Engineering Microvascular Networks in LED Light-cured Cell-Laden

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Tissue Engineering and Regenerative Medicine

Engineering Microvascular Networks in LED Light-cured Cell-Laden Hydrogels Nelson Monteiro, Wenting He, Cristiane Miranda Franca, Avathamsa Athirasala, and Luiz Bertassoni ACS Biomater. Sci. Eng., Just Accepted Manuscript • DOI: 10.1021/acsbiomaterials.8b00502 • Publication Date (Web): 08 Jun 2018 Downloaded from http://pubs.acs.org on June 12, 2018

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ACS Biomaterials Science & Engineering

Engineering Microvascular Networks in LED Light-cured Cell-Laden Hydrogels 1 Nelson Monteiro , Wenting He1, Cristiane Miranda Franca1, Avathamsa Athirasala1, Luiz E. Bertassoni 1,2,3* 1

Division of Biomaterials and Biomechanics, Department of Restorative

Dentistry, OHSU School of Dentistry, 2730 SW Moody Avenue, Portland, Oregon, 97201. 2

Department of Biomedical Engineering, School of Medicine, Oregon Health

and Science University, 3181 SW Sam Jackson Park Rd, Portland, OR, 97239. 3

Center for Regenerative Medicine, Oregon Health and Science University,

3181 SW Sam Jackson Park Rd, Portland, OR, 97239. * Corresponding author Email: [email protected] Abstract The success of tissue engineering inevitably depends on the fabrication of tissue constructs that can be vascularized and that anastomose with the host vasculature. In this report we studied the effects of light-emitting diode (LED)

photopolymerized

gelatin

methacryloyl

hydrogels

(GelMA),

encapsulated with stem cells from the apical papilla (SCAP) and human umbilical vein endothelial cells (HUVECs), in promoting vasculature network formation as a function of hydrogel physical and mechanical properties, as well as total cell density. Lithium acylphosphinate (LAP) was used as the photoinitiator in concentrations of 0.05, 0.075, 0.1% (w/v). GelMA hydrogel precursors of 5% (w/v) were encapsulated with co-cultures of SCAPs and HUVECs at different cell densities (1x, 5x and 10x106 cells/ml) and

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photocrosslinked for 5 s. Results suggested that the compressive modulus of GelMA hydrogels increased as a function of LAP concentration, and had a maximum stiffness of 3.2 kPa. GelMA hydrogels photopolymerized using 0.05 or 0.075% LAP, which had an average of 1.5 and 1.6 kPa of elastic modulus respectively, had the most efficient vasculature formation after 5 days, and these results were further enhanced when the highest cell density (10x106 cells/ml) was used. Immunofluorescence images showed that SCAP cells spread in close contact with endothelial networks and expressed alpha smooth muscle actin (αSMA), which is suggestive of their differentiation into pericyte-like cells. αSMA expression was also apparently higher in hydrogels polymerized

with

0.05%

LAP

and

10x106

cells/ml.

In

conclusion,

photopolymerization of GelMA hydrogels using an LED-light source can be an effective method for potential chair-side/in-situ procedures for engineering of vascularized tissue constructs in regenerative medicine. Keywords:

Vascularization,

GelMA,

SCAP,

HUVEC,

lithium

acylphosphinate, LED, visible light 1. Introduction Vascularization is the most basic requirement for tissue survival. A functional vasculature guarantees adequate oxygenation, nutrient delivery, and removal of waste products to all cells in the body. Therefore, the lack of a functional vascular supply in engineered tissue constructs results in the formation of a necrotic core with dead cells and unviable new tissue formation

1-2

. Inevitably, therefore, the success of regenerative medicine

relies on our ability to understand and controllably engineer 3D vascularized tissues. Thus, the development of strategies for scaffold vascularization in2 ACS Paragon Plus Environment

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vitro is a challenge that must be met before tissue engineering can effectively be translated into clinical use - a reality that remains elusive despite significant progress in the field 1, 3. The process of angiogenesis and vasculogenesis, relies on complexly orchestrated biological events, such as the morphogenesis of endothelial cells into cord-like structures, followed by maturation into hollow capillaries, and recruitment of perivascular mural cells (or pericytes) 4-5. Altogether, these cells remodel the newly-formed vascular network via angiogenic sprouting, resulting in a stable and dense vascular plexus; a process that has been extensively studied previously

2-3

. Several studies have shown that the

combination of endothelial cells (ECs) with stem cells from multiple sources can result in the formation of vascular networks both in-vitro and in-vivo

6-13

.

