Development of a Novel Vascular Simulator and Injury Model to Evaluate Smooth Muscle Cell Response following Balloon Angioplasty

Article Outline

• Abstract
• Introduction
• Materials and Methods
  • Simulator Design
  • Cell Culture
  • Cell Seeding for Mechanical Testing
  • Simulation Experiment Groups
  • Mechanical Testing
  • Evaluation of Cellular Response
    • Cell proliferation
  • Phenotypic Marker
  • Statistical Analysis
• Results
  • Proliferation
  • Western Blot
• Discussion
• Conclusion
• Acknowledgment
• References
• Copyright

Following balloon angioplasty, denudation of endothelial cells exposes vascular smooth muscle cells (SMCs) to normally unseen shear forces from blood flow. In vivo studies investigate the response to angioplasty injury, but limited studies have been performed using in vitro systems. In order to study SMC response in vitro, a concurrent shear and tensile forces simulator has been developed to provide clinically significant levels of strain and shear stresses in addition to simulating forces similar to those during balloon angioplasty. In this acute study (8 hr), rat aortic SMCs demonstrated significant cell proliferation following applied increased tensile forces of angioplasty injury and shear exposure when compared to lower levels of tensile exposure similar to a normal physiological level, with an average 75% increase in the number of cells of the injury group compared to the normal dynamic group. SMCs exposed to balloon angioplasty injury and concurrent shear and tensile mechanical forces demonstrated decreased expression of the contractile phenotypic marker smooth muscle α-actin. These findings demonstrate the efficacy of the developed model for in vitro angioplasty and the simulated mechanical environment to the cells. This provides an in vitro model to isolate the effects of concurrent mechanical forces and could also potentially act as a preliminary step toward use in pharmaceutical research for reduction or prevention of SMC proliferation due to altered mechanical forces during endovascular procedures.

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Introduction 

Following balloon angioplasty, the denudation of the endothelial cell layer allows exposure of smooth muscle cells (SMCs) to normally unseen shear forces from blood flow. The response of the SMCs to these mechanical forces includes apoptosis, differentiation, migration, hypertrophy, and proliferation. Balloon angioplasty–induced injury can be the etiology for the SMC change from a contractile to a synthetic phenotype.¹ Smooth muscle (SM) proliferation is believed to be an acute response to injury and can occur up to 7 days following balloon angioplasty, with most of the cells entering the growth cycle within the first 2-3 days following injury.² Following the acute response, the number of SMCs within the intima stays relatively constant up to a year following the endovascular procedure. Due to the stability of the SMC population with time, further increases in intimal hyperplasia are believed to be due to the production of extracellular matrix and connective tissue and to cellular hypertrophy.³ Maximum intimal thickening in humans usually occurs within 2-3 months of injury.⁴

Normal nondiseased SMCs reside primarily in the media layer of the blood vessel and are of a contractile phenotype. Following balloon angioplasty, medial SMCs can change from a contractile to a synthetic phenotype.¹ These synthetic cells migrate to the intima and proliferate and excrete extracellular matrix components such as elastin, collagen, and proteoglycans.⁵ The extracellular matrix production of synthetic SMCs is four to five times that of contractile SMCs. It is estimated that 50% of the SMCs that migrate from the media due to balloon angioplasty divide three times.² These cells make up eight-ninths of the resultant intimal SMC population.² The remaining one-ninth of the resultant intimal SMC population consists of the remaining 50% of SMCs activated from balloon angioplasty which simply migrated to the intima without further replication.² Synthetic SMCs are characterized by increased number of synthetic organelles such as rough endoplasmic reticulum, increased proliferation, decreased α-actin, loss of contraction ability, and increased cell volume.⁶ In contrast, contractile SMCs are generally spindle-shaped, nonproliferative, and able to contract and contain relatively more α-actin than synthetic SMCs.⁶ In summary, it is hypothesized that the degree of activation of SMCs following angioplasty affects the migration and proliferation characteristics of SMCs and directly contributes to the degree of intimal thickening and hyperplasia.

