Article pubs.acs.org/journal/abseba
Flow-Induced Vascular Network Formation and Maturation in ThreeDimensional Engineered Tissue Barak Zohar,†,∥ Yaron Blinder,†,§,∥ David J. Mooney,‡,§ and Shulamit Levenberg*,† †
Department of Biomedical Engineering, Technion-Israel Institute of Technology, Haifa 3200003, Israel School of Engineering and Applied Sciences, Harvard University, 29 Oxford Street, Cambridge, Massachusetts 02138, United States § Wyss Institute for Biologically Inspired Engineering at Harvard University, 3 Blackfan Circle, Boston, Massachusetts 02115, United States ‡
S Supporting Information *
ABSTRACT: Engineered three-dimensional (3D) constructs have received much attention as in vitro tools for the study of cell−cell and cell−matrix interactions, and have been explored for potential use as experimental models or therapeutic human tissue substitutes. Yet, due to diffusion limitations, the lack of stable and perfusable blood vessel networks jeopardizes cell viability once the tissue dimensions extend beyond several hundred microns. Direct perfusion of 3D scaffold cultures has been shown to enhance oxygen and nutrient availability. Additionally, flow-induced shear stress at physiologically relevant levels, positively impacted endothelial cell migration and alignment in various two-dimensional (2D) culture models and promoted angiogenic sprouting in microfluidic systems. However, little is known about the effect of flow on vascularization in implantable 3D engineered tissue models. The present study investigated the effect of direct flow-induced shear stress on vascularization in implantable 3D tissue. The differential effect of various levels of shear stress, applied while maintaining constant culture conditions, on vascular parameters was measured. Samples grown under direct flow conditions showed significant increases (>100%) in vessel network morphogenesis parameters and increases in vessel and extracellular matrix (ECM) protein depth distribution, as compared to those grown under static conditions. Enhanced vascular network morphogenesis parameters and higher colocalization of alpha-smooth muscle actin (α-SMA) with endothelial vessel networks characterized the specific contribution of direct flow to vessel network complexity and maturation. These observations suggest that flow conditions promote 3D neovascularization and may be advantageous in attempts to create large-volume, clinically relevant tissue substitutes. KEYWORDS: engineered tissue, vascular networks, endothelial cells, flow bioreactors, fluid shear stress maturation, and stabilization both in in vitro9,18−20 and in vivo5,21 models, presumably by triggering mechanical stimulation and enhancing nutrient transport. Flow-induced shear stress applied to 2D monolayer “flow-over” in vitro models22,23 has been shown to regulate endothelial migration, apoptosis, proliferation, and alignment. Furthermore, in various microfluidic assemblies,24−28 physiologically relevant shear stress (1− 10 dyn/cm2) has been shown to regulate morphological processes such as angiogenic sprouting. In line with these findings, direct flow stimulated vascular morphogenesis and angiogenesis-related gene expression in 3D collagen and alginate constructs embedded with endothelial cells.29,30 The effect of flow has proven beneficial in the culture of engineered tissues such as cardiac,31,32 hepatic,33,34 cartilage,35 and bone tissue.36 However, the effect of shear stress on vascularization in 3D engineered constructs remains question-
1. INTRODUCTION Engineered implantable constructs generally contain cells seeded on three-dimensional (3D) polymeric scaffolds. These scaffolds are designed to support tissue formation and mimic the cell−cell and cell−extracellular matrix (ECM) interactions of naturally occurring 3D niches.1,2 Currently, construct thickness, limited by inadequate nutrient perfusion, is one of the main obstacles in tissue engineering.3−5 Following transplantation, insufficient blood perfusion further threatens construct integrity and can result in failed integration and collapse of the engineered tissue.6 Coculture of endothelial cells (ECs) with supportive cells, such as fibroblasts, skeletal muscle cells,7 or cardiomyocytes,8 on 3D scaffolds has been shown to vascularize the engineered tissue in vitro and consequently facilitate its survival and viability upon implantation.7 This approach relies on natural mechanisms, such as vasculogenesis and angiogenesis, which drive the de novo formation of blood vessel networks.9−11 These mechanisms are highly regulated by a variety of chemical molecules12−14 and mechanical forces, such as strain15,16 and shear stress.17 Fluid shear stress has been demonstrated to play a crucial role in vessel formation, © 2017 American Chemical Society