Accordingly, it has been demonstrated that while ECs form the luminal aspect of growing vessels, stem cells can differentiate adjacent to the growing vasculature to function as perivascular α-SMA-expressing pericytelike mural cells 8. A critical factor in engineering vascularized tissues by co-culturing endothelial and mesenchyme-derived stem cells is that the physical properties of the matrix where cells are seeded is known to play a critical role in influencing stem cell fate decisions and overall response

14-16

. We have

demonstrated that ECs respond drastically to variations in stiffness when seeded in monolayer in planar hydrogel substrates hydrogel microchannels

17

, or in 3D printed

18

. Similarly, EC proliferation, network formation and

stem cell differentiation have been shown to be influenced by the degree of methacrylation of gelatin methacryloyl (GelMA) hydrogels,

6

which interferes 3

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with matrix degradation and mechanical properties, among other things19-20. Despite these relevant observations, direct evidence of the contributions of matrix physical and mechanical properties to the formation of mature vascular networks in 3D hydrogels remains scarce. We hypothesize that hydrogels of relative lower stiffness, greater porosity and faster degradation rates, would enhance microvasculature formation in cell-laden 3D hydrogels irrespective of degree of methacrylation. Similarly, although the influence of paracrine signaling of mesenchymal cells and ECs has long been appreciated, it remains unclear whether modulating cell-cell communication in a paracrine fashion by simply increasing cell seeding density could enhance the formation of vascular networks in engineered tissue constructs. To address these knowledge gaps, here we sought to determine the influence of matrix physical properties and cell seeding density on the formation of microvascular networks in engineered tissue constructs. To that end, we co-cultured human umbilical vein endothelial cells (HUVEC) with stem cells from the dental apical papilla (SCAP), the latter of which has demonstrated a strong trophic effect over many other cell types in vitro 21. In addition, here, we present a novel method for photopolymerization of cell-laden GelMA hydrogels using an LED visible-light source that has remained virtually unexplored in the field of regenerative medicine. These light sources are not only widely utilized clinically, but are also known to overcome the detrimental consequences associated with ultra violet (UV) light photopolymerization, such as the formation of free radicals upon longer UV exposure, DNA damage and impaired cellular function we

argue

that

controlling

vasculature

22-26

formation

. In summary,

in

LED-light 4

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photopolymerized cell-laden hydrogels could represent an important step towards effectively translating the concept of engineering vascularized tissues in-situ for regenerative applications. 2. Materials and methods 2.1 . Cell culture SCAPs were donated by Dr. Anibal Diogenes (University of Texas) and were obtained as previously published

27

. Briefly, fragments of the apical papilla

from third molars were digested by incubation with a solution of 3 mg/ml collagenase type I (Worthlington Biomedical, Lakewood, NJ) and 4 mg/ml dispase (Sigma, St Louis, MO) for 30 minutes. Cells were centrifuged (1500 rpm, 2 min) and re-suspended in alpha-minimum essential medium (a-MEM; Gibco, Grand Island, NY) supplemented with1x L-glutamine (Gibco), 10% fetal bovine, penicillin (100 U/ml; Gibco), and streptomycin (100 mg/ml; Gibco). Cells were seeded and expanded to 70%–80% confluence and used up to passage 10. GFP expressing HUVECs (cAP-0001GFP, Angioproteomie, USA) were expanded up to passage 6 in endothelial cell growth medium (EGM2, Lonza, USA) with EGM2 BulletKit (cc3162, Lonza, USA) on 0.1% gelatin coated substrates. All cells were cultured in humidified 5% CO2 at 37°C, and were cryopreserved in 10% dimethylsulfoxide (DMSO) in appropriate culture media until use. 2.2 . Gelatin methacryloyl (GelMA) synthesis The synthesis of GelMA was performed as described previously

28

. In brief,

porcine skin type A gelatin (10% w/v) (Sigma, St Louis, MO, USA) was dissolved in 50°C Dulbecco’s phosphate buffered saline (DPBS, Sigma). 8% (v/v) methacrylic anhydride (Sigma) was added to the solution dropwise, 5 ACS Paragon Plus Environment


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allowing the reaction to proceed for 2 hours, and subsequently stopping it with a 5x dilution of DPBS. The solution was dialyzed against distilled water using a 12-14 kDa dialysis tubing at 45±5°C for five days with two water changes per day. The resulting macromer was then lyophilized for 5 days and stored at room temperature until further use. 2.3. Physical and mechanical properties GelMA macromer at a concentration of 10% (w/v) was dissolved in DPBS with lithium acylphosphinate (LAP, Tokyo Chemical Industry, L0290) photoinitiator of 0.05, 0.075, 0.1% (w/v). GelMA hydrogels were fabricated by dispensing the gel precursors into Poly(dimethylsiloxane) (PDMS, Sigma) molds measuring 5 mm in diameter and 2.5 mm in height and exposing the samples to a visible LED light using a clinically available light curing device (VALO, Ultradent) which emits light in the visible range between 395 and 480 nm through a 10.5 mm diameter curing tip. Hydrogel degradation was determined by incubating GelMA hydrogel disks (n=4) for 2, 5, 8 and 16 hours at 37˚C in a 1 U/ml collagenase solution (MP Biomedical). After incubation, the non-degraded hydrogel fragments were retrieved and lyophilized overnight. Degradation percentage was determined by calculating the weight ratio of degraded versus intact hydrogel samples at each time point. Compressive modulus of the hydrogel samples was determined using an unconfined compression testing method on a universal mechanical testing machine (MTS Criterion), at a loading rate of 1 mm/min, and determining the slope of the linear region corresponding to 0%–10% strain (n=4).