Percutaneous transluminal angioplasty (PTA) or balloon angioplasty specifications depend on the disease state of the vasculature, surgical preference, balloon diameter, injury length, and other contributing factors. The extent of the injury in terms of the transmural pressure exerted on the vessel wall and on the cells within the vessel wall can affect the cellular response to injury. Prior experiments in the literature have shown that a large injury without media damage initiated medial SMC replication but no intimal proliferation.⁷ The injury from balloon angioplasty can cause several disturbances within the blood vessel including platelet aggregation, endothelial denudation, growth factor production from injured and dead SMCs, attraction of leukocytes and macrophages from the inflammatory response, and mechanical sustained stretch of SMCs, which can all contribute to SMC proliferation.⁸

To date, there have been many studies designed to investigate the in vivo response of SMCs to injury following balloon angioplasty. Due to difficulty in maintaining concurrent shear and tensile forces to evaluate their effect on SMCs in vitro, limited studies have been performed using an in vitro model. We report the first use of an in vitro model of SMC response to mechanical injury with a concurrent shear and tensile force simulator developed to provide clinically significant levels of strain and shear stresses with statistical applicability.

With this model, the mechanical forces can be isolated from the biochemical signals and their influence on the SMC response to injury. The use of this model/simulator allows the acquisition of data that can act as a preliminary step in the pharmaceutical manipulation of SMCs and research into the reduction, and eventually prevention, of SMC hypertrophy and proliferation following endovascular procedures.

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Materials and Methods 

Simulator Design 

This simulator system consists of six independent channels, allowing for six independent cell populations to be assessed for increased statistical applicability and multiple analysis techniques. The system is comprised of two apparatuses, each with three channels, as demonstrated in Figure 1a. Design parameters were considered to resemble physiological conditions with minimal compromises, including laminar flow, developed flow at the cells, and applied shear stresses. Cells are grown on a flexible silicone membrane, which oscillates at 60 cycles/min (1 Hz) during mechanical testing.

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• Fig. 1 

  Demonstration of individual channel of the vascular mechanics simulator. a System configuration. b One of the vascular mechanics simulators during experimentation. c Top view of individual channel with entrance length and location of cell culture on silicone. d Side view of entrance side indicating cross-sectional area of channel.

A parallel plate geometry was chosen for this application, to expose a flat sample of cells to both concurrent shear and tensile forces. Figure 1 provides a pictorial explanation of the designed mechanical simulator. The entrance length of the channel to ensure fully developed flow at the site of the cells was governed by the resultant Reynolds number of 151 from the volumetric inlet flow (Q_in) of 350 mL/min. The Reynolds number (R_e) was determined from the equation applied to a rectangular channel:⁹

where b is the width of the channel (cm), ρ is the density of the fluid (kg/cm³), and μ is the fluid viscosity ([dynes ^∗ sec]/cm²). The Reynolds number was also used to establish the channel flow as laminar. This was considered highly laminar inflow since the Reynolds number was much less than 2,000 for this rectangular channel.⁹ This equation assumes that the fluid behaves as an incompressible Newtonian fluid.

The minimum entrance length parameter for a fully developed velocity profile, L_e (cm), was estimated as follows:¹⁰

The calculated required minimum entrance length of 7.55 cm was almost double at 13.95 cm to assure a fully developed velocity profile at the site of the cells, as shown in Figure 1d.

An equibiaxial strain profile was applied along the silicone membrane using vacuum similar to the technology implemented by Flexcell International (Hillsborough, NC). The Flexcell 3000 software was used to control strain and regulate vacuum pressure. Others have shown¹¹ that approximately the outer 20% of the silicone membranes located over the static post in the Flexcell system does not follow constant radial and circumferential strain profiles. Since the designed system utilized in this study implemented a similar static post geometry, cells were confined to the center of the membrane to promote constant radial and circumferential strain at the site of the cultured cells during mechanical testing by using a Teflon ring with an inner cross-sectional area of 2.75 cm² and an approximate height of 1.75 cm centered on the silicone surface over the center lexan post during the period of cell culture. This Teflon ring was removed prior to mechanical testing and only acted to confine the cells during cell culture.

Cell Culture 

Rat aortic smooth muscle cells (RASMCs, passage 11-14) obtained from VEC Technologies (Rensselaer, NY) were maintained with Dulbecco's modification of Eagle medium (DMEM) (10-013-CV; Mediatech, Herndon, VA) supplemented with 10% heat-inactivated fetal bovine serum (FBS, F-4135; Sigma-Aldrich, St. Louis, MO) and 1% antibiotic-antimycotic (A5955, Sigma-Aldrich). Cells were trypsinized with 0.25% trypsin with 0.2 mg/mL ethylenediaminetetraacetic acid (EDTA, E6511; Sigma, St. Louis, MO).

