In vitro system to study realistic pulsatile flow and stretch
signaling in cultured vascular cells
Xinqi
Peng,
Fabio A.
Recchia,
Barry J.
Byrne,
Ilan
S.
Wittstein,
Roy C.
Ziegelstein, and
David A.
Kass
Division of Cardiology, Departments of Medicine and Biomedical
Engineering, Johns Hopkins Medical Institutions, Baltimore,
Maryland 21287-5500
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ABSTRACT |
We developed a novel real-time
servo-controlled perfusion system that exposes endothelial cells grown
in nondistensible or distensible tubes to realistic pulse pressures and
phasic shears at physiological mean pressures. A rate-controlled flow
pump and linear servo-motor are controlled by digital
proportional-integral-derivative feedback that employs
previously digitized aortic pressure waves as a command signal. The
resulting pressure mirrors the recorded waveform and can be digitally
modified to yield any desired mean and pulse pressure amplitude,
typically 0-150 mmHg at shears of 0.5-15 dyn/cm2.
The system accurately reproduces the desired arterial pressure waveform
and cogenerates physiological flow and shears by the interaction of
pressure with the tubing impedance. Rectangular glass capillary tubes
[1-mm inside diameter (ID)] are used for real-time fluorescent
imaging studies (i.e., pHi, NO, Ca2+), whereas
silicon distensible tubes (4-mm ID) are used for more chronic (i.e.,
2-24 h) studies regarding signal transduction and gene
expression. The latter have an elastic modulus of
12.4 · 106 dyn/cm2 similar to in vivo
vessels of this size and are studied with the use of a benchtop system.
The new approach provides the first in vitro application of realistic
mechanical pulsatile forces on vascular cells and should facilitate
studies of phasic shear and distension interaction and pulsatile signal transduction.
shear stress; vessels; nitric oxide; pulse pressure; method
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INTRODUCTION |
VASCULAR ENDOTHELIAL
CELLS are normally exposed to oscillatory distending and shearing
forces due to the pulsatility of circulating blood. This pulsatility
may play an important role in regulating vessel tone (20,
21) and remodeling (4) and is a dominant risk
factor for cardiovascular disease (6, 15, 22). Both shear
and large-amplitude cyclic stretch have been shown to individually stimulate nitric oxide synthase 3 (NOS3) expression and
activity (2, 3), influence cell and F-actin realignment
(12, 14, 23, 24, 26), and modulate various intracellular
signals, such as Ca2+ and pH (9, 10, 25). As
recently suggested by Ziegler et al. (30), however, cells
exposed to both stimuli simultaneously do not necessarily respond as if
both signaling cascades were simply added together. To date, very few
in vitro models have been developed to study the effects of concomitant
pulse perfusion and distension (5, 26), and none have
employed realistic pressure and flow waveforms, but rather sinusoidal
oscillations. This distinction may be important because in vivo studies
have found that the rate of shear development as opposed to mean shear itself can have an important influence on vascular signaling from pathways such as nitric oxide release (7).
To improve in vitro methods for the study of mechanical-vascular
signaling, we developed a novel benchtop system in which endothelial
cells grown inside rigid (glass capillary tubes) or distensible
(silastic) tubes could be exposed to physiological pressure/flow
waveforms generated by a real-time computer-controlled digital feedback
servo-pump system. The apparatus enables control over strain and shear
stress and their combination. Studies can be performed with the use of
fluorescent microscopy to examine real-time signaling or, over several
days, to evaluate molecular expression changes. Here we
describe both systems and their performance characteristics and
demonstrate their utility for studying short and longer-term molecular
and cellular signaling in endothelial cells.
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METHODS |
Servo-perfusion apparatus.
The perfusion system is schematically shown in Fig.
1. Specific details (tubing
lengths, dimensions, etc.) are provided in APPENDIXES A AND B. A sealed reservoir flask (A) contains
physiological buffer solution (i.e., HEPES) or cell culture medium
(e.g., DMEM + 10% FCS without antibiotics). Both are warmed and
bubbled with 95%O2-5%CO2 to yield a pH of 7.3 at 37°C. Medium is then withdrawn from the reservoir by a
rate-controlled flow pump (B; Ismatec 7617-60 for
capillary tube studies; Masterflex-7523-20 with 7519-06 or
7519-10 pump heads for the larger distensible tube studies). The
fluid passes through a compliance chamber (C) to reduce
pump-generated pulsations and/or enhance servo-control stability. For
distensible tube studies, a larger compliance is required that must,
therefore, be isolated from the downstream servo-pump by an in-line
hydraulic resistor and one-way valve (D). Fluid then passes
into a chamber with a movable base (E). For capillary tube
studies, this chamber is a 20-ml glass syringe with the plunger affixed
to a linear motor (517-9-3.6; Applied Engineering, San Jose, CA).
