Electrospun Decellularized Lung Matrix Scaffold for Airway Smooth

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An Electrospun Decellularized Lung Matrix Scaffold for Airway Smooth Muscle Culture Bethany M. Young, Keerthana Shankar, Brittany P. Allen, Robert A. Pouliot, Matthew B. Schneck, Nabil Mikhaiel, and Rebecca L. Heise ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.7b00384 • Publication Date (Web): 13 Nov 2017 Downloaded from http://pubs.acs.org on November 18, 2017

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ACS Biomaterials Science & Engineering

Submitted to ACS Biomaterials Science & Engineering as an Article

An Electrospun Decellularized Lung Matrix Scaffold for Airway Smooth Muscle Culture Bethany M. Young1, Keerthana Shankar1, Brittany P. Allen1, Robert A. Pouliot1, Matthew B. Schneck1, Nabil Mikhaiel1, Rebecca L. Heise* 1,2 1

Department of Biomedical Engineering, Virginia Commonwealth University, 800 E. Leigh St,

Room 1071, Richmond, VA 23219 2

Department of Physiology and Biophysics, Virginia Commonwealth University School of

Medicine, 1101 East Marshall St, Richmond, Virginia 23298 *Corresponding author. Email: [email protected] KEYWORDS: electrospinning, decellularized extracellular matrix, airway smooth muscle, in vitro airway

ACS Paragon Plus Environment

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ABSTRACT

Chronic respiratory disease affects many people worldwide with little known about the intricate mechanisms driving the pathology, making it difficult to develop novel therapies. Improving the understanding of airway smooth muscle and extracellular matrix (ECM) interactions is key to developing treatments for this leading cause of death. With currently no relevant or controllable in vivo or in vitro models to investigate cell-ECM interactions in the small airways, the development of a biomimetic in vitro model with cell attachment, signaling, and organization is needed. The goal of this study was to create a biologically and structurally relevant in vitro model of small airway smooth muscle. In order to achieve this goal, a scaffold was engineered from synthetic poly-L-lactic acid (PLLA) and decellularized pig lung ECM (PLECM). PLECM scaffolds have improved physical characteristics over synthetic scaffolds, by exhibiting a significant decrease in the elastic modulus and an increase in hydrophilicity. Histological staining and SDS-PAGE showed that essential proteins or protein fragments found in natural ECM were present after processing. Human bronchial smooth muscle cells (HBSMCs) seeded onto PLECM 3D scaffolds formed confluent layers and maintained a contractile phenotype, as demonstrated by the organized arrangement of actin filaments within the cell and expected contractile protein expression of calponin 1. HBSMCs cultured on electrospun PLECM scaffold also increased alpha-1 type 1 collagen compared to those cultured on PLLA scaffolds. In summary, this research demonstrates that a PLLA/PLECM composite electrospun mat is a promising tool to produce an in vitro model of the airway with the potential for a better understanding of bronchiole smooth muscle behavior in diseased or normal states.


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INTRODUCTION Airway smooth muscle (ASM) plays a key role in lung homeostasis and disease pathologies. In cases of asthma and chronic obstructive pulmonary disease (COPD), ASM is responsible for significant airway narrowing caused by airway hyperresponsiveness, severe hyperplasia, hypertrophy 1, and airway wall remodeling 2. It is less understood how these pathologies are triggered and what molecular responses can be used as a therapeutic target

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. A debilitating

consequence of airway disease is disruption of normal ASM repair which causes excess extracellular matrix (ECM) production and development of fibrotic tissue over time. Further investigation into the connection between ASM and ECM production requires a physiologically relevant culture model 4 that mimics the natural airway smooth muscle cell environment. Up to now, airway smooth muscle models rely on in vivo whole airways, ex vivo airway segments, or in vitro substrates such as tissue culture plastic (TCP) or protein coatings 2. There are limitations with most of these current approaches, including the absence of biological and mechanical control with in vivo whole airway assessments Investigations using TCP

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and ex vivo segments

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.