Cell Seeding for Mechanical Testing 

Silicone membranes (5.25 cm in diameter) for each station were cut from biomedical-grade silicone sheets (lot SM04106803; Specialty Manufacturing, Saginaw, MI) of 0.015-inch thickness and 40 durometer. These were sonicated in deionized water and assembled with the designed station plates. The assembled plates were steam-sterilized (3870M; Tuttnauer Brinkmann, Jerusalem, Israel) at 121°C for 30 min and then placed into the simulator apparatus over approximately 0.2 cc of silicone lubricant (51360; Loctite, Rocky Hill, CT). The lubricant was distributed using a sterile silicone stopper and application of 5% mechanical strain from the FlexerCell strain system (FlexCell International). The assembled system was then ultraviolet (UV)-sterilized for 3 hr. Following sterilization, sterile Teflon rings were centered on the silicone surface. The center area of silicone encompassed by the Teflon rings was coated with type I collagen (Vitrogen 100, 3 mg/mL; Cohesion Tech, Palo Alta, CA) diluted with sterile water to 50.0 μg/mL and allowed to dry for 72 hr in a sterile laminar flow hood according to Cohesion Tech protocol for formation of a thin collagen film. Prior to cell seeding, the membranes were rinsed with 1.0 mL of Dulbecco's phosphate-buffered saline (DPBS, 21-031 CM; Mediatech). Cells were counted using a hemocytometer and trypan blue, seeded at a density of 7.5 × 10⁴/mL and incubated for 48 hr to allow them to spread and adhere. One additional milliliter of medium was supplemented after 24 hr.

Simulation Experiment Groups 

Two experimental groups simulating in vivo conditions were investigated in this study in addition to a control group, where cells where not exposed to shear or tension. The injury model (IM) group consisted of cells exposed to high strains generated during balloon angioplasty followed by concurrent shear and tensile exposure. These cells represented the SMCs that are directly exposed to blood flow upon denudation of the endothelial monolayer. The cyclic tension (CT) group simulated a physiological control where cells are cyclically strained from 0% to 4% and not directly exposed to flow. Unloaded static controls (U) were used as adhesion and phenotypic behavior controls. Adhesion controls (A), which were fixed and counted prior to testing, assisted in determining problems of adhesion due to exposure of mechanical forces.

Mechanical Testing 

Prior to mechanical testing, confluence was assessed by microscopy of the unloaded static controls. The medium with 10% FBS was removed, and the RASMCs were maintained in quiescent medium consisting of DMEM without FBS supplementation for the duration of mechanical testing. Medium without FBS supplementation is commonly used in dynamic testing to investigate effects of changes in proliferation. During the application of mechanical testing, both dynamic experimental groups were first subjected to a preconditioning regimen of 0-4% cyclic strain for 2 hr at gradually increasing frequencies of 0.1 (30 min), 0.5 (30 min), and 1.0 (60 min) Hz. Following this preconditioning regimen, the CT group (n = 6) was subjected to 0-4% cyclic strain for 8 hr and the IM group (n = 6) was subjected to two 12% static circumferential streches held for 75 sec each with a 30 sec interval between stretches as commonly performed during balloon angioplasty. Following simulated balloon deployment, shear flow was resumed at 350 mL/min and cyclic strain was applied at 0-4% for 8 hr. The flow rate of 350 mL/min provided low wall shear stress (0.25 dynes/cm²) to promote cell proliferation. Static U samples, which were not subjected to dynamic force exposure, served as a control (n = 6) and A controls (n = 6), which were fixed prior to dynamic testing, served to determine the number of cells lost to problems of cellular adhesion.

Evaluation of Cellular Response 

Cell proliferation 

Each experimental group allowed for six samples. Cell proliferation was assessed using a quantitative (n = 3) and a qualitative (n = 3) method. The quantitative approach involved staining of cell nuclei with 4′,6-diamidino-2-phenylindole (DAPI, D-1306; Molecular Probes, Eugene, OR). For this method, cells were rinsed with DPBS following mechanical testing, fixed with 100% ethanol for 30 min, and rinsed again with DPBS. Samples were covered with 0.5 mL DAPI solution for 10 min and rinsed with DPBS. Samples were imaged with fluorescent microscopy (Diaphot 300; Nikon, Melville, NY) at 10 locations distributed throughout the sample. ImagePro Plus Analysis software (version 5.1; Media Cybernetics, Silver Spring, MD) was used to quantify the number of cells present.

