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Hierarchical scaffold design for mesenchymal stem cell-based gene therapy of hemophilia B
Biomaterials 32 (2011) 295e305
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Hierarchical scaffold design for mesenchymal stem cell-based gene therapy of hemophilia B Daniel L. Coutu a, Jessica Cuerquis a, Rouwayda El Ayoubi b, Kathy-Ann Forner a, Ranjan Roy c, Moïra François a, May Griffith d, David Lillicrap e, Azizeh-Mitra Yousefi b, Mark D. Blostein f, Jacques Galipeau f, * a
Lady Davis Institute for Medical Research, McGill University, Montreal, Canada Industrial Material Institute, National Research Council of Canada, Boucherville, Canada Chemical Engineering Department, McGill University, Montreal, Canada d Eye Institute, Ottawa Health Research Institute, Ottawa, Canada e Department of Pathology and Molecular Medicine, Queen’s University, Kingston, Canada f Division of Hematology, Jewish General Hospital, McGill University, Montreal, Canada b c
a r t i c l e i n f o
a b s t r a c t
Article history: Received 18 August 2010 Accepted 29 August 2010 Available online 23 September 2010
Gene therapy for hemophilia B and other hereditary plasma protein deficiencies showed great promise in pre-clinical and early clinical trials. However, safety concerns about in vivo delivery of viral vectors and poor post-transplant survival of ex vivo modified cells remain key hurdles for clinical translation of gene therapy. We here describe a 3D scaffold system based on porous hydroxyapatiteePLGA composites coated with biomineralized collagen 1. When combined with autologous gene-engineered factor IX (hFIX) positive mesenchymal stem cells (MSCs) and implanted in hemophilic mice, these scaffolds supported long-term engraftment and systemic protein delivery by MSCs in vivo. Optimization of the scaffolds at the macro-, micro- and nanoscales provided efficient cell delivery capacity, MSC self-renewal and osteogenesis respectively, concurrent with sustained delivery of hFIX. In conclusion, the use of geneenhanced MSC-seeded scaffolds may be of practical use for treatment of hemophilia B and other plasma protein deficiencies. Ó 2010 Elsevier Ltd. All rights reserved.
Keywords: Mesenchymal stem cells Gene therapy Hemophilia Biomineralization Self-renewal Osteogenesis
1. Introduction Gene therapy for inherited life-threatening or severely debilitating diseases has shown promises in pre-clinical and early clinical trials [1e3]. Its clinical translation has however been limited by: 1) the poor safety profile of viral vectors used for in vivo gene transfer, and 2) poor survival of ex vivo modified cells upon transplantation. Although the use of ex vivo gene-engineered cells increases the safety of gene transfer and new gene transfer technologies are safer than earlier viral vectors (zinc-finger nucleases for instance), cell survival post-transplantation remains a challenge. Gene therapy for inherited protein deficiencies such as hemophilia A, B and some lysosomal storage disorders require long-term systemic protein delivery with no tissue-specificity [4]. In essence, any autologous cell type available in sufficient numbers can be used * Corresponding author. Department of Hematology and Medical Oncology, Pediatrics & Medicine, Emory University, 1365 Clifton Road, Clinic B, Suite 5100 e 5117, Atlanta, GA 30322, USA. Tel.: þ1 404 778 1779. E-mail address: [email protected] (J. Galipeau). 0142-9612/$ e see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2010.08.094
to deliver the transgene product ectopically and systemically. Our laboratory has used mesenchymal stem cells (MSCs) to deliver various therapeutic proteins in animal models [5e9]. These cells are attractive candidates for gene and cell therapy, tissue engineering and regenerative medicine, owing to their availability from autologous sources, their intrinsic mesodermal plasticity and their pro-angiogenic, pro-regenerative and immuno-modulatory properties [10e16]. Over the years, we have used a wide variety of natural polymers (Matrigel, collagen and hyaluronic acid) to deliver MSCs in vivo [5e9]. We always experienced cytotoxicity issues when using chemically cross-linked polymers (not shown) and fast resorption rate/loss of protein production using enzymatically cross-linked polymers (Supplemental Fig. 1). This was consistent with data from other groups using a similar strategy, who typically observed systemic protein delivery for only 5e10 days [17e21]. We concluded that these biomaterials were too short-lived in vivo and lacked the adhesion cues necessary to MSCs. Moreover, these transplantation methods being essentially isotropic, they lacked precise control over cell capacity and fate. These findings suggest
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that clinical translation of gene therapy for monogenic diseases requires a multidisciplinary feat of an unforeseen complexity. The ideal biomaterial should have sufficient cell capacity to achieve therapeutic benefit and sustain long-term cell survival, self-renewal, proliferation and differentiation. It should also incorporate biochemical and spatial cues mimicking complex tissue composition, architecture and mechanical modulus. This requires a hierarchical design of the scaffold from the nanoscale to the macroscale, where cell fate decisions are controlled at the nanoand micro-scale levels, whereas the micro- to macroscale organization levels influence cell capacity, mass transport, vascularization and mechanical strength [4,22e24]. Biomaterials should also: 1) be made of clinically-approved materials and degrade into non-toxic waste products; 2) be easily and reproducibly processed, at low cost and using commercially available technologies; and 3) be designed for off-the-shelf use (with a long shelf-life). Fig. 1 illustrates the approach used in the present study to develop a nanopatterned macroporous scaffold custom-designed
to support MSC survival, self-renewal and coagulation factor IX (FIX) delivery in hemophilic mice. First, MSCs isolation and culture conditions were optimized to ensure maintenance of multilineage potential and culture expansion (Fig. 1a). After standard gene transfer, protein biochemistry was used to measure hFIX functionality. The main focus of the current study (illustrated in Fig. 1b and c) relates to the material science and engineering aspect of the platform and the translational application in a suitable animal model (scaffold production, tissue engineering, and blood physiology post-transplantation). Briefly, we used rapid prototyping to create polylactic-co-glycolic acid (PLGA) composite scaffolds with a simple 0 /90 strand architecture maximizing cell capacity, mass transport and tissue ingrowth. We then identified calciumephosphate ceramic (e.g. hydroxyapatite, HA) as a potent biomaterial promoting MSC survival post-transplantation. In composite HAePLGA 3D scaffolds, optimization of HA crystal size and concentration as well as modification of the microtopology and biochemical surface properties with biomineralized
Fig. 1. Multidisciplinary approach to gene therapy of hemophilia B. a) Stem cell biology and engineering: isolated MSCs are characterized by their multipotent lineage potential and expression of cell surface markers. Following hFIX-retrovirus cloning, virus production and MSCs transduction, maintenance of MSC properties is confirmed and hFIX production and functionality is quantified. Right panel: crystal structure of FIX Gla domain compexed with Ca2þ is shown (PDB ID: 1J35). b) Material science and engineering: biomaterial scaffolds are engineered from the nanoscale to the macroscale to ensure appropriate MSC differentiation, self-renewal and hFIX production in vitro and in vivo. Left panel: SEM image of an HA crystal embedded in PLGA polymer; overlays show chemical structures of HA and PLGA. Scale bar ¼ 1 mm. Middle panel: SEM image of a biomineralized collagen fiber and coating on HAePLGA scaffolds; overlay shows FTIR spectrum confirming the presence of biomineralized collagen. Scale bar ¼ 50 mm. Right panel: 3D architecture and macroporosity of scaffolds, as used in this study, measured by confocal microscopy. c) Translational study: scaffolds are prepared using rapid prototyping, sterilized and coated with biomineralized collagen I. The scaffolds are then seeded with hFIX þ MSCs which are allowed to adhere, proliferate and infiltrate the scaffold. The scaffolds are transplanted by microsurgery in R333Q hemophilia B mice and hFIX activity measured weekly using an aPTT assay specially designed to detect hFIX activity in mouse plasma. Hemophilic phenotype is monitored over time and MSC fate and persistence is assessed. Some images were adapted and modified with permission from Coutu, D. L. et al. J. Cell Biochem. 108, 537e546 (2009) [4].