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2.4. Cell encapsulation – Effect of LAP concentration and Cell Seeding Density For quantification of SCAPs and HUVECs response as a function of hydrogel physical properties, cells were trypsinized and counted using Countess™ II FL automated cell counter (Life Technologies), and re-suspended in 10% GelMA hydrogel prepolymer at a cell density of 5 x 106 cells/ml in a 1:4 cell ratio. In order to adjust the hydrogel physical and mechanical properties without interfering with the density of cell adhesion ligands naturally present in gelatin, the hydrogel prepolymer was mixed in DPBS containing either 0.05, 0.075 or 0.1% (w/v) LAP. Cell-laden hydrogel constructs were fabricated by dispensing 5 µl of a cell-laden hydrogel precursor on plastic Petri dish, and compressing the cell-laden droplet with a TMSPMA ([3(Methacryloyloxy)propyl]trimethoxysilane,

Sigma)

coated

glass

slide

supported by two parallel cover slips, to form 100 µm thick GelMA constructs, as described previously

18

. Photocrosslinking was achieved by exposing all

samples to light for 5 seconds, as explained above (N=8). To test the effect of cell seeding density, SCAP and HUVECs at 1:4 cell ratio were re-suspended in the GelMA hydrogel prepolymer (10%, 0.075% PI), at densities of 1x106 cells/ml, 5x106 cells/ml and 10x106 cells/ml. Samples were photopolymerized as above. Cell-laden hydrogel constructs were cultured in endothelial cell growth medium (EGM2, Lonza, USA) supplemented with EGM2 BulletKit (cc-3162, Lonza, USA) for 7 days, and the medium was changed every two days.



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2.5 Vascular network analysis Analysis of vasculature formation was performed by imaging the GFPexpressing HUVECs in the cell-laden hydrogels at different time points using an automated ﬂuorescence microscope (EVOS FL Auto, Life Technologies). Images were processed using Fiji (ImageJ, NIH), and vasculature formation was quantified using AngioTool (NIH) following a pre-optimized routine of image segmentation, skeletonization and determination of default thresholds for vessel diameter, signal intensity, and removal of small particles. Vessels were compared as a function of vessel percent area, total vessel length, average vessel length, and vascular branching index. 2.6 Immunofluorescence Cell-laden hydrogels were fixed in paraformaldehyde (4%) and permeabilized with 0.1% Triton X-100 in DPBS for 25 min. The hydrogels were then blocked with bovine serum albumin (BSA) (1.5%, Sigma–Aldrich) in DPBS for 1 h. After washing with PBS, samples were incubated with primary mouse monoclonal antibody (anti-α-SMA, Abcam, 1:400) overnight at 4°C. Samples were washed with PBS and incubated with secondary antibody (1:250, goat anti-mouse Alexa Fluor 555, Invitrogen) for 3 h. This was followed by rinsing in 0.1% PBS, staining of the nuclei using a NucBlue staining kit (ThermoFisher Scientific, Waltham, MA USA) for 20 min at 37 °C. Samples were examined using a ﬂuorescence microscope (EVOS FL Auto, Life Technologies) and images were further processed using Fiji (ImageJ, NIH).


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2.7 Data Analysis Statistical analysis was performed using GraphPad Prism 6. The values represent averages ± standard deviations. A two-way ANOVA followed by Tukey post-hoc test (α = 0.05) was used to analyze the differences between different LAP concentrations, and cell densities. 3. Results

3.1 Physical and mechanical properties of LED-light photopolymerized

Figure 1 – Physical and mechanical characterization of GelMA hydrogels. (a) Stress-strain curves of GelMA hydrogels polymerized with 0.05, 0.075 and 0.1% LAP showed an increase in (b) compressive modulus with increase in photoinitiator concentrations from 0.05% to 0.1%. No significant differences were observed between 0.05% and 0.075% (w/v) LAP. Hydrogels with 0.1% LAP followed (c) slower degradation kinetics than those with 0.05% and 0.075% LAP. P

Engineering Microvascular Networks in LED Light-cured Cell-Laden

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