow velocity in the direction of x is inversely proportional tothe width of the channel y(x). If we define the flow velocity in theregion of nonstenosis as c and the width of the channel in the re-gion of nonstenosis as w, then the flow velocity v(x) is as follows:

vðxÞ= cwyðxÞ

or

vðxÞ= cw

ðr+ bÞ−ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffir2 − ðx− aÞ2

q :

The derivative of thisflow velocity equals the flow elongation rate.Calculating the derivative with Mathematica gives the following:

dvdx

=cwða− xÞffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi

r2 − ða− xÞ2q �

ðr+ bÞ−ffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffir2 − ða− xÞ2

q ��2:

The elongation rates of an 80% stenosed channel were calcu-lated by entering the following parameters into the equation: w =300 μm, r = 300 μm, a = 120√6 μm, b = 60 μm, c = 9,259 μm/s.The highest elongation rate is found at around x = 270 μm. Thiselongation rate was multiplied by a factor of 2 to account for theparabolic nature of the flow in rectangular channels with an aspectratio close to 1, yielding the best estimate of the maximumelongation rate in the channel. Similarly, the maximum elongationrate was estimated in a typical 80% stenosed human carotid artery(parameters: w = 6.5 mm, r = 8 mm, a = 7.5 mm, b = 1.3 mm, c =800 mm/s) and in a 75% indented mouse artery (parameters: w =600 μm, r = 625 μm, a = 600 μm, b = 150 μm, c = 92000 μm/s).Both the computational fluid dynamical simulation and the

calculations of fluid dynamical and geometrical parameters werecarried out with the assumption that the channels and vessels werefilled with Newtonian fluids. This assumption is valid because theshear rates are in a regime in which the non-Newtonian behaviorof blood is limited (3).

1. Kuijpers MJ, et al. (2009) Complementary roles of platelets and coagulation inthrombus formation on plaques acutely ruptured by targeted ultrasound treatment: Anovel intravital model. J Thromb Haemost 7(1):152–161.

2. Gilio K, et al. (2010) Roles of platelet STIM1 and Orai1 in glycoprotein VI- and thrombin-dependent procoagulant activity and thrombus formation. J Biol Chem 285(31):23629–23638.

3. Chien S, et al. (1967) Blood viscosity: influence of erythrocyte aggregation. Science157(3790):829–831.

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Fig. S1. Manufacturing and characterization of microfluidic chips. (A) (Left) A PDMS microfluidic chip with a meandering channel configuration placed ona 24 × 60-mm glass coverslip to create a pattern of immobilized vWF/fibrinogen patches. (Center) After removal of the first PDMS chip, a second chip withstenosis channel configuration is placed over the ligand pattern. (Right) Bright-field and fluorescence images showing immobilized vWF/fibrinogen in a ste-nosis channel. (B) Computational fluid dynamics analysis of wall shear rates in the stenosis channels at input shear rates of 600–2,000 s−1 and stenosis degrees of20–80%. (C) Calculated wall shear rate profile experienced by a platelet passing through this stenotic channel at 2 μm from the channel bottom. (D) Singleplatelet velocity measurements visualized in a color map, demonstrating symmetrical distribution of increased platelet velocity through the stenosis.

Fig. S2. Asymmetric platelet aggregation in microfluidic channels at stenosis outlet regions is inhibited by autocrine receptor antagonists. Representativeconfocal fluorescence images of platelet aggregate formation at a 80% stenosis at 120 s of blood flow in the presence of autocrine receptor antagonists (ARA):cangrelor, 30 μM; MRS2179, 100 μM; indomethacin, 10 μM.

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Fig. S3. Control flow studies in microchannels. (A) Platelet density in stenosis outlet region at input wall shear rate of 1,000 s−1 such in comparison with straightchannels at wall shear rate of 1,000 and 8,000 s−1. (B) Stenosis-dependent platelet aggregation at input wall shear rate of 1,000 s−1 with blood samples anti-coagulated with 40 μM PPACK or 20 μg/mL hirudin. (C) Flow-induced vWF secretion in endothelial cell-seeded microchannels. Confluent HUVECs in stenoticchambers were subjected to 1-min buffer perfusion at input wall shear rate of 1,000 or 8,000 s−1. (Left) Staining for secreted vWF, localized in straight channels.(Right) Quantified staining for vWF at the stenosis outlet region. Note the absence of significant vWF staining at 8,000 s−1 in straight channels. Data are means ± SD;n = 3–4 (***P < 0.05).

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Fig. S4. Direct adhesion and aggregation of platelets in stenotic outlet region. Platelets in whole blood were labeled with DiOC6 and imaged with fastscanning confocal fluorescence microscopy during perfusion through stenotic channels at 1,000 s−1 input wall shear rate. (A) Limited translocation of plateletson vWF/fibrinogen patches. Highlighted is a single platelet translocating over 10 μm in 6.79 s. Note the majority of stationary platelets during this interval. (B)Example of an incidental platelet aggregate dislodged from a ligand patch (white arrow). After embolization (4.39 and 5.59 s), the aggregate did not adhereto ligand patches in the stenosis outlet region.

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Fig. S5. (A) Computational analysis of flow elongational rate for stenosis microchannels. (B) Schematic overview of the parameters used in the calculations (SIMethods). (C) Calculated elongational flow in the 80% stenosis microchannels, a human carotid artery with 80% stenosis, and a mouse carotid artery with 75%indentation.

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Movie S1. Time-lapse fluorescence microscopy of platelet aggregation in a microfluidic stenosis channel with discontinuous vWF/fibrinogen coating. Wholeblood labeled with DiOC6 (2 μg/mL) was perfused through the channel at 1,000 s−1. Shown are conditions with a stenotic geometries of 600-μm diameter anda lumen reduction of 80% and 60%. Note the adhesion of single platelets in patches upstream of the stenosis, and the robust formation of platelet aggregatesin the stenotic outlet region.

Movie S1

Table S1. Comparison of typical geometrical and fluid dynamical parameters among an 80%stenotic microfluidic channel, a human carotid artery with 80% stenosis, and a mouse carotidartery with 75% indentation

Parameter Channel Human carotid Mouse carotid

Upstream width, mm 0.3 6.5 0.6Stenotic width, mm 0.06 1.3 0.15Stenotic length, mm 0.6 15 1.2Upstream mean flow velocity, mm/s 9.3 800* 92*Stenotic mean flow velocity, mm/s 46.5 4,000* 368*Reynolds number, upstream/stenosis 0.4/1.2† 2,600/4,300† 28/44†

Upstream mean wall shear rate, s−1 1,000 820 920Peak mean wall shear rate, s−1 5,000 3,692 3,680Stenosis outlet slope, ° 39 35 37Maximum flow elongation rate, s−1 335 1,100 1,200

In all cases, flow velocity profiles are parabolic, but only mean fluid velocities and the resultant wall shearrates are given. The channels were designed and perfused in a way that main fluid mechanical and geometricalparameters matched those of the in vivo situation (i.e., the angle of the stenosis inlets, the wall shear rates instenotic and nonstenotic areas, and the flow elongation rates). Obvious differences between the systems are theflow rates and the channel dimensions. As a result of these differences, the Reynolds numbers differ consider-able. The lack of turbulence in the in vitro microchannels means that the observed effects are due to changes insimple flow parameters and not obscured by additive turbulent effects.*Systolic values.†Assuming a dynamic viscosity of blood of 2 mPa·s.

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