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collagen was necessary to obtain long-term MSC engraftment, selfrenewal and differentiation, along with sustained hFIX delivery in hemophilic mice. We thus present a platform for gene therapy of inherited or acquired protein deficiencies using a combination of autologous cells and clinically available biomaterials. The simple fabrication process, ease of use and long shelf-life of our scaffolds make them readily translatable to the clinic and useful for clinical and pre-clinical applications. Our study also provides insightful data on the behavior, fate and requirements of MSCs on biomaterials in vivo. 2. Materials and methods 2.1. Generation of hFIX retroviral vector plasmid and virus The cDNA of human FIX from pCMVFIX (ATCC) was excised by restriction digest and inserted into the multiple cloning site of a retroviral vector pIRES-GFP. This vector was then digested to remove the IRES-GFP. Sequencing of the plasmid by Bio S&T Inc. (Montreal, Quebec) confirmed correct orientation and sequence. For virus production, hFIX retroviral construct was co-transfected with pVSV-G plasmid into 293GP2 packaging cells (ATCC) in the presence of Polyfect Reagent (Qiagen). Virus was collected for 3 days, filtered and frozen until used.
2.2. Cell culture MSCs from 4 to 8 weeks old R333Q male mice were isolated by whole bone crush and plastic adherence, cultured in DMEM with 10% fetal bovine serum (Wisent) and antibiotics and allowed to expand for 2e4 passages to obtain a homogenous population before retrovector transduction. Multipotentiality was determined as described [25,26]. Vitamin K1 (7 mg/mL) was added to stimulate g-carboxylation. For retroviral transduction, medium was replaced with virus containing conditioned medium from packaging cells diluted 1:1 in growth medium, in the presence of polybrene (2 mg/mL).
2.3. Animals All animal work was approved by the Animal Research Ethics Committee of the Lady Davis Institute for Medical Research. R333Q hemophilia B mice [27] were a kind gift from Dr Darrel Stafford. Only males were used for transplantation. For collagen gel injection, mice were injected subcutaneously on the left flank. For surgical implantation, mice were anesthetized using isoflurane, a subcutaneous pocket was created by blunt dissection on the left flank where the implants were placed and the wound closed with surgical staples. When necessary, animals were sacrificed by CO2 inhalation and implants retrieved and processed for histology. For blood sampling, uncoated capillary tubes were used to draw blood from the saphenous vein. Samples were immediately placed in 3.9% sodium citrate (blood: citrate ratio of 10:1) then centrifuged. Citrated plasma was stored frozen until tested.
2.4. Flow cytometry Cells were harvested, washed with PBS and fixed with paraformaldehyde. Cells were then incubated with the following primary antibodies: anti-mouse CD45, CD31, CD44, CD73, CD105, CD34 or isotypic controls. Appropriate fluorescent secondary antibodies were then applied and analysis performed on a FACS Calibur instrument.
2.5. Western blotting hFIX þ MSC were grown up to 80% confluence, washed extensively with PBS and serum free DMEM was added. Conditioned medium (CM) was harvested after 48 h. After 50-fold concentration by ultrafiltration (MWCO 10 kD), samples were separated by SDS-PAGE, transferred on PVDF membranes and probed with anti-hFIX (Haematologic Technologies Inc) or anti-Gla (American Diagnostica) antibodies and appropriate secondary antibody, then detected using ECL.
2.6. hFIX chromogenic assay To activate hFIX, 1 U/mL hFXIa was added to concentrated conditioned medium from wild type or hFIX þ MSC in the presence of 5 mM CaCl2 and incubated 45 min at 37 C. hFIXa activity was determined using the chromogenic substrate 229 (American Diagnostica). Buffer used was 100 mM Tris, 100 mM imidazole, 100 mM NaCl, 0.1% BSA, 5 mM CaCl2 and 33% ethylene glycol. Samples were compared to a standard curve prepared by serial dilution of purified hFIXa (Haematologic Technologies) in growth medium. Activation of hFIX was confirmed by western blot.