The qualitative analysis involved the use of the Cell Titer 96 AQ_ueous ONE Solution Cell Proliferation Assay (3-(4,5-dimethylthiazol-2-yl)-5-(3-carboxymethoxyphenyl)-2-(4-sulfophenyl)-2H-tetrazolium inner salt [MTS], G-3580; Promega, Madison, WI), in which 1.25 mL was added to each well and incubated for 3 hr. The supernatant was collected individually, dispensed into cuvettes, and gently vortexed to promote sample uniformity. Four 125 μL samples from each supernatant were placed into a 96-well UV spectrophotometer plate. Optical density was determined at 490 nm (model DU^® 640B; Beckman Instruments, Fullerton, CA).

Phenotypic Marker 

SM α−actin as a marker cellular phenotype⁶ was assessed in cell extracts using standard Western blotting techniques. Expression of glyceraldehyde-3-phosphate dehydrogenase (GAPDH) was also evaluated and used as a normalization and control of protein loading. Cells were removed from the silicone membrane and lysed using an extraction buffer consisting of 0.5% Triton X-100 (T-9284, Sigma-Aldrich), 1% sodium dodecyl sulfate (SDS, L-4522, Sigma-Aldrich), 20 mM Tris (161-0716, Bio Rad, Hercules, CA), and 10 μL/mL protease cocktail inhibitor (P8340, Sigma-Aldrich). The BCA Protein Assay Kit (23225; Pierce, Rockford, IL) was then used to determine the total protein content of each sample. These samples were then diluted to 1 μg/μL using sample buffer consisting of β-mercaptoethanol, 0.5 M Tris-HCl, 10% SDS, 0.5% bromophenol blue, glycerol, and distilled water. Each sample was normalized to 10 μg/mL. Prior to gel electrophoresis, the samples were boiled for 5 min to linearize the proteins. A 10% polyacrylamide gel (161-0158, Bio Rad) was used in electrophoresis with 17.5 μL cell solution loaded per lane. Western blot analysis was performed using transfer to a polyvinylidene difluoride (PVDF) membrane (162-0184, Bio Rad). The WesternBreeze^® chromogenic Western Blot Immunodetection Kit (WB7103; Invitrogen, Carlsbad, CA) was used to detect levels of α-actin and GAPDH expression. The WesternBreeze Chromogenic Immunodetection Protocol was followed with the primary antibody mouse monoclonal to SM α-actin (ab18460-1; Abcam, Cambridge, MA) at a dilution of 1:1,000 and with the primary antibody anti-GAPDH mouse monoclonal (6C5) (CB1001; Calbiochem, San Diego, CA) at a dilution of 1:1,500. Densitometry was performed on the developed Western blots to evaluate the SM α-actin expression differences between the U (n = 3), CT (n = 3), and IM (n = 6) groups. The integrated optical density (IOD) of the bands was determined using an imaging densitometer (GS-700, Bio Rad) paired with Quantity One analysis software (version 4.1.1, 170-9600; Bio Rad). The ratios of SM α-actin to GAPDH were evaluated to normalize the data, and these values were used in further statistical analysis.

Statistical Analysis 

A statistical analysis of variance paired with Tukey's analysis was performed using SigmaStat (Systat Software, San Jose, CA) to examine the data on cell proliferation and SM α-actin, with p < 0.05 indicating a significant difference.

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Results 

Proliferation 

Acute studies (8 hr) using RASMCs demonstrated significant cell proliferation following angioplasty injury and shear exposure (IM group) when compared to “physiological function” cyclic tensile exposure (CT group). DAPI analysis indicated a 75% increase in the number of cells of the IM group compared to the CT group (p = 0.001). The cell number averages for the DAPI cell count were 3.45 × 10⁴ (± 1.7 × 10³ standard deviation [SD]) cells/mL for the CT group, 6.07 × 10⁴ (± 5.0 × 10³ SD) cells/mL for the IM group, 10.0 × 10⁴ (± 15.8 × 10³ SD) cells/mL for the combined averages of the static U group, and 8.2 × 10⁴ (± 16.1 × 10³ SD) cells/mL for the A group, as shown in Figure 2a. As shown in Figure 2b, MTS results supported the trends of the DAPI results, with mean resultant absorbance levels at 490 nm of 0.062 (± 0.028 SD) for the CT, 0.153 (± 0.04 SD) for the IM, and 0.345 (± 0.034 SD) for the U groups.