For Silastic tube studies, the chamber is a 30-ml plastic inverted
Plexiglas cone, and the chamber base is a silastic-coated flexible
diaphragm (4-200-81-CB900; Bellofram, Newell, WV) linked to a
linear servo-motor (F; no. 411, Ling Apparatus, CT; and
model 4020-LS 630D1064 REV F, Aerotech, PA). For both systems, the
linear motor is controlled by a P5-60 MHz computer (G)
that runs real-time custom digital software. A micromanometer catheter (H; SPC-350; Millar Instruments) measures luminal
pressure and provides the feedback signal for the servo-pump to
generate a desired mean and pulse pressure. Pulsatile perfusion is then directed into nondistensible tubing of a sufficient length to establish
laminar flow and is then passed through the tubes containing cultured
cells (J). An in-line ultrasound flowmeter (I;
T101, 2N or 1N probe; Transonic) records phasic and mean flow.
Cells can be concomitantly imaged by fluorescent or light microscopy. After exiting the tubes, perfusate passes through a distal terminal hydraulic resistor (K) to set the mean pressure. Effluent is
then either discarded (when studied at low flow rates) or recirculated through the reservoir (A; for studies at higher flow rates).
In chronic experiments conducted at low flow rates, an in-line heat exchanger is interposed between D and G to help
maintain constant temperature.

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Fig. 1.
Schematic diagram of the pulse-perfusion system.
Perfusate is withdrawn from a reservoir (A) at a flow rate
regulated by a pump (B) under computer feedback control.
Flow then passes through a compliance chamber (C) to dampen
pump-generated oscillations and/or stabilize the feedback system. In
the application in which the flow pump is pulsatile and a larger
compliance required, flow then passes through an in-line hydraulic
resistor and one-wave valve (D) to help isolate the
compliance from the downstream servo-pump. Perfusate then enters a
chamber (E) with a movable floor linked to a linear motor
(F). The motor is controlled by digital feedback
(G) control. Pressure is measured by micromanometer
(H) near the exit port of the linear motor, and this signal
is used to generate the servo-command. Pulsatile flow then passes
through tubing that matches the diameter of that which contains
cultured cells and is of an appropriate length to assure laminar flow,
passes through an in-line flow probe (I) and then
cell-seeded tubes (J) that can be imaged in real-time with
the use of light (silicon tubes) and/or fluorescent (glass tubes)
microscopy. After it exits these tubes, flow passes through a resistor
(K) to establish mean pressure and is either discarded or
recirculated through the reservoir (A).
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The pump system is regulated by custom-developed
proportional-derivative-integral feedback control software (available
on request). The flow pump is controlled by both a proportional (open loop) gain and an integral gain to maintain a given flow rate. A second
proportional open loop and the derivative gain controls the linear
motor. Digital sampling and feedback calculations are performed at a
frequency of 1 kHz. To establish a mean flow rate, the servo-program is
first set at the desired mean perfusion pressure (i.e., 90 mmHg), and
the distal hydraulic resistor is adjusted until mean flow reaches the
desired level (0.5-275 ml/min, corresponding to mean shears that
vary with tubing size, ranging from 0.3 to 15 dyn/cm2) at
this mean pressure. Mean pressure is constant due to integral feedback
regulation. Pulse pressure is generated by digitally varying the
amplitude and/or waveform shape of the command signal and then
generating the identical pressure by real-time feedback.
Distensible tubing.
To study concomitant stretch and shear signaling, distensible tubes
were custom designed in collaboration with Specialty Manufacturing, Saginaw, MI. They are composed of dimethyl silicone applied over a
bovine gelatin-coated glass or highly polished (electropolished) stainless steel mandrels (radius/thickness ratio = 8.34, i.e., for
2.0-mm radius, thickness is 0.24 mm, with 18.5-cm-length). Manufacturing tolerance for wall thickness is within 0.05 mm, which can
result in some modest variability in tubing elasticity. This can be
calibrated by assessing pulse pressure-strain dependence for each lot
before study. The tubing internal diameter and surface area do not vary
(a 4-mm-diameter tube is 23.2 cm2 and yields 150-200
µg of total protein from cell lysate).
The tubes currently in use have an incremental elastic modulus of
12.4 · 106 dyn/cm2, commensurate with
in vivo vessels of their size (17). The pressure strain
elastic modulus (Ddia ·
P/
D) where
Ddia is end-diastolic diameter is
1.4 · 106 dyn/cm2, which is also
comparable to in vivo values (1). Thus pulse pressures of
0-150 mmHg translate to 0-16% strain, similar to strain
levels obtained in distensible membrane systems. Unlike membrane
systems, however, pulse strain is uniformly radially applied to the
side wall and thus cells lining the tube. Strains of ~4-6%
(i.e., PP: 40-90 mmHg) are typical in vivo (1,
18). Tubing distension is assessed by real-time video-edge
detection of the outer diameter (Crescent Electronics). For
luminal shear (
) calculations, we employed Womersely's modified
formula (5) for pulsatile flow
=
/
× µQ/
r3; where
= r(2

/µ)1/2, r is vessel radius in cm, µ is
viscosity in poise,
is fluid density in gm/ml,
is pulsation
frequency (s
1), and Q is mean flow in ml/s. For steady
shear,
= µQ/
r3.