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2.7. hFIX-specific ELISA Assay was performed using the anti-hFIX paired antibodies kit (Cedarlane Laboratories) and performed under the manufacturer’s instructions on concentrated serum free conditioned medium. A standard curve was prepared with purified hFIX (Haematologic Technologies Inc.). 2.8. Scaffolds fabrication and cell seeding Scaffolds were prepared as described [28] with modifications, using 8515DLHighIV PLGA (Lakeshore Biomaterials) with or without the incorporation of hydroxyapatite ceramic particles (Berkeley Advanced Biomaterials). The plotting material was prepared by dissolving PLGA in methyl-ethyl-ketone and mixing with hydroxyapatite particles when necessary. Optimal concentration of the polymer in solvent was determined based on viscosity constraints of the 3D plotter and optimal syringe deposition. Three-dimensional (3D) scaffolds were fabricated by rapid prototyping using a Bioplotter (Envisiontec, Gladbeck, Germany) and computer-assisted design (CAD). Briefly, the plotting material was transferred to the plotting cartridge and was dispensed layer by layer, forming a 0 /90 strand structure. Different topologies were fabricated (strand size and distance between strands) as specified in the text. Plotting needles had 250 and 400 mm inner diameter for PLGA and HAePLGA scaffolds, respectively. Bricks of 20 20 1.1 mm were fabricated. Discs of 6 mm diameter were punched out and vacuum-dried to allow solvent evaporation. Scaffolds were sterilized in 70% ethanol, washed in PBS and coated with 1% collagen type I (BD Bioscience) or biomineralized collagen I [29,30]. For cell seeding, 30 mL of cell suspension (approximately 250 000 cells) was loaded onto the scaffolds by capillarity. After 2 h incubation in a humidified incubator, growth medium was added. The cells cultured for 2 weeks prior to their implantation with frequent medium changes (or additionally for 3 weeks in differentiation medium). 2.9. Scanning electron microscopy SEM was done using a Hitachi S-4700 Scanning Electron Microscope. Scaffolds were thoroughly washed with dH2O and ethanol, then lyophilized and coated using gold/palladium. For cell-seeded scaffolds, a fixation step was included before washing (2.5% glutaraldehyde in 0.1 M sodium cacodylate at pH 7.4 for 2 h). 2.10. hFIX-specific aPTT assay on mouse plasma We used a hFIX-specific aPTT assay modified from Arruda et al. [31]. Briefly, a standard curve was prepared by serially diluting hFIX in citrated plasma from R333Q mice. Samples and standards were diluted 1:10 in imidazole buffer. Samples, FIX-deficient pooled human plasma, aPTT activator reagent and CaCl2 were mixed in equal volumes, calcium being added after 3 min incubation at 37 C. Clotting time was measured using a KC4 coagulometer. 2.11. Histology/immunohistochemistry Implants were retrieved, fixed with formalin and processed for regular histology. Implants were decalcified using 10% EDTA between fixation and processing. Alternatively for undecalcified histology, formalin fixed implants were processed for methylmethacrylate embedding. For immunohistochemistry, we used antigen retrieval solution, avidin/biotin blocking kit and DAB reagent from Vector. Anti-mouse osteocalcin (AbDSerotec), appropriate biotinylated secondary antibodies and streptavidin-HRP (Jackson ImmunoResearch) were used. 2.12. Microscopy Photomicrographs were taken using a Leica DM LB2 microscope equipped with a Leica DFC480 camera and acquired using the Leica Application Suite software. Alternatively, mounted slides were scanned using a NanoZoomer (Hamamatsu) and analyzed using the NanoZoomer Digital Pathology (NDP View) software. 2.13. Confocal microscopy Paraffin sections were stained as described [32] using antigen retrieval solution from Vector, anti-hFIX and anti-Ki67 (Dako) antibodies, appropriate isotypic controls (not shown) and Alexa Fluor 488 or 647 conjugated secondary antibodies and DAPI (Molecular Probes). Confocal imaging was done on a WaveFX/Leica spinning disk confocal microscope. Image acquisition, 3D reconstruction and deconvolution were done using Volocity or Imaris softwares, when necessary. 2.14. Collagen gel preparation Rat tail collagen type I (10 mg/mL) was from BD Biosciences. 1 M199 buffer was supplemented with 0.75% NaHCO2 and 0.01 M HEPES. The injections consisted of 75% collagen gel (with or without 40 mg/mL HA ceramic particles, see below) and 25%
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cell suspension with the following final concentrations: 4.9 mg/mL collagen, 1 mM DTT and 0.24 U/mL tissue non-specific transglutaminase (Sigma). PBS diluted collagen was mixed with complete M199 buffer (60:40 ratio). The pH was adjusted using 0.1 M NaOH. Cells, DTT and transglutaminase were then added. Kept liquid on ice, the gel is loaded into a syringe and injected subcutaneously. 2.15. MSCs implantation on HA particles Was performed as described [33,34] with modifications. MSC suspended in growth medium were incubated with 40 mg HA/106 cells (particle size 250 mm diameter, Plasma Biotal) with slow rotation at 37 C for 2 h. The cell/ceramic suspension was centrifuged and mixed in the collagen I gel described above prior to implantation. Up to 107 cells were transplanted in each animal. 2.16. Fourier transform infra red spectroscopy (FTIR)-microscopy Samples were washed with dH2O and ethanol and then lyophilized. Measurements were done using a Bruker Tensor 27 equipped with a Hyperion 2000 microscope on 35 different points. 2.17. Statistical analyses For aPTT assays, dot plots showing individual mice data overlaid with mean SEM is shown. Two-way ANOVA were performed using GraphPad Prism5. For histograms, data from one of at least three experiments is shown. We did not pool data from different experiments because hFIX production levels varied slightly from one transduced MSC lot to another. All tests were performed at least in triplicates.
3. Results 3.1. MSC characterization, gene-engineering and functionality We used the murine R333Q model of hemophilia B because these mice do not develop inhibitory antibodies against hFIX after gene or protein therapy [35]. These FIX knockout mice express, under the murine FIX promoter, an inactive hFIX carrying the R333Q mutation frequently observed in humans. They are thus human cross-reactive material positive but have a hemophilic phenotype, recapitulating the human condition. As an initial step, MSC were isolated from R333Q mice and transduced with a hFIXretrovirus. They were shown to retain in vitro multilineage differentiation potential into adipocytes, osteocytes and chondrocytes (Fig. 2a). They also presented a typical C57Bl/6 MSC immunophenotype (including CD34) for cell surface markers (Fig. 2b). In our hands, CD105 is rapidly lost with culture expansion of murine MSC. hFIX secretion was quantified by ELISA and found to be maintained throughout multilineage differentiation (Fig. 2c). MSC-derived hFIX functionality was assessed by western blot analysis, confirming g-carboxylation of glutamic acid residues (Glu) to g-carboxyglutamic acid (Gla; Fig. 2d). Functionality was also demonstrated using a chromogenic assay, where MSC-derived hFIX was activated by hFXIa and then quantified using the chromogenic substrate 229. Western blot analysis confirmed the activation (cleavage) of hFIX to hFIXa which proteolytically processed 229, releasing free pNA (Fig. 2e). 3.2. Macro- to micro-scale optimization of 3D porous scaffolds We previously demonstrated that synthetic polymers (such as PLLA/PLGA)-based 3D porous scaffolds created using a combination of rapid prototyping (RP, or solid free forming) and porogen leaching can be optimized for maximal cell capacity as well as mass and oxygen transport in vitro [28]. While RP allows fine control over structure, pore interconnectivity and cell seeding efficiency, microporosity influences cell adhesion, behavior and mass transport. Fig. 1b (right panel) shows a disc-shaped scaffold with a simple 0 /90 architecture similar to those used previously and here. Our initial data demonstrated that these scaffolds supported MSCs adhesion, proliferation and protein production in vitro.