or engineered platforms such as electrospun materials or coatings

currently lack either 3D architecture or tissue-specific protein profiles

9–11

. Platforms with both

3D topographies that mimic the ECM structure and tissue-specific biological signals on the micro and nano levels will give more accurate ASM response 12, leading to more translatable and reliable therapies for airway smooth muscle pathologies. Incorporating natural ECM components into engineered scaffolds have shown to have the potential to promote constructive tissue remodeling after injury in many areas of tissue engineering including vascular 13, skeletal muscle 14, central nervous systems 15, and abdominal


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wall repair 9,16. These materials offer cellular attachment and tissue-specific signaling inherent to natural ECM proteins (i.e., collagen, elastin, fibronectin, and laminin). Specific proteins can influence particular cell behaviors that are important to the complete cell response seen in vivo. Fibronectin plays a key role in cell attachment and ASM spreading, the presence of both fibronectin and collagen 1 can affect the proliferation of ASM important in airway repair, and laminin plays a key role in maintaining a contractile phenotype with increased expression of myosin heavy chain, calponin, and alpha-smooth muscle actin

9,12

. Interactions of cellular

integrins with ECM ligands mediate contractility and ECM production of ASM by mechanotransduction, gathering information on ECM quantity and component concentration

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.

Mechanical properties and ECM profile can be altered by disease 5, modulating the ECM turnover and expression of phenotype-specific genes

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to propagate diseased ECM

characteristics. ECM is not uniform across all tissues within the body, and the dissimilarities determine cell behavior including the activation or deactivation of internal cell signaling

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. For example,

several studies have shown that organ-specific ECM alone will transition MSC phenotype towards a lineage native to the original organ 17,19,20. There is an ongoing cycle between the cells laying down ECM specific to cell phenotype and the ECM signaling the cells towards a specific tissue phenotype

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. For these reasons, organ-specific decellularized tissues have been used to

create an appropriate microenvironment for proper phenotypic expression of native cells 21–23. ECM structure may also be mimicked by fabricating the scaffold with electrospun 3D microfibers that maintain cellular morphology, polarity, and differentiation compared to flat substrates

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. Electrospinning allows for the ability to tailor diameter size, density, and porosity

by manipulating parameters such as flow rate, needle gauge, or mandrel speed

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. When


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layered into a three-dimensional mat, the resulting scaffold has a highly porous architecture allowing nutrient movement and multiaxial mechanical strength. Studies investigating the importance of 3D structure similar to ECM, specifically fiber size and distribution have shown that these spatial ques alone can maintain healthy, tissue specific cell phenotype and organization 11,27,28

to give a more appropriate in vitro study. These advantages have been utilized in this study by incorporating ECM into

electrospinning of an FDA approved, biodegradable synthetic polymer 29. Electrospinning with ECM proteins will offer customizable nanostructure, mechanical profile, and degradation properties with practical biological signaling optimized for an accurate human analog for smooth muscle function in the airways. The objective of this research was to create a biomimetic platform for in vitro studies of the ASM. This in vitro model will not only give control to research experiments but also decrease the reliance on animal models. In the present study, we characterized the mechanical and microstructural properties of a naturally derived lung ECM and Poly-L Lactic Acid (PLLA) electrospun nanofibrous fabric. The basic cell-scaffold interactions were examined using primary human airway smooth muscle cells. MATERIALS AND METHODS Tissue Decellularization Pig lungs (donated from Smithfield Hams) were decellularized by established protocol

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. All

chemical perfusion steps were done with a hand pump at room temperature with unspecific pressure. The lung was perfused through the trachea and vasculature with ultrapure water before being drained passively. The lungs were then perfused with 0.1% Triton X-100 (Fisher Scientific) and submerged for 24 h at 4° C. After overnight incubation, the tissue was rinsed and