Flow instability and ultimately turbulence can be expected at a mean
flow rate of
400 ml/min (Reynolds number > 1,500). This flow
yields an estimated mean shear of 14 dyn/cm2, but this
would be high for a 4-mm in vivo vessel. Nonturbulent higher shears can
be obtained by reducing the tube diameter. However, we found the thin
wall required to maintain distensibility in 2- to 3-mm tubes often led
to aneurysm formation when tubes were exposed to pulse perfusion. Thus
4 mm represents the current practical size limit for reliable
mechanical performance. For pulsatile experiments, nonturbulent flow is
easily achieved at 245 ml/min (mean shear 7 dyn/cm2),
commensurate with values for conductance vessels.
Tube preparation and cell seeding.
Silastic tubes are covered with plastic caps over which a rubber gasket
(Quest Medical) is placed and then autoclaved. A solution of 0.01%
fibronectin (Sigma) is introduced into the tube, and tubes rotate for
1-2 h at 10 rpm at room temperature with the use of a custom
rotisserie apparatus. At this point, fluid is withdrawn and replaced by
medium for 1 h before cell seeding. Bovine aortic endothelial
cells (BAEC; passages 4-8; Coriell Cell Repositories) are cultured to confluence in standard dishes
using DMEM supplemented with 10% fetal bovine serum and 2 mm
glutamine. Tubes are filled with
2-4 × 105
cells/ml. Tubes are incubated at slow rotation of 10 rpm to
achieve uniform seeding. Medium is changed every other day, and
confluent monolayers are achieved after 1-2 days. Tubes are
mounted on the perfusion system (Fig. 1). To assure sterility, the
system is designed so tubes can be interposed with perfusion tubes
within a sterile hood, and the entire system is mounted onto the
flowmeter and pulse motor without exposure to room air. Glass tubes are seeded as previously described (27).
Application protocols: real-time signaling studies.
The glass capillary tubes have a cell surface area <1% of a standard
culture dish and are therefore of limited use for molecular signaling
analysis. However, they can be used for real-time fluorescent imaging
to study signaling responses. One example of this application is for
the study of endothelial pHi. BAECs are grown to confluence within glass capillary tubes and then loaded over 30 min with 10 M of
esterified cSNARF 1-AM, a fluorescent H+-sensitive
indicator (28). The cSNARF-1-loaded cells are washed with
buffer, mounted on the stage of an inverted epifluorescence microscope
(Nikon Diaphot-3), and excited with a Xenon short-arc lamp (UXL-75 XE;
Ushio) at 530 ± 5 nm. The emission ratio of 590 ± 5/640 ± 5 nm measures pHi (27). Data are
obtained under nonpulsatile conditions at a constant flow of 0.5 ml/min
and 90 mmHg mean perfusion pressure. Flow is then increased to a
constant value of 6 ml/min (12 dyn/cm2) or to a pulsatile
flow (peak 20 dyn/cm2) with the use of a pulse-pressure
waveform of 75 mmHg. cSNARF-1 fluorescence is monitored continuously to
yield real-time pHi change in response to each stimuli.
Distensible tubing studies: molecular signaling.
The distensible Silastic tubes provide a sufficiently large surface
area for molecular signaling analysis and can also be employed in
chronic (2-48 h) exposure studies. An example of this application
is a 24-h examination of NOS3 expression in response to varying levels
of flow pulsatility. Tubes are subjected to two levels of shear (0.16 and 1.6 dyn/cm2) with pulse pressure ranging from 0 to 150 mmHg (maximal strain of 16%). After 24 h, the majority of the
tube is used for obtaining cell lysate. A 3-cm portion is cut
longitudinally, flattened, and mounted on a glass slide, fixed in 3.7%
formaldehyde/PBS, and stained with either 0.1% crystal violet for cell
morphology analysis or FITC-conjugated rhodamine phalloidin (Molecular
Probes) for actin filament analysis. Microscopic examination is made on the open cut surface.
For these studies, total cell protein was analyzed by immunoblot to
assess NOS3 expression. Lysis buffer contained (in mM) 150 NaCl, 50 Tris · HCl (pH 7.6), 1 EDTA, 0.1% SDS, 1% sodium deoxycholate, 1% Triton X-100, 1 phenylmethylsulfonyl fluoride, 50 NaF, 0.5 sodium orthovanadate, 1 µg/ml pepstatin A, 1 µg/ml leupeptin, and 2 µg/ml aprotinin. Tubes were manually squeezed, and
contents were centrifuged at 12,000 rpm for 20 min at 4°C. Protein
quantified by the Bradford assay (Bio-Rad Laboratories, Hercules, CA)
was loaded (10 µg/lane) into 7.5% SDS-polyacrylamide minigels.