To further optimize tissue ingrowth by MSCs on scaffolds, we here modified scaffold macroporosity by varying strand diameter (150e300 mm), distance between strands (150e300 mm) and layer number (0.6e1.3 mm total thickness). We also varied microporosity using porogen leaching, where NaCl (0e30%) was incorporated into the polymer and then leached in water. Scanning electron microscopy (SEM) demonstrated the drastic effect of micro- and macroporosity modification on MSC ingrowth. Scaffold topologies of 200 250 (strand diameter and distance between strands respectively, in micrometers) and 15e30% porogen provided the best MSC proliferation and tissue ingrowth after two weeks in culture (Fig. 3a). This was surprising since we had previously demonstrated that a scaffold topology of 150 150 with 30% porogen provided the highest total surface area and maximal effective surface area. However, the current results indicate that larger pores facilitate MSC invasion of pore space in 3D scaffolds. We confirmed by ELISA that MSCs seeded on PLGA scaffolds maintained hFIX production in vitro throughout mesodermal differentiation (Fig. 3b). We next implanted these optimized constructs in hemophilic mice. We followed systemic hFIX delivery by weekly aPTT assays and retrieved the scaffolds after 8 weeks for histological analysis. Alcian blue stain (for cartilage-specific proteoglycans) revealed that the implants contained mostly fibrous tissue, with evidence of mesenchymal condensations staining weakly with alcian blue and extensive vascularisation (Fig. 3c). Toluidine blue/von Kossa staining (for osteoid matrix and mineralized phosphate, respectively) showed a complete absence of mineralization and osteoid deposition (Supplemental Fig. 3, left panel). Systemic hFIX activity reached therapeutic levels for a short two week period before returning to baseline. These results identified the optimal 3D architecture and macroporosity of scaffolds for in vitro MSC proliferation and tissue ingrowth. However, the PLGA substrate clearly did not allow long-term in vivo MSC persistence (self-renewal) and longterm systemic hFIX delivery at therapeutic levels. 3.3. Calciumephosphate ceramics for MSC delivery Calciumephosphate ceramics (such as hydroxyapatite, HA) have a long history of clinical use and have been demonstrated by many groups to sustain long-term engraftment of MSC in vivo [33,34,36,37]. However, ceramics are not easily processed into complex 3D structures. We first tested the ability of isotropically transplanted 175e250 mm diameter HA ceramic particles to promote MSC self-renewal and protein delivery in vivo. We used a transglutaminase cross-linked collagen I hydrogel to hold the particles together (see Supplemental Fig. 1). Fig. 4a illustrates the sequence of bone development by murine MSC on HA. At 2 weeks post-implantation, the implants are very cellular (with non-hematopoietic, presumably donor-derived cells) with evidence of de novo ECM deposition. At week 4, the implants show a decreased cellularity and predominance of fibrous tissue. At week 6, an increase in cell content is observed, coincident with vascular invasion. Within 7e8 weeks, bone tissue is formed within the implant, with an outer periosteum-like capsule that eventually mineralizes and contains hFIX þ cells (not shown). A hematopoietic marrow-like cavity is observed and the implants are densely irrigated by large, sinusoidlike blood vessels. They also contain extravascular erythrocytes suggestive of erythropoiesis (Supplemental Fig. 2). At 10 weeks post-implantation, the presence of hFIX þ MSC closely attached to the carrier particles can still be detected by immnohistochemistry (Fig. 4b). HA-seeded MSC produced detectable systemic hFIX for up to 12 weeks, although the levels remained below the 2% therapeutic threshold (Fig. 4c). The low total surface area of the ceramic particles (low cell capacity) likely explains these results. Moreover, 40 mg of particles is required for every 106 cells [38], making scaling up
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Fig. 2. MSCs characterization, gene modification and hFIX production. a) MSCs from R333Q hemophilic mice displayed typical mesodermal plasticity and could be differentiated into osteocytes, adipocytes and chondrocytes (left to right: alizarin red S, oil red O and alcian blue stains) after retroviral transduction. b) Flow cytometry immunophenotyping of R333Q MSCs reveals expression of murine MSC markers CD34, CD44 and CD73 with the absence of hematopoietic/endothelial markers CD45 and CD31. c) MSCs maintained hFIX secretion after their differentiation into mesodermal lineages as measured using a hFIX-specific ELISA. (UD: undifferentiated). d) Western blot analysis of wild type (WT) or hFIX þ MSCs conditioned medium demonstrates the presence of g-carboxyglutamic acid on MSC-derived hFIX. e) Upper panel: MSC-derived hFIX functionality was measured using a hFIX-specific in vitro chromogenic assay. hFIX in MSC conditioned medium was activated with hFXIa and its activity measured with the hFIX-specific chromogenic substrate 229. Lower panel: hFIX activation was confirmed by western blot. (FL: full length, HC: heavy chain).
impractical due to the large size of the implants, which also limits mass transport. These results show that HA particles promote long-term engraftment and hFIX delivery by MSCs and that osteogenesis is
required for both processes. However, isotropically transplanted particles lack the 3D architecture maximizing cell capacity and hFIX delivery in vivo. We thus used rapid prototyping to create composite HAePLGA macroporous 3D scaffolds.
