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There is a need of new materials and architectures for tissue engineering and regenerative medicine. Based upon our recent results developing novel scaffold architecture, we hypothesized that this new architecture would foster vascularization, a particular need for tissue engineering. We report on the potential of superporous hydrogel (SPH) scaffolds for in vivo cellular infiltration and vascularization. Poly(ethylene glycol) diacrylate (PEGDA) SPH scaffolds were implanted in the dorsum of severe combined immunodeficient (SCID) mice and harvested after four weeks of in vivo implantation. The SPHs were visibly red and vascularized as apparent when compared to the non-porous hydrogel controls which were macroscopically avascular. Host cell infiltration was observed throughout the SPHs. Blood cells and vascular structures, confirmed through staining for CD 34 and smooth muscle alpha actin, were observed throughout the scaffolds. This novel soft material may be utilized for cell transplantation, tissue engineering, and in combination with cell therapies. The neovasularization and limited fibritic response suggest that the architecture may be conducive to cell survival and rapid vessel development.
Tissue engineering, which applies the principles of biology and engineering to the development of functional substitutes for damaged tissues or organs (Langer, 1997), has been increasingly discussed as an option to overcome limitations of transplanting tissues. While in vitro tissue engineered constructs consist of cells seeded on suitable scaffolds with minimal mortality, in vivo cells within these constructs cannot survive by diffusion alone (Folkman, 1995) and constructs must support angiogenesis. Angiogenesis within scaffolds also allows successful engraftment of the construct into the surrounding host tissue (Rucker et al., 2006). To this end, porous hydrogel scaffolds are being investigated (Druecke et al., 2004; Ford et al., 2006; Wake et al., 1994).
The inherent hydrated architecture of synthetic hydrogels imparts mechanical properties similar to soft tissues and extracellular matrix (Peppas et al., 2000) which are natural hydrogels. Natural and synthetic hydrogels have been extensively examined and applied to hard and soft tissue even when the hydrogel is significantly weaker than the bulk tissue (Lee and Mooney, 2001; Yamamoto et al., 2000). The hydrophilicity, favorable biological recognition properties, and the ability to encapsulate cells within the hydrated network (Burdick and Anseth, 2002; Elisseeff et al., 1999) have suggested the applicability of poly(ethylene glycol) diacrylate (PEGDA) hydrogels for tissue engineering constructs (Nguyen and West, 2002). Additionally, various strategies have been reported to modulate the degradation of PEGDA scaffolds which typically aim for degradation of the scaffold in concert with new tissue growth (Hudalla et al., 2008; Kim et al., 2008; Kraehenbuehl et al., 2008). However, the photopolymerization techniques which are commonly used for preparation of non-porous hydrogels, have the disadvantages of relatively large diffusion distances and limited cell-cell interactions within the scaffolds (Nuttelman et al., 2004). While the rate and depth of vascular ingrowth into scaffolds are influenced by the presence, size, and interconnectivity of pores (Druecke et al., 2004; Gerecht et al., 2007; Landers et al., 2002), engineering interconnected porous architectures within hydrogels is a challenge. Various techniques have been employed for generating a porous architecture (Flynn et al., 2003; Ford et al., 2006; Landers et al., 2002; Liu et al., 2000; Weinand et al., 2007; Yoon et al., 2004), but interconnected pores are not always formed. Without interconnected pores or inherent degradability, cell penetration and proliferation within the scaffolds is inhibited (Chirila et al., 1993). In addition to pores being necessary for cell invasion, surface modification, either by conjugating peptides with cell adhesive sequences (Hern and Hubbell, 1998; Hersel et al., 2003) or infiltrating the hydrophobic scaffold with a hydrophilic polymer, have been proposed to be required to make hydrophilic scaffolds conducive to cell attachment and migration (Mooney et al., 1995). To this end, hydrophilic polyvinyl alcohol (PVA) sponges have been extensively studied as in vivo models for wound healing and foreign body reaction (Diegelmann et al., 1986; Kyriakides and Bornstein, 2003). While early stages (1-3 weeks) of cell invasion into the PVA sponge is characterized by neo-vascularization, the later stages (3 weeks) of the response are typical of a foreign body response leading to a fibrotic encapsulation (Kyriakides and Bornstein, 2003; Nam et al., 2004). Formation of this avascular and acellular fibrous capsule can be deleterious to the function of a tissue engineered construct. The implant cannot further vascularize and the cells within the implant undergo ischemic death (Anderson, 2006). Additionally, PVA sponges as implant scaffolds require pro-angiogenic factors (Yamamoto et al., 2000) or implantation in the highly vascular mesenteric space of a larger animal model to ensure vascularization (Takeda et al., 1995). This limits the use of PVA sponges for long term implantation and cell encapsulation.