Coomassie blue staining confirmed equal loading. Blots were incubated
with 1:1,000 monoclonal anti-NOS3 antibody (Transduction Laboratories)
and 1:4,000 monoclonal
-actin antibody (Sigma, A4700) at room
temperature. After being washed, goat anti-mouse secondary antibody
conjugated to horseradish peroxidase was applied, NOS3 and
-actin
immunoreactivity was detected by chemiluminescence (Amersham), and band
density was quantified (Adobe Photoshop). Data were derived from
four separate tubes for each flow/pulsatility combination, with gels
run in duplicate and results averaged. To compensate for interrun
variability, nonpulsatile flow lanes were run in each gel, and
expression at various pulse pressures was normalized to this control.
Data were analyzed by two-way ANOVA, with pulse pressure and mean shear
as the grouping variables.
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RESULTS |
Mechanical performance of perfusion system.
Figure 2 demonstrates the performance
features of the servo-system used with distensible tubing. The system
that employs glass capillary tubes has very similar characteristics.
The frequency response (A) was flat to at least 15 Hz, more
than twice that which is needed to accurately reproduce an arterial
pressure waveform at 60 cycles/s (8). Figure 2B
displays the input command (left), the generated pressure
waveform (middle), and a plot of one vs. the other (i.e.,
input-output, right). The highly accurate waveform reproduction with minimal deviation of amplitude or phase is
demonstrated (right).

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Fig. 2.
A: frequency response of servo-perfusion
system. The input was a sine wave at constant amplitude and 0.5-15
Hz frequency. Output pressures generated by the servo-system (at a mean
flow of 200 ml/min) are shown, revealing a flat frequency response.
B: example of pressure input command signal
(left), the real-time generated pressure waveform
(middle), and plot of input vs. output (right).
The solid line (right) is the line of identity.
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Figure 3 provides typical pressure and
flow signals studied with both perfusion systems. Data are shown using
1 Hz command signals. Figure 3A displays luminal pulse
pressure, flow, and vessel diameter tracings for perfusion combinations
in 5-mm-ID distensible tubes. The servo-generated arterial pressure
waveforms mirrored the prerecorded signals (A), and
corresponding phasic flows were determined by the interaction with the
downstream impedance. The latter waveforms are similar to in vivo flows
for moderate-sized arteries (17). Vessel diameter changes
mirror the pressure waves. At 100 ml/min, peak shear at the highest
pulse pressure was 2.5 dyn/cm2. Four-millimeter-ID tubes at
275 ml/min flow yield peak shears of 13.4 dyn/cm2 (not
shown). At lower flows, i.e., 10 ml/min (right), flow
reversal is observed as pulsatility increases. Pressure and vessel
distension signals are identical to those at higher flow, but shear
varies from +1.02 to
0.61 dyn/cm2. Such flow patterns are
common when enhanced pulsatility occurs at low shears in vivo. However,
this differs from signaling generated in the more widely used
closed-chamber distensible systems where concomitant shear
fluctuation is minimal.

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Fig. 3.
A: hemodynamics of distensible tubes pulse perfused by
computer-servo-controlled system for 2 different levels of mean shear.
Increasing pulsatility generated identical pressures and vessel
distension waveforms at both levels of flow, whereas pulsatile flow
(and shear) was very different, with shear reversal in the low-flow
state. B: hemodynamics of flow and pressure profiles
generated in rigid capillary tube system. The pressure is
servo-controlled, and the realistic flow waveform is the product of its
interaction with the perfusion tubing.
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Figure 3B displays signals generated in the capillary tubing
system. The pressure again closely duplicates the stored command signal, whereas the physiological flow waveform results from an interaction of this pressure with the downstream impedance of the
tubing. Data are shown for mean shear of 0.83 dyn/cm2 but
are easily obtained over a broad range of flow rates, encompassing the
full range of physiological shears. Pulsatile perfusion was generated
with the use of a 75-mmHg pulse pressure, producing phasic shear with a
peak of 20 dyn/cm2.
Applications: NOS3 expression and cellular pHi.
Figure 4A (top)
shows NOS3 immunoblots for cells exposed to 1.6 or 0.16 dyn/cm2 mean shear at varying levels of pulsatility. At low
nonpulsatile shear, NOS3 expression was unchanged compared with
incubator control, and this persisted with modest increases in
pulsatility (PP = 40 mmHg). Expression rose by 36.8 ± 6.9%
(P = 0.01) at higher pulsatility (90 mmHg, 9% strain)
but declined back to baseline levels at higher strains
(P = 0.009 16% vs. 9% strain) as flow reversal
increased (cf. Fig. 3). At 10-fold higher nonpulsatile flow, NOS3
expression rose +23.1 ± 12% (P < 0.001, n = 8) compared with control. Under these conditions,
increasing pulsatility induced significant further rises in
NOS3 expression (i.e., +23.6 ± 11.1% at PP = 40 vs. 0, P < 0.01), but this was little changed by higher levels of pulsatility. Summary data are shown in Fig. 4A
(bottom). At even higher nonpulsatile shear rates, there was
less additive effect from pulsatility (data not shown) consistent with
recent studies (30).