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Fig. 3. Macro- and microporosity optimization of 3D porous scaffolds. a) PLGA scaffolds were created using solid free forming processing. Macrotopology was modified by varying strand diameter and distance between strands. Microtopology was modified using porogen leaching. Scaffolds seeding and tissue ingrowth by MSCs in vitro was assessed by SEM two weeks after seeding. Images demonstrate that both macro- and micro-porosity influence the capacity of MSCs to adhere, proliferate and invade scaffold pores. Best tissue ingrowth was observed with the 200 250 (30%) topology, where 200 is strand diameter, 250 is distance between strands, and 30% the porogen concentration. Scale bar ¼ 100 mm b) MSCs maintained hFIX secretion after their differentiation into mesodermal lineages on porous PLGA scaffolds, as measured using a hFIX-specific ELISA. c) Histological analysis of retrieved implants 5 weeks post-implantation reveals poor chondrocytic differentiation (alcian blue stain shown). The implants contained few mesenchymal condensations (arrowheads) and large blood vessels (arrow). Scale bar ¼ 100 mm. C: carrier. d) MSCs were seeded on the best scaffold topology and differentiated into osteoblast, adipocytes or chondrocytes before transplantation. hFIX delivery was monitored by weekly aPTT assays. Only the chondrocytic group (shown) reached hFIX levels above therapeutic threshold, very transiently.
3.4. Nano- to micro-scale optimization of 3D scaffolds To create HAePLGA composite scaffolds, we first needed to optimize the composition, concentrations, and formulation of both components within the scaffolds. HA physico-chemical properties (crystal size and perfection for instance) influence its osteoconductivity [39]. Moreover, crystal size also affected our capacity to process HAePLGA porous scaffolds using rapid prototyping. For instance, 200 mm particles were poorly miscible in PLGA and were not tested further. To avoid delamination problems, we also had to increase both strand diameter and distance between strands to 300 mm, reducing the effective total surface area by about half (not shown). Nevertheless, we successfully created composite scaffolds using two types of HA, nano-HA and micro-HA (100 nm and 5 mm particle diameter, respectively). Our findings are summarized in Supplemental Figs. 3 and 4. Briefly, we found that nano-HA was more osteoconductive than micro-HA at the same concentration (40%). However, both constructs did not support long-term hFIX delivery by MSCs in vivo. At higher concentration, occlusive aggregation of the nano-HA particles occurred within the printing
syringe whereas we were able to reach 65% micro-HA concentration. Our data shows that micro-HA (65%) is more effective than nano-HA (40%) in supporting osteogenesis and hFIX delivery by MSCs. Cell adhesion to biomaterials such as PLGA and HA is thought to require the presence of binding proteins to bridge between cells and material. Coating with serum, collagens or other extracellular matrix proteins is typically used. Because we observed that osteogenesis was necessary for long-term MSC engraftment and hFIX delivery, we sought to direct MSC adhesion and fate decision by modifying the biochemical surface properties and microtopology of the scaffolds. More specifically, we directly compared two different coatings: collagen type I (the most abundant collagen in bones), and biomineralized collagen I [29,30]. To assess the coating efficiency and modification of microtopology we used a combination of SEM and Fourier Transform Infra Red (FTIR)-spectroscopy and microscopy (Fig. 5 and Supplemental Fig. 5). Uncoated PLGA scaffolds show a very smooth surface texture under SEM whereas HAePLGA composite scaffolds presented a rougher surface (Fig. 5a), with evidence of HA particles exposed
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Fig. 4. MSCs persistence, fate and hFIX production on HA/TCP ceramic particles. hFIX þ MSCs were seeded on 200e500 mm HA/TCP ceramic particles and injected subcutaneously in R333Q mice. Bone formation was assessed by histological analysis of retrieved implants and hFIX systemic delivery by weekly aPTT assays. a) Sequence of bone development by MSCs on HA/TCP in vivo. Donor-derived MSCs are highly present in the implants at 2 weeks post-transplant but are gradually replaced with host-derived hematopoietic marrow at 6 weeks, concomitant with development of vasculature (arrowheads). At 7 weeks, the bone-like implant contains a readily identifiable marrow cavity (*). b) hFIX þ MSCs closely attached to the ceramic particles are still detectable by IHC after 10 weeks. c) Weekly aPTT assays reveal that although hFIX persists in the plasma of the transplanted animals for up to 12 weeks, the levels produced are below the 2% normal activity (100 ng/mL) therapeutic threshold (dashed line). Scale bars ¼ 100 mm.
on the surface (presence of calcium was confirmed using X-ray fluorescence, not shown). Upon coating with collagen or biomineralized collagen, the microtopology and texture of the scaffolds were greatly modified (Fig. 5b). In the former, we can observe large collagen fibrils and bundles of various sizes covering the surface. In the later, both the protein-rich collagen coating and large fibrils are completely mineralized. Higher magnification SEM images (Fig. 5c) demonstrate the presence of free-floating HA crystals (upper panel) and collagen fibrils showing evidence of clubbing and bulging, reflecting nucleation of spherulitic HA within gaps of the fibrils [40] (lower panel). We performed FTIR spectroscopy to confirm the composition of the coating and to measure changes in surface chemistry (Fig. 5d). The characteristic n3 vibration mode of phosphate within HA (1075 cm1) was clearly visible in all HAePLGA scaffolds. The sharp amine band near 1750 cm1 and increased hydroxyl band confirmed efficient collagen coating. The n1 phosphate vibration mode (960 cm1) was further increased by biomineralization, suggesting a higher crystalline HA at the strand surface. Other dominant bands between 800 and 1110 cm1 likely reflect various amorphous and protonated phosphates [41]. FTIR-microscopy
confirmed efficient coating of the whole scaffold on at least three layers (Supplemental Fig. 5). 3.5. MSC self-renewal, fate and hFIX delivery on fully optimized 3D scaffolds As with previous experiments, we tested our optimized scaffolds by seeding them with hFIX þ MSCs and transplanting them subcutaneously in R333Q hemophilic mice. We monitored plasmatic hFIX levels by weekly aPTT assays and MSC fate by histology and immunofluorescence confocal microscopy on retrieved implants (Fig. 6). Toluidine blue/von Kossa staining of implants retrieved at 15 weeks post-transplantation demonstrates the extent of osteogenesis in optimized scaffolds. We observed increased osteogenesis, scaffold remodelling, mineralization and cellular invasion in the biomineralized collagen-coated scaffolds than with any other scaffold used (compare Fig. 6a with Supplemental Figs. 3, 4 and 6). To assess MSC self-renewal within the scaffolds, we used immunofluorescence confocal microscopy and were able to demonstrate the persistence of hFIX þ cells co-expressing the proliferation marker Ki67 15 weeks after transplantation (Fig. 6b). These cells were almost
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Fig. 5. Nano- and micro-topology optimization of 3D porous scaffolds. a) SEM images of two 3D porous scaffolds used in this study, show that PLGA scaffolds have a very smooth surface whereas the surface texture of HAePLGA scaffold is rougher (we can also see HA crystals, which appear brighter on the SEM) protruding at the surface of the strands. b) Modification of scaffolds surface properties (chemistry and topology) as assessed by SEM. Scaffolds were coated with collagen type I or biomineralized collagen type I. Left: After coating with collagen, the proteinaceous layer completely covers the scaffold strands and large collagen fibrils are observed. Right: After subsequent biomineralization of the collagen coating, the microtopology is greatly affected and collagen fibrils of all sizes appear mineralized. Boxed area is shown at higher magnification in c). c) High power SEM images of the biomineralized coating of HAePLGA scaffolds reveals the presence of free HA crystals (arrowhead) attached to the thick layer of biomineralized collagen I. The small collagen fibers show bulging and clubbing, suggesting the presence of spherulitic apatite within fibril gaps. d) FTIR spectra of the various 3D scaffolds used in this study confirm coating efficiency and composition. The apatite-specific n3 vibration mode of phosphate (1075 cm1, double arrow) can be seen in all HAePLGA scaffolds. The amine band is clearly visible after collagen I coating. After collagen biomineralization, the n1 vibration mode of phosphate (960 cm1, single arrow) is increased.