Special Issue: Regenerative Biomaterials Received: January 11, 2017 Accepted: April 10, 2017 Published: April 11, 2017 1265
DOI: 10.1021/acsbiomaterials.7b00025 ACS Biomater. Sci. Eng. 2018, 4, 1265−1271
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ACS Biomaterials Science & Engineering able, as the process can also be affected by accelerated mass transport. In addition, the effect of flow-induced shear stress on vessel network complexity, maturation, and stabilization in implantable 3D engineered tissue remains to be established. Herein, we assessed the biomechanical impact of flow-induced shear stress on vascularization of 3D implantable engineered tissue, by applying various flow regimens while maintaining the medium circulation rate. The effect of constant flow-induced shear stress was compared to that of higher, intermittent flowinduced shear stress. The effects elicited by direct versus indirect perfusion were also compared.
2. MATERIALS AND METHODS 2.1. Scaffold Preparation. Poly-L-lactic acid (PLLA) (Polysciences, Warrington) and poly-L-glycolic acid (PLGA) (Boehringer Ingelhein) (1:1) were dissolved in chloroform to yield a 5% polymer solution. The solution (0.24 mL) was loaded into molds containing 0.4 g of sodium chloride particles with a size distribution ranging between 200 and 600 μm (isolated by sieve size-exclusion). The solvent was allowed to evaporate, and the sponges were subsequently immersed, for 8 h, in distilled water (replaced every hour), to leach the salt and create an interconnected pore structure. The sponges were sliced to ∼28 mm3 circles (diameter, 6 mm; width, 1 mm). 2.2. Cell Culture and Scaffold Seeding. Zs-green-expressing human adipose microvascular endothelial cells (HAMECs; ScienceCell), isolated from human adipose tissue, Zs-green-expressing human umbilical vein endothelial cells (HUVECs; Lonza), and neonatal human dermal fibroblasts (HNDFs; Lonza) were grown separately on standard tissue culture plates. HNDFs were cultivated in DMEM (Gibco), supplemented with 10% fetal bovine serum (FBS) (Hyclone), 1% nonessential amino acids, 0.2% β-mercaptoethanol, and 1% penstrep (Sigma-Aldrich). HAMECs were cultivated in endothelial cell medium (ScienceCell), supplemented with 5% FBS (ScienceCell) and endothelial cell growth supplement (ScienceCell), and HUVECs were cultivated in EGM-2 medium, supplemented with a bullet kit containing FBS, hydrocortisone, hFGF-β, VEGF, R3-IGF1, hEGF, GA-1000, and heparin (Lonza). Cells were cultured in a 5% CO2 humidified incubator at 37 °C and harvested during passages 5− 8. Then, ECs and HNDFs were mixed at a ratio of 5:1 (total 6 × 105 cells per scaffold) in 14 μL of human fibrin gel, prepared from a 1:1 mixture of thrombin solution (15 mg/mL, Johnson & Johnson Medical, Israel) and human fibrinogen solution (5 mU/mL, Johnson & Johnson Medical), and then rapidly pipetted and seeded upon the PLLA/PLGA scaffolds. Scaffolds were then incubated (30 min, 37 °C, 5% CO2) on a 12-well nontissue culture plate. Coculture medium (a 1:1 mixture of the two respective cell media) was added (1−3 mL per scaffold) and replaced every 2−3 days. Scaffolds were cultured under static conditions (37 °C, 5% CO2) in a 12-well nontissue culture plate for 5 days before being transferred to the flow bioreactors. 2.3. Flow Bioreactor Configuration. The flow bioreactor contained flow chambers and a perfusion system. Flow chambers were prepared from commercially available sterile scaffold holders (P6D, Ebers Medical) (Figure 1A) or in-house developed poly(methyl methacrylate) (PMMA) chambers that were designed to test the effect of direct flow compared to indirect perfusion (control) conditions (Figure 1C). The perfusion system was composed of a glass reservoir bottle, silicone tubes, and a multichannel peristaltic pump (Ismatec or EBERS TEB1000 pumps) (Figure 1B). Before use, the flow bioreactor was cleaned with 1% Liquinox detergent and distilled water, washed with PBS, and sterilized in an autoclave (121 °C, 30 min). The bioreactor was then incubated with relevant culture media for at least 24 h before scaffolds were inserted. After verifying the integrity of the system, flow chambers were loaded with preseeded scaffolds and plugged into the perfusion system. Flow bioreactors were placed inside the designated EBERS incubator (EBERS TEB1000, highly humidified 37 °C, 5% CO2) for further culturing under different flow conditions. Constant and intermittent flow profiles were programmed using the EBERS TEB1000 software.