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Fig. 4.
A: immunoblots data for nitric oxide synthase
3 (NOS3, 140 kDa) in tubes exposed to lower and higher mean shear.
Compared with preperfusion control (C), laminar flow [pulse pressure
(PP) = 0] increased NOS3 expression at the higher but not lower flow
rate. With increasing perfusion pulsatility, there was a further rise
in expression at all pulse pressures with higher flow, but only a rise
in expression at 90 mmHg at reduced flow. A, bottom,
provides summary data (mean of 4 tubes per bar) comparing the percent
change in NOS3 expression at each pulse pressure with a PP = 0 reference. B: example of pHi responses to
increasing perfusion pulse pressure followed by either steady or
pulsatile flow. The pHi declines with steady flow increases
but rises with pulsatile reversing flow.
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Figure 4B displays example data from an experiment
evaluating pHi. Exposure of cells to an increase in laminar
steady flow from 0.5 to 6 ml/min at a constant pressure of 90 mmHg
results in acidification as previously reported (27).
However, if mean flow remains constant but pulsatility is added, an
alkalization results. These data support a specific role of phasic
shear stress to endothelial signaling pathways (25).
Cell morphology and F-actin alignment in distensible tubing exposed
to pulsatile perfusion.
Figure 5 displays photomicrographs of
cells perfused in distensible tubes for 24 h at varying levels of
flow pulsatility (all at a mean shear of 1.6 dyn/cm2).
Confluent randomly oriented cells were observed with nonpulsatile or
low-pulsatility flow (Fig. 5, A and B), whereas
cells become aligned in the flow direction at higher pulsatility (Fig.
5, C and D). Figure
6 displays histograms of cell angular
orientation (derived from 4 separate tubes for each condition) relative
to the horizontal flow direction (0°). At higher pulsatility, the orientation clustered parallel to flow, with a narrow angle
distribution. The SD of each distribution significantly declined with
increasing PP (P < 0.0001) from 63.4 ± 4.2 with
nonpulsatile flow to 11.7 ± 1.2 at PPs of 90 or 150 mmHg.

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Fig. 5.
Influence of varying perfusion pulsatility on cell elongation and
alignment. PP: A, 0; B, 40; C, 90; and
D, 150 mmHg. Flow direction is horizontal on each panel.
Each field is 0.8 × 1.4 mm. Cell confluence was observed under
all conditions. With increasing PP (C and D),
cells became elongated and aligned in the flow direction.
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Fig. 6.
Histograms of cell orientation angle relative to the flow direction
(0°). These angles were randomly distributed at 0 (A) and
40 (B) mmHg pulse pressure, but clustered more closely at
0° with higher pulsatility [90 (C) or 150 (D)
PP].
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Figure 7 displays higher power images of
cellular actin fiber realignment, revealing translocation of fibers
from a diffuse cytosolic distribution with nonpulsatile flow
(7A) to peripheral longitudinal orientation with increased
pulsatility (7B), consistent with previous studies
(26).

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Fig. 7.
Influence of increased pulsatility on actin filament alignment.
A: example of cells exposed to 1.5 dyn/cm2 mean
shear with PP = 0, displaying extensive cross-cellular orientation
of actin filaments. B: after exposure to PP = 90 mmHg
(6% strain), cells demonstrate peripherally oriented actin filaments
with central clearing associated with cell elongation. Dimensions are
80 × 140 µm.
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DISCUSSION |
We present two novel servo-controlled in vitro perfusion systems
in which endothelial cells can be exposed to physiological realistic
pulse pressures and flows to study signal transduction interactions.
The major new aspects of this methodology over prior models are the
implementation of real-time servo-systems to generate realistic
pressure/flow waves, adaptation of the pump systems to full benchtop
operation that employs heat exchangers if needed, and the development
of custom-designed Silastic tubing (now commercially available) that
facilitates cell adhesion to the inner surface despite marked
elevations in perfusion pulsatility. Previously reported systems
combining pulse distension and flow (26, 30) employed a
sinusoidal waveform to generate pulsatility. This distinction may be
important, because the physiological waveform yields nearly an order of
magnitude greater rate of strain and shear change when contrasted with
a sine wave of similar amplitude. This may further contribute to NOS3
(7) or other cell signaling.
Our system reproduces cellular alignment and actin cytoskeletal
rearrangement that have been previously reported in response to
pulsatile stretch, flow, and combined signals (26). We
demonstrate by example how the pattern of simultaneous flow can
influence net signaling from pulsatile stretch on proteins such as
NOS3, or how cyclic reversing shear modulates pHi
differently from constant shear. Data presented for both applications
(19, 25) are the focus of ongoing investigations.