exclusively localized closely attached to the scaffold strands. Interestingly, as seen with calciumephosphate ceramic particles we also observed the creation of a hematopoietic marrow-like cavity with the scaffold pores, as evidenced by abundant extravascular erythrocytes (Fig. 6c). The formation of this myelosupportive environment was only observed with these two transplantation vehicles. We were also able to detect long-term hFIX production in hemophilic mice at therapeutic levels in most animals, although we did observe variability in responses to the treatment (Fig. 6d). hFIX levels remained detectable, albeit with reduced levels after 12 weeks, for the duration of the experiment (15 weeks). Interestingly, plasmatic hFIX levels oscillated with a period of 10e20 days. It is
worth mentioning that murine osteoblasts have an estimated halflife of 10e20 days [36,42]. Thus, the oscillating levels of hFIX observed may well represent normal cell turnover and replacement, again suggesting the persistence of self-renewing MSCs within the implants. 4. Discussion In this study, we developed a 3D porous scaffold with a hierarchical design made with clinically-approved biomaterials that, combined with autologous mesenchymal stromal cells, could be used off-the-shelf for personalized cell-based gene therapy of
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Fig. 6. In vivo performance of MSC-seeded, fully optimized 3D porous scaffolds for hFIX delivery in hemophilic mice. Histological and immunofluorescence analysis of retrieved implants at 15 weeks post-transplant. a) von Kossa/toluidine blue stain reveals important remodelling and mineralization (black staining) of the scaffolds (left panel; scale bar ¼ 200 mm). The presence of multinucleated osteoclasts in the mineralized and partially degraded carrier material (c) is shown (*), as well as the deposition of osteoid-like matrix on the scaffold strands (deep blue staining and arrowheads; right panel). b) MSCs persistence, survival and self-renewal within the implants was assessed by immunofluorescence and confocal microscopy. Image shows co-localization of hFIX and the proliferation marker Ki67 in cells closely attached to the scaffold strands (arrowheads). Dashed line represents the tissue/scaffold interface. Scale bar ¼ 50 mm c) Hematoxylin/eosin staining shows the presence of hematopoietic tissue within scaffold pores, as evidenced by cell morphology and the presence of extravascular erythrocytes (arrowheads). Scale bar ¼ 50 mm d) Weekly aPTT assays on the plasma of transplanted animals demonstrates reversal of the hemophilic phenotype of most animals for up to12 weeks, with detectable hFIX persisting for the duration of the experiment (15 weeks).
inherited and acquired protein deficiencies. We used rapid prototyping because it can be performed with commercially available technologies and offers great control over pore size, orientation, interconnectivity and total surface area in the scaffolds. When combined with porogen leaching, it also allows maximal cell capacity and mass transport through the implant [28]. As described by others [43e46], our study confirms that substrate macro- and micro-porosity as well as nano- and micro-texture greatly influences cell behavior in vitro and in vivo. Furthermore, we show that only calciumephosphate ceramics provide the necessary signals promoting long-term MSCs survival and self-renewal. These could be further increased by modification of the surface chemistry and nano-topology of the scaffold, and by modifying particle size and concentration of the ceramic. The complete reversal of the hemophilic phenotype in mice demonstrated in this study strongly suggests that this approach has clinical relevance. The fact that we were able to cure the mice for over 12 weeks is a considerable improvement over previous attempts using similar strategies [17e21]. Although some animals
eventually lost therapeutic hFIX activity over time, additional rounds of transplantation could easily be contemplated. Since R333Q mice are not known to mount immune responses to hFIX in gene and protein therapy, potential explanations for loss of hFIX activity include promoter silencing (we used retroviral gene transfer and LTRs are subject to inactivating methylation in vivo). Furthermore, as seen with hematopoietic stem cells it is not unlikely that subpopulations of MSCs exist that possess either short- or long-term “reconstituting” (self-renewal) potential. These subpopulations have however not yet been described for MSCs. Nevertheless, human MSCs of good quality are much easier to isolate and expand ex vivo than murine MSCs (especially in the C57Bl/6 background) [47,48] suggesting that our strategy would also be successful in humans. Finally, it should be mentioned that FIX is a very abundant protein (with systemic levels of nearly 5 mg/mL). Thus, reaching the relatively low 2% therapeutic threshold still requires 100 ng/mL systemic FIX, a huge protein load. In comparison, therapeutic efficacy for hemophilia A would only require 4 ng/mL systemic factor VIII.