Figure 1. Flow bioreactor configurations. (A) Schematic drawing of a commercially available scaffold holder for direct perfusion (Ebers Medical). (B) Schematic view of a typical perfusion bioreactor in which culture medium flows directly through a porous cellular scaffold. Reproduced with permission from ref 37. Copyright 2009 Wiley Periodicals, Inc. (C) Schematic drawing of the in-house-designed direct flow and control chambers. (D) Flow velocity characterization by the CFD model (FLUENT software) in the direct flow and control chambers while applying a flow rate of 0.1 mL/min. 2.4. Shear Stress Estimation Using a Predesigned Computational Fluid Dynamics Model. A computational fluid dynamics (CFD) model for direct flow through the complex 3D microstructure of a porous scaffold was designed37 to predict physiological shear stress levels sensed by endothelial cells. The model considered the accumulation of cells and ECM proteins in a scaffold 7 days postseeding by introducing geometrical modifications to differences in cell layer thicknesses. Flow-associated parameters were extracted from the model (Table 1). Both intermittent and constant flow rates were
Table 1. Mean Velocity and Shear Stress through a Porous 3D Scaffold, at the Tested Flow Rates, As Estimated by the CFD Model flow profile
flow rate (mL/min)
mean velocity (cm/s)
mean/maximum shear stress (dyn/cm2)
constant intermittent
0.1 0.6
0.093 0.55
0.75/2 4.5/13
set to generate identical medium circulation rate of 3 scaffold volume replacements in a minute. The direct intermittent flow was set as 0.6 mL/min, corresponding to a mean shear stress of 4.5 dyn/cm2 in a duty cycle of 16.7% (10 s), applied for 1 min. The constant flow was set at 0.1 mL/min for both direct and indirect perfusion (control) conditions, corresponding to a mean shear stress of 0.75 dyn/cm2 for the direct flow conditions (Table 1 and Figure 2A). Flow velocity was characterized inside the direct flow and control chambers by the CFD model using the FLUENT software to evaluate media circulation rates and to ensure negligible shear stress under indirect perfusion (control) conditions (Figure 1D). Medium circulation rates were estimated by 1266
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Figure 2. Vessel network formation in constructs cultured for 5 days under static conditions, followed by 2 days under static, constant (0.1 mL/min), or intermittent (0.6 mL/min) flow conditions. (A) Shear stress profiles corresponding to constant and intermittent volumetric flow rates, as estimated by the CFD model. (B) EC epifluorescence (green) and respective Angiotool segmentation (green = EC epifluorescence, red = segmented vessels, and light blue = segmented junctions) (scale bar −500 μm). (C) Normalized mean vessel network morphological measures in the explant area, vessel area and density, junction count and density, overall vessel length, average vessel length, number of vessel end points, and mean E lacunarity as determined by Angiotool (n = 4). The results were normalized to measurements made under static conditions. Error bars represent the standard error of the mean (SEM). * indicates statistical significance (*p < 0.05). the CFD model to be 3 and 0.3 scaffold volume replacements in a minute for direct flow and control conditions, respectively. Cells exposed to constant and intermittent flow were cultured for 2 days, and those exposed to direct flow and indirect perfusion (control) were cultured for 6 days. 2.5. Scaffold Sectioning and Immunohistochemical Staining. Constructs were fixed in 4% paraformaldehyde (Electron Microscopy Sciences) and then embedded in an agar gel (5% low melting temperature SeaPlaque agarose in distilled water, Lonza). Sections (200 μm-thick) were obtained from the agar-embedded constructs using a vibratome (VT1000S, Leica) and then permeabilized with 0.5% Triton X-100 (Bio Lab Ltd.), rinsed several times with PBS, and blocked with 5% bovine serum albumin (BSA) (w/v, Millipore). Subsequently, samples were incubated with primary monoclonal rabbit antihuman von Willebrand factor (vWF) antibody