Our goal was to develop a system whereby realistic (physiological) flow
and pressure waveforms could be imposed on vascular cells. The
particular applications described with both versions of this method
focused solely on endothelial cell signaling. This has limitations,
because coculture studies have demonstrated that interactions between
vascular smooth muscle and the endothelium can alter the nature of
cellular signaling to both agonist and mechanical stimulation
(11, 16, 29). Furthermore, cultured endothelial cells
often undergo phenotypic changes that can influence mechanosignaling
responses, and our method does not overcome this limitation. However,
the method can be adapted to ex vivo vessel analysis. There is only one
previous study employing pulse perfusion in isolated ex vivo vessels
(13). This system employed classical linear, small-signal
theory as opposed to rapid-response real-time PID feedback
that is used in the present study. The result would appear to be
improved realization of the precise pressure waveform and a
physiological flow waveform with few high-frequency oscillations (i.e., Fig. 3) with the present system. Furthermore, if distensible tubes are replaced by a sterile chamber filled with filtered
humidified air, a vascular segment could be interposed and perfused ex
vivo with the same control over flow and pulsatility. Coculture is also
feasible in Silastic tubes and could be used to study interactions of
both cell types.
Increasing evidence suggests a critical role of flow pulsatility
on morphological signaling of the vascular wall, on
endothelial-mediated regulation of smooth muscle tone, and on platelet
adhesion. The current method enables these signaling mechanisms to be
accurately studied in vitro employing realistic mechanical stimuli and
should thereby provide a valuable addition of available methods for
these investigations.
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APPENDIX A |
Specifications for Pulse-Perfusion System Used With
Distensible Silastic Tubing
All perfusion tubing is Phar-Med7 (Masterflex,
Cole-Parmer). From reservoir (Fig. 1A) to flow pump (Fig.
1B), L/S-18 tubing (40 cm) connects via
y-connector(s) (PGC Scientifics, 82-2520-24) to two or four
L/S-17 (35 cm), mounted within the flow pump. The tubing network
reconverges into a single 9-cm L/S-18 tube that leads to the compliance
chamber (Fig. 1C). This chamber is in parallel with the flow
line (8.5-cm diameter × 11-cm height) and filled with 9 cm
air above the fluid level. Flow then traverses a 12-cm-long L/S-17 tube
to the pulse-generating chamber (Fig. 1F), passing through a
screw-clamp hydraulic resistor and unidirectional flow valve (St. Jude
aortic prosthetic valve, 1 cm) to isolate the servo-pump from upstream
compliance. The pulse generator (Fig. 1F) is cone shaped
with a 5.1-cm open base and 3.2-cm height. The base is sealed by a
4-cm-diameter flexible top-hat style silicon-coated diaphragm
linked via a plastic cylinder (4-cm base, 3-cm height) to the linear
servo-motor. The outport at the tip of the cone couples to L/S-18
tubing, with a side port for micromanometer pressure recording. Flow
continues through L/S-18 tubing (22 cm) that bifurcates (Cole-Parmer,
06295-30) into two L/STM-25 (Cole-Parmer P-06485-25,
20-30 cm) tubes that match the impedance of the cell-seeded distensible tubes to assure laminar flow. Medium exits via parallel 15-cm L/STM-25 tubes, one fitted with an in-line flow probe (Transonic, 2N), which then converge via y-adaptors to a single 20-cm return tube
(L/S-17) recirculating medium to the reservoir. A screw-type hydraulic
resistor is placed just prior to reentry into the reservoir and
adjusted to set mean pressure for any flow level.
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APPENDIX B |
Specifications for Pulse-Perfusion System Used With Glass
Capillary Tubes
As with the system designed for use with larger distensible
tubing, almost all perfusion tubing is Phar-Med (MasterFlex,
Cole-Parmer). Perfusate is withdrawn into L/S-16 tubing (45-cm length)
and flows through a t-connector (Cole-Parmer, P-06478-40, 3.16-mm
ID). The trunk of the t-connector is linked to the compliance chamber, which in this setup is also in parallel (rather than in series) with
the servo-motor. This compliance chamber is 10 ml, filled to all except
1-2 ml by perfusate. The remaining port of the t-connector is
linked to a P-06365-66 (Cole-Parmer) connector to transit to L/S-14 tubing (12-cm) fitted with an in-line flow probe (Transonic, 1N). Flow enters another 12-cm-length of L/S-14 tubing and then into
L/S-13 tubing (5.5-cm) via another reducing connector (P-06365-55) with an internal diameter of 0.8 mm. This is linked to the glass capillary tube seeded with endothelial cells. The exit port of the
glass tube is again linked to L/S-13 tubing (15-cm) and then to larger
L/S-16 tubing (silastic, 60-cm) to provide a distal compliance. The
terminal tube is linked via a multiple stopcock manifold terminating
with rotating diaphragm resistors (Tuhey-Borst adaptors) that enable
resistance to be varied to achieve a desired mean pressure for a given flow.
 |
ACKNOWLEDGEMENTS |
This study was supported by National Heart, Lung, and Blood
Institute Grant HL-47511 (to D. A. Kass) and an American Heart Association-Maryland Chapter fellowship grant (to F. Recchia).
 |
FOOTNOTES |
Present address of F. A. Recchia: New York Medical College, Dept.
of Physiology, BSB636, Valhalla, NY 10595.