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5. Conclusion In summary, we here present a clinically relevant platform for mesenchymal stem cell-based gene therapy of protein deficiencies such as hemophilia B. Our data illustrates how long-term selfrenewal and differentiation of MSCs can be manipulated through hierarchical scaffold design from the nanoscale to the macroscale. Our use of commercially available technologies and clinically accepted materials combined with autologous cells make our approach easily translatable to the clinic. Our data also provides key insights into the physico-chemical environment necessary for longterm MSCs survival and self-renewal in vivo. Acknowledgements The authors would like to thank Yoon Kow Young and the WEBC Imaging Facility for assistance with confocal microscopy; the McGill Bone Center for MMA embedding and sectioning; Micheline Fortin and the IRIC Histology Platform for assistance with histology and digitalization of slides; Christian Degrandpré for scaffold fabrication; Dr Ian B. Copland and Dr Janet E. Henderson for useful discussions and technical help; the Canadian Stem Cell Network and Fonds de Recherche en Santé du Québec for funding. Appendix. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.biomaterials.2010.08.094. Appendix Figures with essential color discrimination. Figs. 1e4 and 6 in this article are difficult to interpret in black and white. The full color images can be found in the online version, at doi:10.1016/j. biomaterials.2010.08.094. References [1] Edelstein ML, Abedi MR, Wixon J. Gene therapy clinical trials worldwide to 2007ean update. J Gene Med 2007;9:833e42. [2] Chuah MK, Collen D, VandenDriessche T. Preclinical and clinical gene therapy for haemophilia. Haemophilia 2004;10(Suppl. 4):119e25. [3] Murphy SL, High KA. Gene therapy for haemophilia. Br J Haematol 2008;140:479e87. [4] Coutu DL, Yousefi AM, Galipeau J. Three-dimensional porous scaffolds at the crossroads of tissue engineering and cell-based gene therapy. J Cell Biochem 2009;108:537e46. [5] Copland IB, Jolicoeur EM, Gillis MA, Cuerquis J, Eliopoulos N, Annabi B, et al. Coupling erythropoietin secretion to mesenchymal stromal cells enhances their regenerative properties. Cardiovasc Res 2008;79:405e15. [6] Kucic T, Copland IB, Cuerquis J, Coutu DL, Chalifour LE, Gagnon RF, et al. Mesenchymal stromal cells genetically engineered to overexpress IGF-I enhance cell-based gene therapy of renal failure-induced anemia. Am J Physiol Renal Physiol 2008;295:F488e96. [7] Eliopoulos N, Lejeune L, Martineau D, Galipeau J. Human-compatible collagen matrix for prolonged and reversible systemic delivery of erythropoietin in mice from gene-modified marrow stromal cells. Mol Ther 2004;10:741e8. [8] Stagg J, Lejeune L, Paquin A, Galipeau J. Marrow stromal cells for interleukin-2 delivery in cancer immunotherapy. Hum Gene Ther 2004; 15:597e608. [9] Eliopoulos N, Al Khaldi A, Crosato M, Lachapelle K, Galipeau J. A neovascularized organoid derived from retrovirally engineered bone marrow stroma leads to prolonged in vivo systemic delivery of erythropoietin in nonmyeloablated, immunocompetent mice. Gene Ther 2003;10:478e89. [10] Prockop DJ. “Stemness” does not explain the repair of many tissues by mesenchymal stem/multipotent stromal cells (MSCs). Clin Pharmacol Ther 2007;82:241e3. [11] Prockop DJ, Gregory CA, Spees JL. One strategy for cell and gene therapy: harnessing the power of adult stem cells to repair tissues. Proc Natl Acad Sci U S A 2003;100(Suppl. 1):11917e23. [12] Stagg J, Galipeau J. Immune plasticity of bone marrow-derived mesenchymal stromal cells. Handb Exp Pharmacol 2007;180:45e66.
[13] Romieu-Mourez R, Francois M, Boivin MN, Stagg J, Galipeau J. Regulation of MHC class II expression and antigen processing in murine and human mesenchymal stromal cells by IFN-gamma, TGF-beta, and cell density. J Immunol 2007;179:1549e58. [14] Francois M, Romieu-Mourez R, Stock-Martineau S, Boivin MN, Bramson JL, Galipeau J. Mesenchymal stromal cells cross-present soluble exogenous antigens as part of their antigen-presenting cell properties. Blood 2009; 114:2632e8. [15] Romieu-Mourez R, Francois M, Boivin MN, Bouchentouf M, Spaner DE, Galipeau J. Cytokine modulation of TLR expression and activation in mesenchymal stromal cells leads to a proinflammatory phenotype. J Immunol 2009; 182:7963e73. [16] Bianco P, Robey PG, Simmons PJ. Mesenchymal stem cells: revisiting history, concepts, and assays. Cell Stem Cell 2008;2:313e9. [17] Wen J, Xu N, Li A, Bourgeois J, Ofosu FA, Hortelano G. Encapsulated human primary myoblasts deliver functional hFIX in hemophilic mice. 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J Bone Joint Surg Am 2004;86-A:1541e58. [23] Griffith LG, Swartz MA. Capturing complex 3D tissue physiology in vitro. Nat Rev Mol Cell Biol 2006;7:211e24. [24] Griffith LG. Emerging design principles in biomaterials and scaffolds for tissue engineering. Ann N Y Acad Sci 2002;961:83e95. [25] Pittenger MF, Mackay AM, Beck SC, Jaiswal RK, Douglas R, Mosca JD, et al. Multilineage potential of adult human mesenchymal stem cells. Science 1999;284:143e7. [26] Mackay AM, Beck SC, Murphy JM, Barry FP, Chichester CO, Pittenger MF. Chondrogenic differentiation of cultured human mesenchymal stem cells from marrow. Tissue Eng 1998;4:415e28. [27] Jin DY, Zhang TP, Gui T, Stafford DW, Monahan PE. Creation of a mouse expressing defective human factor IX. Blood 2004;104:1733e9. [28] El Ayoubi R, Eliopoulos N, Diraddo R, Galipeau J, Yousefi AM. Design and fabrication of 3D porous scaffolds to facilitate cell-based gene therapy. Tissue Eng Part A 2008;14:1037e48. [29] Chen Y, Mak AF, Wang M, Li J. Composite coating of bonelike apatite particles and collagen fibers on poly L-lactic acid formed through an accelerated biomimetic coprecipitation process. J Biomed Mater Res B Appl Biomater 2006;77:315e22. [30] Kim SS, Park MS, Gwak SJ, Choi CY, Kim BS. Accelerated bonelike apatite growth on porous polymer/ceramic composite scaffolds in vitro. Tissue Eng 2006;12:2997e3006. [31] Arruda VR, Hagstrom JN, Deitch J, Heiman-Patterson T, Camire RM, Chu K, et al. Posttranslational modifications of recombinant myotube-synthesized human factor IX. Blood 2001;97:130e8. [32] Robertson D, Savage K, Reis-Filho JS, Isacke CM. Multiple immunofluorescence labelling of formalin-fixed paraffin-embedded (FFPE) tissue. BMC Cell Biol 2008;9:13. [33] Sacchetti B, Funari A, Michienzi S, Di Cesare S, Piersanti S, Saggio I, et al. Selfrenewing osteoprogenitors in bone marrow sinusoids can organize a hematopoietic microenvironment. Cell 2007;131:324e36. [34] Mankani MH, Kuznetsov SA, Marshall GW, Robey PG. 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Contents lists available at ScienceDirect