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perfused with a 2% sodium deoxycholate (Sigma) solution and left in sodium deoxycholate for 24 h at 4° C. The lung was rinsed again then filled with 1M sodium chloride (Fisher Scientific) solution for 1 h at room temperature. After one h, the lung was then rinsed before adding DNase 1 (Sigma) for another hour at room temperature. All remaining debris and agents were removed by rinsing with ultrapure water and then 1X PBS five times at room temperature. All tissue was then dissected to remove larger, cartilaginous airways and cut into smaller pieces for processing. Processing of the decellularized tissue to achieve a fine powder was done by first lyophilizing, then freezer milling using a SPEX 6700, as described previously 30. The powder was stored at 80° C until used for the electrospinning solution. Scaffold Fabrication A solution of 35 mg/mL pig lung extracellular matrix (PLECM) was found to be the maximum concentration of PLECM that could be successfully electrospun. A lesser concentration was also tested (17.5 mg/mL) to determine if the largest amount of PLECM would be detrimental to the biological response, but it was determined that the cell viability was improved with higher PLECM concentrations quantified with picogreen DNA quantification (ThermoFisher, Figure S1). Therefore 35 mg/mL PLECM was used for all further testing. The 35 mg/mL PLECM solution was made by adding 46.67 mg/mL of powdered PLECM to three-quarters of the final volume of 1,1,1,3,3,3-Hexafluoro-2-propanol (HFP, Oakwood Chemical) and mixed for 24 h at room temperature on a vortexer. At the same time, a 400 mg/mL PLLA (Goodfellow Cambridge Limited, part number: LS402531) in HFP solution was made with a quarter of the final volume and mixed for the same period of time. The final 35 mg/mL PLECM solution was completed by straining the first 46.67 mg/mL PLECM solution through a stainless-steel wire cloth (150x150 count, type 304) into the 400 mg/mL PLLA solution to remove any large aggregates of ECM that


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had not gone into solution. The mixture was then vortexed for one h at room temperature to create the final 35 mg/mL PLECM solution. PLLA only control solution was made by mixing 100 mg/mL PLLA with the final desired volume of HFP and vortexed for one h at room temperature. Some solutions were applied directly onto tissue culture plates, and the HFP was evaporated under a hood for 24 h at room temperature to be used as coatings. Both solutions were electrospun onto a rotating rectangular mandrel that was also moving side to side for full mandrel coverage. Electrospinning was performed with the mandrel 27 cm away from the needle and at a flow rate of 4.5 mL/hr. 35 mg/mL PLECM solutions were electrospun with a voltage of 27kV and the PLLA only solutions were electrospun with a voltage of 15kV. Voltages were arrived at by the establishment of a stable Taylor cone.

Scaffold Architecture Scaffolds with and without PLECM were dried in a desiccator for a minimum of 24 h before sputter coating with platinum for 60 sec. Scaffolds were imaged with Scanning Electron Microscopy (SEM, Hitachi SU-70 FE-SEM) to show fiber size and architecture. Decellularized airway tissue before milling was dissected away from non-airway tissue and platinum coated for SEM imaging. Airways were identified prior to SEM preparation by the presence of cartilage. Average fiber diameter and distribution of the scaffolds and decellularized airway submucosa/adventitia layers were determined by taking measurements of 200 plus fibers per image using NIH Image J. Fiber and Cell Alignment


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Fiber alignment of the scaffolds and human bronchial smooth muscle cells (HBSMCs, Lonza) cultured onto scaffolds were determined from analysis of SEM images using Orientation J within NIH Image J software. First, the mandrel spinning orientation of each sample was identified for comparison and normalization with overall scaffold fiber alignment. Three or more SEM images of each sample were measured to find overall average fiber orientation (°) using a Gaussian distribution32. Each measurement of fiber orientation is given an alignment index that correlates to how accurately the scaffold is aligned. An alignment index of 1 would indicate that all of the fibers are oriented in the same direction and an index closer to 0 would correlate to a random fiber distribution. Contact Angle Measurement and Porosity Wettability of the composite nanofiber mat was determined with the sessile drop method using an OCA 15 Goniometer with controlled automatic liquid deposition and a computer-based image processing system. 5 µL of dH2O were deposited onto the fiber scaffolds. The percent porosity of both scaffolds was determined using a liquid displacement method33. The chosen displacement liquid was ethanol because it is not a solvent for PLLA, it will allow for full penetration into the sample, and is used as the disinfection method used for cell culture so any changes in porosity after ethanol exposure would be the same used in cell culture studies. A sample of known weight was submerged into ethanol and put on a shaker for 24 h to ensure full impregnation of the scaffold. Scaffolds with ethanol were briefly blotted with a Kimwipe before weighing the samples once again. Volumes of the scaffold and the ethanol within the scaffold were determined using the densities of ethanol, PLLA, and the 35 mg/mL PLECM mixture. For 35 mg/mL PLECM solution density calculations, the loading efficiency of