(1:150; Abcam) or CD31 (1:50, Cell Marque), monoclonal mouse antihuman Alpha-smooth muscle actin (α-SMA) (1:50, Dako) antibody, or with collagen I (1:200; Abcam) and collagen IV (1:200;Dako) antibodies overnight. Samples were then rinsed several times with PBS and then incubated (3 h, room temperature) with Cy3-conjugated goat antimouse (1:100, Jackson Immuno-research laboratory) and Alexa 488-conjugated antirabbit (1:800, Thermo Fisher Scientific) antibodies, diluted in staining buffer and then stored in PBS until imaging. 2.6. Confocal Imaging and Analysis. Flow-oriented sides and cross-sections of the scaffolds were scanned with an LSM700 confocal microscope (Zeiss). The depth of imaging was 300−500 μm, split into at least 20 z-stacks. Three-dimensional (XYZ) confocal z-stacks were converted to 2D TIFF stacks, by performing high-intensity z1267
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Figure 3. Colocalization of collagens I and IV with EC structures in scaffold cross-sections. (A) A cross-section of a scaffold cultured for 7 days under static conditions (scale bar −500 μm). (B) A higher magnification of the ECM structure of the cross-sectioned scaffold sample shown in (A) (scale bar −100 μm). (C) A cross-section of a scaffold cultured for 5 days under static conditions, followed by 2 days under direct constant flow conditions (scale bar −500 μm). (D) A higher magnification of the ECM structure of the cross-sectioned scaffold sample shown in C (scale bar −100 μm).
Figure 4. Effect of direct flow on vessel distribution in 3D scaffolds. (A) Endothelial cell epifluorescence, as observed in static- and indirect-flowcultured scaffold cross-sections (scale bar −500 μm). (B) The percentage of epifluorescence area at increasing distances (30 μm intervals) from the surface of scaffolds subjected to static versus flow conditions ((n = 4, different scaffolds for each condition). projections using NIH ImageJ software. Stacks were then prepared for morphological analysis by contrast enhancement (stack histogram equalization and normalization, 0.4% saturated pixels). Each frame was processed using the Angiotool interface, as described in the online user manual (see http://ccrod.cancer.gov/confluence/display/ROB2/ Quick+Guide), to quantify vessel percentage area, total number of junctions, junction density, total and average vessel lengths, total number of vessel end points, and lacunarity. EC bulk distribution was quantified by isolating and masking 30 μm-thick layers at increasing distances from the exterior edge of each cross-section image and quantifying the degree of endothelial area coverage by applying an image analysis algorithm (MATLAB code). Colocalization of vWF (ECs) and α-SMA was quantified by the Colocalization Threshold plugin (ImageJ), as described in the online user manual (see http:// imagej.net/Colocalization_Analysis#Colocalization_Threshold). Colocalization percentage is presented as the percentage of abovethreshold vWF (ECs) intensity colocalized with above-threshold α-