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• Fig. 2 

  Cell proliferation comparison between cyclic tensile (CT), injury model (IM), unloaded (U), and adhesion (A) (for DAPI only) test groups. a DAPI cell quantification indicated significant difference, with IM and U being different from all other groups (^∗p = 0.001). b MTS cell proliferation comparison, confirming significant difference between experimental groups (^∗p < 0.001).

Western Blot 

Western blot analysis was performed to detect levels of the contractile phenotype marker SM α-actin. Decreased levels of SM α-actin expression would indicate a more synthetic phenotype of the SMCs. The static unloaded (U1-U6) samples expressed a normalized SM α-actin IOD of 1.038 (± 0.103 SD). The normal function simulation in the CT group (CT1-CT3) demonstrated a 12% increase in SM α-actin, with IOD levels of 1.162 (± 0.098 SD); however, these results were not significantly different. Significant differences were observed with the balloon angioplasty injury model (IM1-IM3), with decreased normalized SM α-actin expression at an IOD level of 0.841 (± 0.046 SD) (p = 0.006) and a 19% decrease in SM α-actin expression when comparing the IM to the static U samples.

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Discussion 

The effects of individual mechanical forces on SMC behavior in an in vitro setting have been well documented by others. However, the presented model has been developed with the capability to more accurately represent concurrent mechanical forces due to the additional injury response from balloon angioplasty deployment on the reaction of SMCs to low wall shear stress exposure that stimulates SMC proliferation.

One purpose of the injury model was to simulate increased balloon deployment forces on cells in comparison to normal levels due to blood vessel expansion due to blood flow. During balloon angioplasty, high deployment forces of 5-10 atmospheres can negatively affect surrounding SMCs. Often, the procedure of balloon angioplasty is described as the compression of arterial plaque within the blood vessel to restore nominal lumen diameter. However, the higher rigidity of the plaque within the vessel walls compared to surrounding tissue results in greater transfer of the deployment force to surrounding tissue and SMCs.⁸ Further exposure of the SMCs to a physiological tensile dynamic force of 0-4% for less elastic arteries at 1 Hz resembled the normal tensile forces that SMCs are subjected to in vivo.

The shift of SMCs from a contractile to a more synthetic phenotype has a significant role in intimal hyperplasia following endovascular procedures. In vivo, normal nondiseased SMCs reside primarily in the media layer of the blood vessel and are of a contractile phenotype. Following balloon angioplasty, medial SMCs can change from a contractile to a synthetic phenotype.¹ These changes in the cell result in the cellular loss of ability to contract, increased protein secretion, and increased reaction to autocrine and paracrine growth factors.¹², ¹³, ¹⁴ Contractile SMCs contain relatively more SM α-actin than synthetic SMCs, which is why this was the chosen marker in this study.⁶ Mechanically, the stiffness of contractile SMCs has been shown to be higher than that of synthetic SMCs, which is proposed to be from an increased amount of actin within the cell.¹⁵

In vitro cell lines are generally characterized as a synthetic phenotype. When transferred to static culture, contractile vascular SMCs tend to change their phenotype to a synthetic state. Increased levels of applied tensile forces (>10%) have been shown to result in increased SMC proliferation¹⁶, ¹⁷ in addition to decreased expression of markers such as SM α-actin.¹⁷ When examining the results of this study, the static unloaded SMCs, normal to in vitro culture, served as a baseline for SM α-actin expression. Even at a short time point of 8 hr following injury, a low level of tensile forces similar to in vivo forces resulted in a higher level of SM α-actin expression, as seen in Figure 3. Although not significant, this suggests that application of a low level of tensile forces could stimulate differentiation toward a more contractile phenotype over time. Investigation of increased time points would be needed to explore this hypothesis. The significant decrease in SM α-actin expression shown in Figure 3 at the early time point of 8 hr exposure to shear and tensile forces following in vitro balloon angioplasty provides valuable insight to the early phenotypic shift due to applied mechanical forces.

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• Fig. 3 

  Western blot results for SM α-actin for cyclic tensile (CT), injury model (IM), and unloaded (U) test groups. a Densitometric picture of SM α-actin used to evaluate integrated optical density to compare expression. b Densitometric picture of GAPDH acting as protein loading control. c Normalization of SM α-actin expression to loading control GAPDH, indicating significant difference of SM α-actin expression in the injury model (^∗p = 0.006).