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PLECM into the electrospun scaffold was found by calculating the ratio of PLLA and PLECM in the solution before electrospinning to the final weight of the scaffold. It was assumed that density of the pure PLECM would be mostly dictated by collagen and therefore the final density of 1.25 g/cm-3 was arrived at by the ratio of PLLA to PLECM 34. The final porosity was then calculated with: Є = (ܸா௧ைு )/(ܸா௧ைு + ܸ௦௖௔௙௙௢௟ௗ )

(1)

Mechanical testing Elastic properties of the scaffolds with and without cells, as well as intact pig lung parenchyma and small cartilaginous airways, were tested using MTS Bionix 200 with TestWorks 4.0 Software, using established tensile testing protocols. Electrospun samples were taken from various scaffolds produced on different dates to cover batch to batch variability. Scaffolds without cells were submerged for one week in Smooth Muscle Basal Media (SmBM, Lonza) to simulate wet physiological conditions. Scaffolds with cells were cultured for one week in a 24 well plate with dog bone shaped scaffolds held to the bottom of the wells with O-rings (Aflas). 50,000 HBSMCs were seeded onto each dog bone in Smooth Muscle Growth Media (SmGM, Lonza), changed every two days. Distal lung tissue was separated from non-cartilaginous airways and cut into dog bone shapes with similar thicknesses. Small cartilaginous airways were removed, cut longitudinally, and opened before punching. SDS-PAGE SDS-PAGE was performed to determine which ECM proteins and the amount of each protein present in the scaffolds post decellularization and electrospinning. 100 mg of intact tissue, 100


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mg of decellularized tissue, and 200 mg of each scaffold were put into a mixture of 1% SDS, 25 nM Tris, and 4.5 M Urea and heated to 60 ⁰C for one h, to increase the amount of protein resuspended. The loading efficiency of PLECM into the electrospun scaffolds was used to determine the amount of PLECM powder needed to have similar amounts of PLECM compared to 200 mg of 35 mg/mL PLECM electrospun scaffold. The PLECM powder treated with HFP was allowed to evaporate fully after dissolving in HFP. Both were then put in 1 mL of 1% SDS, 25 nM Tris, and 4.5 M Urea solution and heated to 60⁰C for one h. Protein concentration was measured using a Peirce™ BCA protein assay kit. All samples except PLLA were balanced to the protein concentration of the 35 mg/mL PLECM/PLLA scaffold. 5 ug/uL of protein loaded into each well of a 15 well, Mini-PROTEAN TGX stain-free gel (BIO-RAD) and run for 25 minutes in a Mini-PROTEAN tetra system (BIO-RAD) with running buffer. Resulting gels were imaged using a BIO-RAD gel imager. Histology Scaffolds frozen in OCT (Tissue-Tek) were cryosectioned at a 10 to 20 µm thickness with a Thermo FSE Cryostat and stored frozen. Before staining, slides were heated to 100°C to increase specimen attachment. Sections of 35 mg/mL and PLLA only were stained with Masson’s trichrome (Sigma) including hematoxylin (Sigma) following manufacturers protocol. Imaging of ECM proteins within the scaffolds was done with the use of a light microscope. Collagen and Elastin quantification Collagen and elastin quantification procedure followed an established protocol

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. 10 mg of

electrospun scaffolds (PLLA and 35 mg/mL PLECM) and 10 mg of PLECM powder were used for both quantifications. The electrospun scaffolds and PLECM powder were solubilized in 0.1


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M NaOH at 98°C for 45 min. The supernatant was removed and lyophilized before hydrolyzing for 24 h in 6M HCl at 110°C. The samples were lyophilized again and resuspended in 2 mL of DI water. For the collagen and elastin quantification, the protein was examined with a hydroxyproline assay (Sigma) and a ninhydrin assay according to previously published methods 30