Our group has recently reported that unmodified PEGDA superporous hydrogels (SPH) allow human mesenchymal stem cell loading, survival for up to a month and stimuli-induced differentiation (Keskar et al.). PEGDA SPHs are highly permeable due to its high water content that which may hold advantages over hydrophobic polymer-based scaffolds. With these observations, we hypothesized that the highly interconnected porous architecture would be favorable for cellular infiltration and angiogenesis in vivo. We further hypothesized that the limited fibrotic response to nonporous PEGDA-based hydrogels would be similar when the porous network was implanted.
PEGDA (MW 3400 g/mol) SPH scaffolds were prepared and characterized as previously reported (Keskar et al.). Briefly, PEGDA solution, foam stabilizer (Pluronic™ F127, double distilled water, the initiator pair, N,N,N’,N’- tetramethylethylene-diamine (TEMED) and ammonium persulfate, were added sequentially to a glass vial. Saturated citric acid solution was used for pH adjustment. The precursor solution was mixed and heated gently to 37°C for approximately 2 minutes. Sodium bicarbonate, 200 mg, was added with constant stirring to evenly distribute the salt and evolving gas. The SPHs were then removed from the vial and allowed to swell in double distilled water to remove traces of unpolymerized monomers and salt before dehydrating in 80% ethanol followed by overnight dehydration in absolute ethanol. The hydrogels were then dried in a food dessicator and stored in an airtight container for further use. To make the non-porous PEGDA hydrogels (NPH), sodium bicarbonate, 200 mg, was replaced with sodium hydroxide solution. The precursor solution was pipetted into 96 well plates with each well containing an equivalent volume to the size of the total volume of the porous hydrogel. Polymerization was allowed to proceed for half an hour. The NPHs were then rinsed with double distilled water to remove traces of unpolymerized monomers, dried and stored in an airtight container for further use. Scanning electron microscopy of the interior surface of the dehydrated SPHs revealed interconnected pores ranging from 100 μm to 600 μm with an average pore size of 250 ± 94 μm (Fig. 1a). The hydrated SPH had a larger pore diameter and broader distribution in pore diameter, 395 ± 107 μm, as estimated from bright field images (Fig. 1b).
To investigate the potential of the SPHs for in vivo angiogenesis, the SPHs were implanted in the dorsal skin fold of SCID mice (Fox Chase SCID, Charles River Laboratories). SCID mice were chosen as the model for vascularization as part of a larger experimental design involving human cells incorporated within the SPH where the tissue growth and vascularization within the SPH will be evaluated. All animal experiments were approved by the institutional animal care committee at the University of Illinois at Chicago. Experimental design for the study was developed using power analysis of published data of porous polymer implants (Arinzeh et al., 2005; Hidetsugu et al., 2007). Briefly, the mice were divided into 2 groups of 7 mice each. The mice were anesthetized with intraperitoneal injection of 100 mg/kg ketamine and 5 mg/kg xylazine. Subcutaneous pockets were opened in the back of the mice using a blunt probe. Two hydrogels, either SPH or NPH, were inserted within either side of the pocket. The incision was closed and the animals were monitored for four weeks. At the end of four weeks, the mice were sacrificed by an overdose of carbon dioxide followed by cervical dislocation. The hydrogels were removed en bloc. Fixed samples were placed in paraffin, sectioned and processed further for histological evaluation of cellular infiltration and vascular ingrowths.