SMA intensity. Autothreshold was determined using the Costes method. 2.7. Statistics. Measurements were performed in triplicate, at minimum, and images were scanned, processed, and analyzed using an identical setup each time. For streamline vascular parameters presentation, data were normalized to the relevant control treatment. Normalized means were plotted, with error bars representing the standard error of the mean (SEM). Statistical comparisons were performed using the Student’s t test with a 95% confidence limit (twotailed and unequal variance). Differences with a p-value 100%), junction count and density (∼200%), overall vessel length (>100%), average vessel length, and number of vessel end points (∼100%) as compared to those of static-cultured constructs (Figure 2C). Although intermittent flow-cultured constructs showed higher vascular measures, as compared to those of static-cultured constructs, the mean differences were all smaller than those measured for constant flow-cultured constructs. 3.2. Vessel Depth-Distribution and ECM Protein Structures. Examination of scaffold cross-sections revealed a clear difference in vascular and ECM protein structures in constructs cultured under static versus flow conditions. Flowcultured constructs displayed higher EC, collagen I, and collagen IV epifluorescence intensities within the scaffold (Figure 3A,C). High-magnification images demonstrated the higher colocalization of collagens I and IV with endothelial vascular structures formed following exposure to flow versus static conditions (Figure 3B,D). In addition, constructs cultured under static conditions showed high endothelial densities in the exterior 60 μm of the construct, which decreased at increasing distances from the scaffold perimeter. Conversely, constructs cultured under flow conditions showed a lower endothelial density in the outermost layers but a higher and relatively constant density at greater depths (Figure 4B). 3.3. Effect of Direct Flow on Vessel Formation and Maturation. To isolate the effect of flow-induced shear stress on vascularization, we compared vessel formation and maturation under direct flow versus indirect perfusion (control) conditions (Figure 1C), while applying identical perfusion rates in both flow chambers. At 11 days postseeding, endothelial network branching was observed for both flow-culture regimens (Figure 5A). Direct flow-cultured constructs displayed significantly lower mean lacunarity and significantly higher vascular measures, such as vessel area and density, junction count and density (>50% increase), and overall vessel length, when compared to those of the control (Figure 5B). Although the increase in average vessel length was statistically insignificant, it was 2-fold higher following exposure to direct flow versus control conditions (Figure 5B). In addition, vessel network maturity, as reflected by α-SMA expression, and its colocalization with endothelial vascular structures were greatest in vascularized areas of direct flow-cultured samples (Figure 6E) and were considerably low in static- or control-cultured samples (Figure 6C,D). More specifically, ∼ 70% of the ECs in constructs subjected to direct flow colocalized with α-SMA, while only ∼40% colocalized with α-SMA following exposure to control or static conditions (n = 3, p < 0.05).
Figure 5. Vessel network formation in constructs cultured for 5 days under static conditions, followed by 6 days under direct flow (0.1 mL/ min) or control conditions. (A) EC epifluorescence (green) and respective Angiotool segmentation (green = EC epifluorescence, red = segmented vessels, and light blue = segmented junctions) (scale bar −500 μm). (B) Normalized mean vessel network morphological measures of the explant area, vessel area and density, junction count and density, overall vessel length, average vessel length, number of vessel end points, and mean E lacunarity, as determined by Angiotool (n = 3). The results were normalized to the measures obtained following exposure to control conditions. Error bars represent the standard error of the mean (SEM). * indicates statistical significance (*p < 0.05).
