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Am J Physiol Cell Physiol 279: C797-C805, 2000;
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Vol. 279, Issue 3, C797-C805, September 2000

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


    ABSTRACT
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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


    INTRODUCTION
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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.


    METHODS
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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).

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 · Delta P/Delta 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 (tau ) calculations, we employed Womersely's modified formula (5) for pulsatile flow tau  = alpha /<RAD><RCD>2</RCD></RAD> × µQ/pi r3; where alpha  = r(2pi rho upsilon /µ)1/2, r is vessel radius in cm, µ is viscosity in poise, rho  is fluid density in gm/ml, upsilon  is pulsation frequency (s-1), and Q is mean flow in ml/s. For steady shear, tau  = µQ/pi 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 approx 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 beta -actin antibody (Sigma, A4700) at room temperature. After being washed, goat anti-mouse secondary antibody conjugated to horseradish peroxidase was applied, NOS3 and beta -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.


    RESULTS
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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.

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.

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.

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].

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.


    DISCUSSION
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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.


    APPENDIX A
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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.


    APPENDIX B
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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
TOP
ABSTRACT
INTRODUCTION
METHODS
RESULTS
DISCUSSION
APPENDIX A
APPENDIX B
REFERENCES

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Am J Physiol Cell Physiol 279(3):C797-C805
0363-6143/00 $5.00 Copyright © 2000 the American Physiological Society


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