Upon implantation, the hydrogels could be palpated easily. Daily monitoring did not reveal any weight loss or any apparent signs of toxicity like inflammation or reddening of skin in the test animals. At the end of four weeks, the hydrogels could be seen attached to the inside of the dorsal skin. Gross appearance revealed the red, vascularized superporous hydrogels (Fig. 2a) and the pale yellow avascular non-porous hydrogels (Fig. 2b). Hematoxylin and eosin (H&E) staining of the SPH sections revealed host cell infiltration throughout the scaffold (Fig. 3a) which was absent in the nonporous hydrogels. Thus, the porous architecture of the SPH seemed to provide a favorable environment for micro-vessel formation. While it has previously been demonstrated that the rate and the depth of vascular ingrowths within scaffolds are influenced by the presence of pores (Gerecht et al., 2007) the number of vascular ingrowths are limited by the interconnectivity of the pores within the scaffolds (Druecke et al., 2004; Landers et al., 2002).
To further elucidate the presence of host cells, paraffin embedded sections were de-paraffinized in xylene, hydrated with serial concentrations of ethanol and stained with Hoechst 33258 nuclear stain (Latt and Stetten, 1976) (Fig. 3b). A fibrotic, vascularized capsular layer surrounding the SPH was formed (Fig. 4). Extracted SPH implants showed the presence of blood cells and vascular structures. A representative image taken from the center of the SPH section revealed the presence of these ingrowths to the core of the SPH (Fig. 5a). Lumen-like structures filled with blood cells (Fig. 5b) confirmed the presence of a functional microvasculature.
To further confirm the presence of microvasculature, CD34, an early stage endothelial cell marker (Civin et al., 1990), and alpha smooth muscle actin (α-SMA), a vascular smooth muscle cell marker (Van Gieson et al., 2003), were examined. Briefly, the de-paraffinized and hydrated sections were blocked with 1% bovine serum albumin in phosphate buffered saline (pH 7.4) for 30 minutes. The sections were then incubated with primary rat polyclonal antibodies against CD34 and mouse polyclonal antibodies against smooth muscle α-actin (Santa Cruz Biotechnology) for 2 hrs followed by incubation with FITC conjugated goat anti rat and goat anti mouse (Molecular Probes) respectively for 30 mins. Hoechst 33258 was used as the nuclear stain. CD34 positive endothelial cells were localized throughout the SPHs (Fig. 6a). In addition, vessel lining musculature (Owens, 1995) was observed in the form of α-SMA positive cells, associated with the vascular growths (Fig. 6b). These results suggest that the ingrowths observed within the implanted acellular SPHs are neo-vasculature.
Tissues generated in vitro or in vivo must survive and function in vivo However, in vivo tissue viability and function is limited by the ability of the scaffold to support functional microvasculature and overcome subsequent transport limitations. The unmodified PEGDA SPHs utilized in this study were able to support initial in vivo vascularization upon implantation in SCID mice. The presence of a highly interconnected porous architecture within the PEGDA SPH provided an open network which facilitated cellular infiltration from the host and vascular ingrowth throughout the scaffold; PEGDA scaffolds with discontinuous pores should, however, be employed in future studies to elucidate the importance of pore interconnectivity and the rate of in vivo vascularization. In summary, we present a platform scaffold technology that should be further examined for tissue engineering applications. The approach of in vivo implantation of the acellular SPH or seeding cells within the SPH scaffolds before in vivo implantation, could both benefit from rapid thorough vascularization. Future studies will be directed at further understanding the dynamics of vascular infiltration and the stability of the newly formed vasculature.
This investigation was conducted in a facility constructed with support from Research Facilities Improvement Program Grant C06 RR15482 from the National Center for Research Resources, NIH. The authors also thank Dr. Howard Greisler for insightful comments and suggestions.