Present address of B. J. Byrne: Pediatric Cardiology, Univ. of
Florida, PO Box 100296, Gainesville, FL 32610.
Address for reprint requests and other correspondence: D. A. Kass, Halsted 500, Johns Hopkins Medical Institutions, 600 N. Wolfe
St., Baltimore, MD 21287 (E-mail: dkass{at}bme.jhu.edu).
The costs of publication of this
article were defrayed in part by the
payment of page charges. The article
must therefore be hereby marked
"advertisement"
in accordance with 18 U.S.C. §1734 solely to indicate this fact.
Received 10 February 2000; accepted in final form 3 April 2000.
 |
REFERENCES |
1.
Armentano, RL,
Barra JG,
Levenson J,
Simon A,
and
Pichel RH.
Arterial wall mechanics in conscious dogs.
Circ Res
76:
468-478,
1995[Abstract/Free Full Text].
2.
Awolesi, MA,
Sessa WC,
and
Sumpio BE.
Cyclic strain upregulates nitric oxide synthase in cultured bovine aortic endothelial cells.
J Clin Invest
96:
1449-1454,
1995[ISI][Medline].
3.
Awolesi, MA,
Widmann MD,
Sessa WC,
and
Sumpio BE.
Cyclic strain increases endothelial nitric oxide synthase activity.
Surgery
116:
439-444,
1994[ISI][Medline].
4.
Baumbach, GL.
Is pulse pressure a stimulus for altered vascular structure in chronic hypertension?
Hypertension
18:
728-729,
1991[ISI][Medline].
5.
Benbrahim, A,
L'Italien GJ,
Milinazzo BB,
Warnock DF,
Dhara S,
Gertler JP,
Orkin RW,
and
Abbott WM.
A compliant tubular device to study the influences of wall strain and fluid shear stress on cells of the vascular wall.
J Vasc Surg
20:
184-194,
1994[ISI][Medline].
6.
Benetos, A,
Safar M,
Rudnichi A,
Smulyan H,
Ducimetieere P,
and
Guize L.
Pulse pressure: a predictor of long-term cardiovascular mortality in a French male population.
Hypertension
30:
1410-1415,
1997[Abstract/Free Full Text].
7.
Canty, JMJ,
and
Schwartz JS.
Nitric oxide mediates flow-dependent epicardial coronary vasodilation to changes in pulse frequency but not mean flow in conscious dogs.
Circulation
89:
375-384,
1994[Abstract].
8.
Chen, CH,
Nevo E,
Fetics B,
Pak PH,
Yin FC,
Maughan WL,
and
Kass DA.
Estimation of central aortic pressure waveform by mathematical transformation of radial tonometry pressure. Validation of generalized transfer function.
Circulation
95:
1827-1836,
1997[Abstract/Free Full Text].
9.
Helmlinger, G,
Berk BC,
and
Nerem RM.
Calcium responses of endothelial cell monolayers subjected to pulsatile and steady laminar flow differ.
Am J Physiol Cell Physiol
269:
C367-C375,
1995[Abstract/Free Full Text].
10.
Helmlinger, G,
Berk BC,
and
Nerem RM.
Pulsatile and steady flow-induced calcium oscillations in single cultured endothelial cells.
J Vasc Res
33:
360-369,
1996[ISI][Medline].
11.
Hendrickson, RJ,
Cappadona C,
Yankah EN,
Sitzmann JV,
and
Redmond EM.
Sustained pulsatile flow regulates endothelial nitric oxide synthase and cyclooxygenase expression in co-cultured vascular endothelial and smooth muscle cells.
J Mol Cell Cardiol
31:
619-629,
1999[ISI][Medline].
12.
Kanai, AJ,
Strauss HC,
Truskey GA,
Crews AL,
Grunfeld S,
and
Malinski T.
Shear stress induces ATP-independent transient nitric oxide release from vascular endothelial cells, measured directly with a porphyrinic microsensor.
Circ Res
77:
284-293,
1995[Abstract/Free Full Text].
13.
Labadie, RF,
Antaki JF,
Williams JL,
Katyal S,
Ligush J,
Watkins SC,
Pham SM,
and
Borovetz HS.
Pulsatile perfusion system for ex vivo investigation of biochemical pathways in intact vascular tissue.
Am J Physiol Heart Circ Physiol
270:
H760-H768,
1996[Abstract/Free Full Text].
14.
Malek, AM,
and
Izumo S.
Mechanism of endothelial cell shape change and cytoskeletal remodeling in response to fluid shear stress.
J Cell Sci
109:
713-726,
1996[Abstract/Free Full Text].