. Resulting data for ECM powder was normalized based on loading efficiency of the PLECM

into the scaffolds. Fiber elemental distribution Scaffolds were dried in a desiccator for a minimum of 24 h and then were carbon coated for SEM Energy Dispersive X-ray Spectrometry (EDX) using a Hitachi SU-70 FE-SEM. Both scaffolds were analyzed by taking 20,000 or more points. Collagen type 1 (Proteintech) was also used to stain cells on the scaffolds according to manufacturer’s protocol and imaged using Zeiss LSM 710. Cell and media specifications Scaffolds and coatings were all disinfected with ethanol for 15 min, then rinsed three times with 1X PBS before cultured with cells. HBSMCs were passaged when confluence reached 80%. HBSMCs between passages 2-8 were expanded in SmGM with medium changed every 48 to 72 h. All scaffolds and coatings were anchored to the bottom of tissue culture plates with Aflas ORings (24-well plates) fitting tightly to the well sides or silicone (6-well plates, Loctite), placed on the underside of the scaffolds or coatings that have become detached from the bottom of the plate during the evaporation process to secured them to the bottom of the well. O-rings were sterilized by autoclave and silicone was assumed to be sterile after curing and were also disinfected with ethanol while attached to the scaffolds. O-rings were tested for changes in cell


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viability using Live/Dead quantifications, and no difference was seen in cell counts with the addition of O-rings. Live/Dead Cell Viability Assay HBSMCs were seeded at a density of 50,000 cells onto PLLA and 35 mg/mL PLECM/PLLA scaffolds in a 24 well plate. After 4 h and 72 h in culture, reagents were added to label live cells (calcein AM, Thermo Fisher Scientific) and dead cells (Ethidium homodimer-1, Thermo Fisher Scientific) according to manufacturer’s protocols. After incubating for 30-45 min, the samples were mounted onto glass slides and imaged with an Olympus IX71 inverted microscope. Three images were taken of each sample and each quantified using Image J. SEM Imaging 200,000 HBSMCs were seeded onto scaffolds in 24-well plates and cultured for one week. All samples were rinsed three times with PBS before fixing with 4% paraformaldehyde for 20 minutes and rinsed again three times with PBS. After the initial fixation, the samples were put through secondary fixation with osmium tetroxide for one h. Samples were ethanol dehydrated for critical point drying. The samples were dried using a Tousimis Critical point dryer, platinum coated for 60 sec, and imaged using a JEOL JSM-5610LV SEM. qPCR Gene expression of 1,000,000 HBSMCs seeded onto scaffolds or coatings of various concentrations PLLA and PLECM for one week were quantified using qPCR. Cells and scaffolds were put into 1 mL of TRIzol (Life Technologies) and homogenized using 0.5 mm zirconium bead homogenization tubes and a microtube homogenizer (Beadbug) and any remaining


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insoluble pieces of the scaffold were removed. RNA was then extracted following manufactures protocol. RNA concentration and quality were quantified using a Take3 microvolume plate (BioTek) and an Epoch reader. RNA was purified using a RNeasy kit (Qiagen) for samples that did not have a 260/280 ratio of 2 +/- 0.2. Using a high capacity cDNA reverse transcription kit (Applied Biosystems), balanced RNA was converted to cDNA. qPCR was run on CFX connect real-time system (BIO-RAD) with powerSYBR Green PCR Master Mix (Applied Biosystems). The fold change in gene expression of α-SMA, CNN1, MYH11, and COL1 A1 (IDT) was calculated relative to two housekeeping genes (GAPDH and 18s) and tissue culture plastic using the delta-delta Cq method. Primer sequences are listed in Table 1. Gene α-SMA

Protein Alpha Smooth Muscle Actin

Forward Primer Sequence 5' CCG ACC GAA TGC AGA AGG A 3'

CNN1

Calponin 1

5' GTC AAC CCA AAA TTG GCA CCA 3'

MYH11

Smooth Muscle Myosin 11

5' CGC CAA GAG ACT CGT CTG G 3'

COL1A1 GAPDH 18S

Alpha-1 Type 1 Collagen 5' GTG CGA TGA CGT GAT CTG TGA 3' Glyceraldehyde 3-phosphate dehydrogenase 5’ CTC TGC TCC TCC TGT TCG AC 3’ Ribosomal RNA 5’ GCA ATT ATT CCC CAT GAA CG 3’ Table 1: Primer sequences of contractile genes used in qPCR.