vessel network morphology, we designed two experiments in which we applied different perfusion approaches. 3D engineered constructs cultured under constant flow conditions exhibited significant increases (100−200%, p < 0.05) in almost all vascular measures, as compared to those grown under static conditions. Direct flow through the 3D construct triggered a significantly different bulk distribution of vascular structures and ECM protein expression, as compared to static conditions. More specifically, endothelial structures in static-cultured constructs were distributed more closely to the exterior construct edges, with a substantial drop in endothelial presence
4. DISCUSSION AND CONCLUSIONS Flow-induced, physiologically relevant levels of shear stress have been shown to influence endothelial alignment in 2D “flow-over” models and to promote vessel sprouting in microfabricated systems. Thus, we hypothesized that exposure of 3D engineered tissue constructs to direct flow conditions would have similar effects. To assess the effect of direct flow on 1269
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media convection through the scaffold that increases shear stress stimuli and accelerates mass transport within the scaffold. To better define how its effect is mediated, we compared the effect of direct flow to indirect perfusion (control) conditions while maintaining similar flow rate. Our results indicate that direct flow contributed (20−50%, p < 0.05) to vessel formation. Moreover, direct flow-culture constructs showed significant higher EC and α-SMA colocalization as compared to controlcultured and static-cultured constructs (∼75%, p < 0.05 and p < 0.01 respectively), indicative of enhanced vascular maturation and stabilization. In conclusion, this work demonstrated that direct flow through implantable engineered 3D tissue induces vessel network formation, increases vessel and ECM depth-distribution, and enhances vessel maturation and stabilization. The positive effect correlated with the convection of media through the scaffold in direct-perfusion mode. Our findings provide a deeper understanding of vascularization processes that are induced and maintained by flow. These observations are expected to promote judicious utilization of flow stimuli in tissue engineering, toward the creation of real-scale, vascularized artificial tissues that may function better and survive longer in vivo.
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ASSOCIATED CONTENT
* Supporting Information S
The Supporting Information is available free of charge on the ACS Publications website at DOI: 10.1021/acsbiomaterials.7b00025. DAPI and α-SMA segmented area analyses in crosssections of scaffolds cultured for 5 days under static conditions, followed by 6 days under direct flow or indirect perfusion (control) conditions (PDF)
Figure 6. vWF (ECs) and α-SMA colocalization in scaffolds subjected to direct flow (0.1 mL/min) versus static and control culture conditions. (A) Top and cross-sectional views of ECs in a scaffold cultured for 6 days under direct flow. (B) Top view in higher magnification (scale bar −50 μm). (C−E) Cross-sectional view of scaffolds cultured for 6 days under direct flow, control, or static conditions (scale bar −100 μm). (F) vWF (ECs) and α-SMA colocalization percentage, as determined from cross-sections of scaffolds subjected to direct flow, control, or static culture conditions (n = 3, different scaffolds for each condition). Error bars represent the standard error of the mean (SEM). * indicates statistical significance (*p < 0.05 and **p < 0.01).
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AUTHOR INFORMATION
Corresponding Author
*Tel:+972-48294810. Fax: +972-48294809. E-mail: shulamit@ bm.technion.ac.il. ORCID
Shulamit Levenberg: 0000-0001-5471-7339 Author Contributions ∥
within the scaffold core. Conversely, flow-cultured constructs exhibited slightly lower endothelial density at the scaffold exterior but increased and uniform distribution deeper within the scaffold volume. These observations demonstrated that medium convection through the 3D scaffold promotes cell viability and vitality, ECM protein secretion, and vessel network formation at depths beyond diffusion limits (∼150 μm). A relatively low constant flow of 0.1 mL/min, inducing an estimated shear stress of 0.75 dyn/cm2, applied for only 48 h, was sufficient to positively impact angiogenic processes. The reduced vascular measures recorded in intermittent versus constant flow-cultured constructs may be the result of a positive effect of continuous shear when in a physiological range. In both cases, there was identical minutely medium exchange; thus, the differences between set-ups were unlikely to be due to major differences in mass transport. The positive effect of flow on vessel formation can be ascribed to the perfusion system that improves mass transport by generating better mixing of media near the scaffold and/or
B.Z. and Y.B. contributed equally to this work.
Notes
The authors declare no competing financial interest.
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ACKNOWLEDGMENTS This research was supported by FP7 European Research Council Grant 281501, ENGVASC (to S.L.)
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REFERENCES
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DOI: 10.1021/acsbiomaterials.7b00025 ACS Biomater. Sci. Eng. 2018, 4, 1265−1271