15.
Mitchell, GF,
Moye LA,
Braunwald E,
Rouleau JL,
Bernstein V,
Geltman EM,
Flaker GC,
and
Pfeffer MA.
Sphygmomanometrically determined pulse pressure is a powerful independent predictor of recurrent events after myocardial infarction in patients with impaired left ventricular function. SAVE investigators. Survival and Ventricular Enlargement.
Circulation
96:
4254-4260,
1997[Abstract/Free Full Text].
16.
Nackman, GB,
Fillinger MF,
Shafritz R,
Wei T,
and
Graham AM.
Flow modulates endothelial regulation of smooth muscle cell proliferation: a new model.
Surgery
124:
353-360,
1998[ISI][Medline].
17.
Nichols, WW,
and
O'Rourke MF.
Contours of pressure and flow waves in arteries.
In: McDonald's Blood Flow in Arteries, edited by Nichols WW,
and O'Rourke MF.. London: Arnold, 1998, p. 170-200.
18.
Nichols, WW,
and
O'Rourke MF.
Properties of the arterial wall: practice.
In: McDonald's Blood Flow in Arteries, edited by Nichols WW,
and O'Rourke MF.. London: Arnold, 1998, p. 73-97.
19.
Peng, X,
Shu W,
and
Kass DA.
Interaction of enhanced pulse pressure and pulse-flow on NOS3 expression in endothelial cells cultured inside distensible tubes (Abstract).
Circulation
98, suppl 1:
I3509,
1998.
20.
Recchia, FA,
Senzaki H,
Saeki A,
Byrne BJ,
and
Kass DA.
Pulse pressure-related changes in coronary flow in vivo are modulated by nitric oxide and adenosine.
Circ Res
79:
849-856,
1996[Abstract/Free Full Text].
21.
Saeki, A,
Recchia F,
and
Kass DA.
Systolic flow augmentation in hearts ejecting into a model of stiff aging vasculature. Influence on myocardial perfusion-demand balance.
Circ Res
76:
132-141,
1995[Abstract/Free Full Text].
22.
Safar, ME,
Siche JP,
Mallion JM,
and
London GM.
Arterial mechanics predict cardiovascular risk in hypertension.
J Hypertens
15:
1605-1611,
1997[ISI][Medline].
23.
Sumpio, BE,
Banes AJ,
Link GW,
and
Iba T.
Modulation of endothelial cell phenotype by cyclic stretch: inhibition of collagen production.
J Surg Res
48:
415-420,
1990[ISI][Medline].
24.
Uematsu, M,
Ohara Y,
Navas JP,
Nishida K,
Murphy TJ,
Alexander RW,
Nerem RM,
and
Harrison DG.
Regulation of endothelial cell nitric oxide synthase mRNA expression by shear stress.
Am J Physiol Cell Physiol
269:
C1371-C1378,
1995[Abstract/Free Full Text].
25.
Wittstein, IS,
Qiu W,
Ziegelstein RC,
Hu Q,
and
Kass DA.
Opposite effects of pressurized steady versus pulsatile perfusion on vascular endothelial cell cytosolic pH: role of tyrosine kinase and mitogen-activated protein kinase signaling.
Circ Res
86:
1230-1236,
2000[Abstract/Free Full Text].
26.
Zhao, S,
Suciu A,
Ziegler T,
Moore JE, Jr,
Burki E,
Meister J-J,
and
Brunner HR.
Synergistic effects of fluid shear stress and cyclic circumferential stretch on vascular endothelial cell morphology and cytoskeleton.
Arterioscler Thromb Vasc Biol
15:
1781-1786,
1995[Abstract/Free Full Text].
27.
Ziegelstein, RC,
Cheng L,
and
Capogrossi MC.
Flow-dependent cytosolic acidification of vascular endothelial cells.
Science
258:
656-659,
1992[ISI][Medline].
28.
Ziegelstein, RC,
Corda S,
Pili R,
Passaniti A,
Lefer D,
Zweier JL,
Fraticelli A,
and
Capogrossi MC.
Initial contact and subsequent adhesion of human neutrophils or monocytes to human aortic endothelial cells releases an endothelial intracellular calcium store.
Circulation
90:
1899-1907,
1994[Abstract].
29.
Ziegler, T,
Alexander RW,
and
Nerem RM.
An endothelial cell-smooth muscle cell co-culture model for use in the investigation of flow effects on vascular biology.
Ann Biomed Eng
23:
216-225,
1995[ISI][Medline].
30.
Ziegler, T,
Bouzourene K,
Harrison VJ,
Brunner HR,
and
Hayoz D.
Influence of oscillatory and unidirectional flow environments on the expression of endothelin and nitric oxide synthase in cultured endothelial cells.
Arterioscler Thromb Vasc Biol
18:
686-692,
1998[Abstract/Free Full Text].
Am J Physiol Cell Physiol 279(3):C797-C805
0363-6143/00 $5.00
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