Immunofluorescence Scaffolds in a 24-well plate were seeded with 100,000 HBSMCs for one week. Scaffolds were fixed with 4% PFA and rinsed with PBS three times. To visualize cell morphology and actin distribution, FITC labeled phalloidin (Cell Signaling) was used to stain the cells on scaffolds according to manufacturer’s protocol and imaged using a Zeiss LSM 710. 2.5 Statistics


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Unless otherwise stated, GraphPad Prism was used for all statistical analysis. One-way ANOVA with Tukey post-hoc comparisons or T-tests were done when appropriate to determine significance. Significance is represented by: *p-value < 0.05, **p-value < 0.01, ***p-value < 0.005 unless otherwise stated. RESULTS Scaffold architecture Electrospun scaffolds were imaged using SEM to examine the characteristics of the fibers. Electrospinning PLLA and PLLA/PLECM composite created a scaffold with both nanometer and micrometer-sized fibers (Figure 1(A-D)). Electrospinning with our parameters created fibers similar to the size of fibers in whole decellularized lung tissue (Figure 2(E-F)). The most uniform fibers in Figure 1(E) are seen in the regions where smooth muscle would most likely be located below the basement membrane of the airways, in the submucosa and adventitia regions (see supplemental figures for a lower magnification image, showing entire airway structure). In order to quantify and compare fiber diameter of the decellularized tissue to the scaffolds, three SEM images were taken of each scaffold or decellularized airway, and 200 or more measurements of fiber width were averaged. Average fiber diameter for PLLA only electrospun scaffolds was 712.4 nm+/- 462.4 nm with a range of 54 nm to 4702.1 nm. 35 mg/mL PLECM scaffolds had an average fiber diameter of 806.4 nm +/- 432.1 nm with a range of 232.5 nm to 4783 nm. The mean fiber diameter within the decellularized airways was 837.25 +/-763.22 which was similar to those of the electrospun scaffolds. The range of the fiber sizes for the decellularized airways was 25 nm to 4477 nm. The only variation in fiber range from decellularized airway fiber range was seen in the minimum fiber diameter of 35 mg/mL PLECM


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scaffolds with a 207.5nm difference. Fiber distribution for electrospun scaffolds (Figure 1(G-H)) shows the majority of fibers to have a diameter from 250 nm to 800 nm, while the decellularized tissue has a bimodal curve with peaks in the 50 nm to 450 nm and 1050 nm to 1450nm ranges. Average scaffold thickness was found to be of 139µm+/- 53µm.

Figure 1. Scaffold and decellularized tissue fiber diameter. SEM of electrospun 35 mg/mL PLECM/PLLA (A-B) and PLLA only(C-D) scaffolds to compare to decellularized ECM fibers (E-F) that relate directly to natural fibers found in the airway environment. Fiber distribution of (G) 35 mg/mL PLECM/PLLA, (H) PLLA only electrospun scaffolds, and (I) decellularized


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airways were analyzed for comparison. Annotations show the location of the airway basement membrane (BM) and submucosa (SM) where ASM would be located and fiber distribution was analyzed. n=3. To characterize the fiber alignment within the scaffolds (Figure 2 (A, C, E, F, I)) was found to be partially dictated by the rotation of the mandrel during electrospinning process causing significant cellular alignment (Figure 2 (B, D, G, H, J). Representative aligned scaffolds (Figure 2 A-H) were used to show fiber and HBSMC orientation that resulted in distributions with large peaks of aligned fibers around the mandrel orientation indicating high fiber alignment (normalized to mandrel orientation at 0°). The average fiber and cell alignments were taken from three different scaffolds to determine if all electrospun scaffolds show overall alignment. 35 mg/mL PLECM scaffolds were on average 28.38 +/- 20.80° away from mandrel spinning direction and PLLA only scaffolds were 14.31 +/- 9.44° away from the mandrel spinning direction. For HBSMC alignment, an average difference from mandrel spinning direction of 10.41 +/- 12.39° for cells on 35 mg/mL PLECM scaffolds and 9.16 +/- 6.56° for cells on PLLA only scaffolds with no significant differences between PLECM scaffolds and PLLA only scaffolds for fiber or cell alignment. The higher the peak of the angle distribution and alignment index, indicates more precisely aligned scaffolds. The average fiber alignment index for 35 mg/mL PLECM and PLLA only scaffolds were 0.16 +/- 0.06° and 0.15 +/- 0.08° respectively and an average cell alignment index of 0.16 +/- 0.04° and 0.24 +/- 0.05° respectively.