Proliferation results indicated in Figure 2 supported the concept of an early phenotypic shift. Comparing the dynamic groups CT and IM, DAPI results indicated a significantly different 75% increase in the number of cells present in the IM compared to the CT group simulating normal arterial function, indicating that the injury and exposure to concurrent shear and tensile forces resulted in increased proliferation of the injury model. As cell adhesion is a concern associated with dynamic experiments in both the CT and IM groups, a direct proliferation comparison of both groups with the static group is irrelevant. This was supported by the data for the A group indicating a significant loss of cells due to dynamic exposure. For this reason, the static U group was not included in this comparison.

Application of shear force to SMCs can have various effects on the cell response. Variations in flow behavior can invoke different responses along with the magnitude of the shear rate. Shear studies demonstrate the inhibition of in vitro SMC proliferation with high levels of shear stress.¹⁸ In general, shear stress levels above 10 dynes/cm² are considered high and atheroprotective.¹⁹ In contrast, areas of low shear stress (<10 dynes/cm²) promote an atherogenic response and increased SMC proliferation.¹⁹ Alterations in arterial geometry can result in areas of flow separation, stagnation, and recirculation. These disturbances in the velocity profile can result in areas of low wall shear stresses, therefore promoting SMC proliferation.

Application of tensile force to cells is usually in the form of circumferential, uniaxial, or biaxial application. The direction is usually adjusted to the application. Vascular cells are cyclically deformed from the dilation and relaxation of the blood vessel, usually represented by either circumferential or uniaxial testing. The cyclic strain response of vascular cells is important to both normal and abnormal cell function. It can affect cell proliferation, migration, apoptosis, morphology, and alignment.²⁰

Cell differentiation is an important issue in development and in vitro testing. When transferred to static culture, contractile vascular SMCs tend to change their phenotypes to a synthetic state. Cyclic tension has been shown to increase the expression of some of the markers of the differentiated state of SMCs. In a study by Birukov and collaborators,¹⁶ in addition to h-caldesmon upregulation, increased proliferation was seen, indicating a contradicting synthetic behavior, but the level of stress on the cells was higher than physiological stress (15%), which could affect the proliferation behavior of the cells. Butcher and collaborators¹⁷ demonstrated a phenotype shift in a three-dimensional collagen construct of RASMCs subjected to tensile forces corresponding to 10% strain at 1 Hz for 48 hr using a platen multiaxial strain system. The strained group results showed increased cell viability, decreased morphological elongation, decreased α-actin and calponin expression, and increased vimentin expression compared to the static control. These results suggest a shift from contractile toward a more synthetic phenotype.

There are different in vitro systems in use to study the effects of applied mechanical forces. For example, shear stresses are generally fluid induced using a parallel plate flow chamber based on pressure differences or by a cone on plate rotational device. In vitro tensile testing is modeled using devices which apply tensile forces in unidirectional, bidirectional, and radial/circumferential directions. The advantage of this system in this study is the combinatory effect of force application through use of combined principles of radial/circumferential tensile application and parallel plate shear application.

Although the developed system is an improved in vitro model, there are limitations of this study that will be addressed in further studies. As previously stated, application of mechanical forces can cause a variety of responses such as apoptosis, migration, hypertrophy, differentiation, and proliferation. Further cell analyses will investigate the possible responses of apoptosis, proliferation, and cell hypertrophy. Additionally, as discussed previously, there are other phenotypic markers that indicate synthetic or contractile behavior. Investigating other markers in addition to SM α-actin will assist in better characterization of the differentiation effects of applied mechanical forces on vascular SMCs.

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Conclusion 

This study presents an in vitro model to simulate in vitro balloon angioplasty followed by applied tensile and shear forces to exposed SMCs. The findings within this study demonstrate the efficacy of the developed model for in vitro angioplasty and the simulated mechanical environment to the cells. Additionally, the significant response at a short point stimulates further investigation for increased forces and time points for increased SMC response.

The clinical significance of this in vitro cell proliferation model is the ability to mimic clinically relevant mechanical stresses and to apply a more complex applied regimen to simulate clinical intervention combined with normal exposure. Hopefully, this model could ultimately be implemented for use in early pharmaceutical research for potential antiproliferative agents for SMCs in an attempt to prevent hyperplasia and restenosis following endovascular intervention.

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The authors acknowledge Cassie Gregory, Rebecca Cribb, and Dr. Agneta Simionescu for research support and the Greenville Hospital System for the funding of this work.

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