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Figure 2. Fiber and HBSMC alignment analysis. Representative images of 35 mg/mL PLECM and PLLA only scaffold fiber alignment (A, C) and corresponding cell alignment (B, D) respectively with the oval representing the average alignment angle. Quantitative analysis of 35 mg/mL and PLLA only fiber (E, F) and cell angle distribution (G, H) respectively, normalized to mandrel spinning angle at 0°. The accuracy of average alignment angle for scaffold fibers (I) and cells on scaffolds (J) represented by alignment index. n=3. Data are presented as mean +/- st. dev with no significance between groups.


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Contact Angle Measurement and Porosity Representative images of the water contact angle with 5µL of dH2O onto 35 mg/mL PLECM and PLLA only scaffolds with the sessile drop method are shown in Figure 4(A-B). Overall the water contact angle decreased with the addition of PLECM by 14.1°. Porosity measurements done by ethanol displacement show no significant difference between PLLA only and 35 mg/mL PLECM scaffolds with a similar average porosity of 84.65 +/- 5.80% and 86.75 +/- 0.65% respectively. Mechanical testing Uni-axial tensile testing was performed on scaffolds after one-week in basal media and scaffolds with HBSMCs cultured for one week to determine elastic modulus compared to natural tissue tested under similar testing parameters. Native tissue testing of non-cartilaginous lung tissue had a smaller average elastic modulus of 0.46 +/- 0,09 MPa than small cartilaginous airways as expected, with an average elastic modulus of 2.31 +/- 0.76 MPa. With respect to the electrospun scaffolds, the addition of PLECM decreased the average elastic modulus by 34.11 MPa compared to PLLA only. HBSMCs significantly changed scaffolds’ mechanical properties by decreasing the average PLLA scaffold modulus to 16.35 +/- 2.94 MPa and the 35 mg/mL PLECM modulus to 14.03 +/- 5.92 MPa. 6 of 11 samples were untested from the 35 mg/mL with HBSMCs group due to loading failure, potentially indicating a further decrease in strength of those samples not represented in the data. Scaffold degradation experiments were also conducted for one week in PBS to rule out natural degradation in a physiological solution as a cause for mechanical property changes, but no degradation was found in either scaffold over that time period. Scaffold degradation experiments were also conducted for one week in PBS to rule out


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natural degradation in a physiological solution as a cause for mechanical property changes. No degradation was found in either scaffold over that period of time. Figure 3(C) compares scaffolds, with and without cells to native tissue. Overall the addition of PLECM significantly decreased the mechanical properties of scaffolds with and without cell intervention to more closely mimic the properties of natural tissue as expected. In order to determine the effect of alignment on scaffold elastic modulus, samples were tested in the transverse and perpendicular directions to the mandrel rotation, and no significant differences were found with respect to testing direction. Mechanical testing was also done to investigate the effect of ethanol disinfection on scaffolds by comparing elastic modulus of scaffolds treated with ethanol and untreated scaffolds and found no significant difference (data not shown).


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Figure 3. Mechanical and physical properties of electrospun scaffolds. Contact angle measurements of 35 mg/mL PLECM (A) with an average contact angle of 129.2° +/- 1.9° and PLLA only (B) with an average contact angle of 129.2° +/- 3.5°. Tensile testing for scaffolds with and without PLECM under wet conditions and with or without cells compared to native lung tissue (C). 35 mg/mL with HBSMCs and PLLA w/ HBSMCS represent respective scaffolds with HBSMCs cultured for one week.35 mg/mL PLECM/PLLA and PLLA only represent scaffolds were soaked in media for one week without cells. Data are presented as mean +/- st. dev. n=4-11. * p

Electrospun Decellularized Lung Matrix Scaffold for Airway Smooth

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