Biomaterials journal homepage: www.elsevier.com/locate/biomaterials
Hierarchical scaffold design for mesenchymal stem cell-based gene therapy of hemophilia B Daniel L. Coutu a, Jessica Cuerquis a, Rouwayda El Ayoubi b, Kathy-Ann Forner a, Ranjan Roy c, Moïra François a, May Griffith d, David Lillicrap e, Azizeh-Mitra Yousefi b, Mark D. Blostein f, Jacques Galipeau f, * a
Lady Davis Institute for Medical Research, McGill University, Montreal, Canada Industrial Material Institute, National Research Council of Canada, Boucherville, Canada Chemical Engineering Department, McGill University, Montreal, Canada d Eye Institute, Ottawa Health Research Institute, Ottawa, Canada e Department of Pathology and Molecular Medicine, Queen’s University, Kingston, Canada f Division of Hematology, Jewish General Hospital, McGill University, Montreal, Canada b c
a r t i c l e i n f o
a b s t r a c t
Article history: Received 18 August 2010 Accepted 29 August 2010 Available online 23 September 2010
Gene therapy for hemophilia B and other hereditary plasma protein deficiencies showed great promise in pre-clinical and early clinical trials. However, safety concerns about in vivo delivery of viral vectors and poor post-transplant survival of ex vivo modified cells remain key hurdles for clinical translation of gene therapy. We here describe a 3D scaffold system based on porous hydroxyapatiteePLGA composites coated with biomineralized collagen 1. When combined with autologous gene-engineered factor IX (hFIX) positive mesenchymal stem cells (MSCs) and implanted in hemophilic mice, these scaffolds supported long-term engraftment and systemic protein delivery by MSCs in vivo. Optimization of the scaffolds at the macro-, micro- and nanoscales provided efficient cell delivery capacity, MSC self-renewal and osteogenesis respectively, concurrent with sustained delivery of hFIX. In conclusion, the use of geneenhanced MSC-seeded scaffolds may be of practical use for treatment of hemophilia B and other plasma protein deficiencies. Ó 2010 Elsevier Ltd. All rights reserved.
Keywords: Mesenchymal stem cells Gene therapy Hemophilia Biomineralization Self-renewal Osteogenesis
1. Introduction Gene therapy for inherited life-threatening or severely debilitating diseases has shown promises in pre-clinical and early clinical trials [1e3]. Its clinical translation has however been limited by: 1) the poor safety profile of viral vectors used for in vivo gene transfer, and 2) poor survival of ex vivo modified cells upon transplantation. Although the use of ex vivo gene-engineered cells increases the safety of gene transfer and new gene transfer technologies are safer than earlier viral vectors (zinc-finger nucleases for instance), cell survival post-transplantation remains a challenge. Gene therapy for inherited protein deficiencies such as hemophilia A, B and some lysosomal storage disorders require long-term systemic protein delivery with no tissue-specificity [4]. In essence, any autologous cell type available in sufficient numbers can be used * Corresponding author. Department of Hematology and Medical Oncology, Pediatrics & Medicine, Emory University, 1365 Clifton Road, Clinic B, Suite 5100 e 5117, Atlanta, GA 30322, USA. Tel.: þ1 404 778 1779. E-mail address: [email protected] (J. Galipeau). 0142-9612/$ e see front matter Ó 2010 Elsevier Ltd. All rights reserved. doi:10.1016/j.biomaterials.2010.08.094
to deliver the transgene product ectopically and systemically. Our laboratory has used mesenchymal stem cells (MSCs) to deliver various therapeutic proteins in animal models [5e9]. These cells are attractive candidates for gene and cell therapy, tissue engineering and regenerative medicine, owing to their availability from autologous sources, their intrinsic mesodermal plasticity and their pro-angiogenic, pro-regenerative and immuno-modulatory properties [10e16]. Over the years, we have used a wide variety of natural polymers (Matrigel, collagen and hyaluronic acid) to deliver MSCs in vivo [5e9]. We always experienced cytotoxicity issues when using chemically cross-linked polymers (not shown) and fast resorption rate/loss of protein production using enzymatically cross-linked polymers (Supplemental Fig. 1). This was consistent with data from other groups using a similar strategy, who typically observed systemic protein delivery for only 5e10 days [17e21]. We concluded that these biomaterials were too short-lived in vivo and lacked the adhesion cues necessary to MSCs. Moreover, these transplantation methods being essentially isotropic, they lacked precise control over cell capacity and fate. These findings suggest
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that clinical translation of gene therapy for monogenic diseases requires a multidisciplinary feat of an unforeseen complexity. The ideal biomaterial should have sufficient cell capacity to achieve therapeutic benefit and sustain long-term cell survival, self-renewal, proliferation and differentiation. It should also incorporate biochemical and spatial cues mimicking complex tissue composition, architecture and mechanical modulus. This requires a hierarchical design of the scaffold from the nanoscale to the macroscale, where cell fate decisions are controlled at the nanoand micro-scale levels, whereas the micro- to macroscale organization levels influence cell capacity, mass transport, vascularization and mechanical strength [4,22e24]. Biomaterials should also: 1) be made of clinically-approved materials and degrade into non-toxic waste products; 2) be easily and reproducibly processed, at low cost and using commercially available technologies; and 3) be designed for off-the-shelf use (with a long shelf-life). Fig. 1 illustrates the approach used in the present study to develop a nanopatterned macroporous scaffold custom-designed
to support MSC survival, self-renewal and coagulation factor IX (FIX) delivery in hemophilic mice. First, MSCs isolation and culture conditions were optimized to ensure maintenance of multilineage potential and culture expansion (Fig. 1a). After standard gene transfer, protein biochemistry was used to measure hFIX functionality. The main focus of the current study (illustrated in Fig. 1b and c) relates to the material science and engineering aspect of the platform and the translational application in a suitable animal model (scaffold production, tissue engineering, and blood physiology post-transplantation). Briefly, we used rapid prototyping to create polylactic-co-glycolic acid (PLGA) composite scaffolds with a simple 0 /90 strand architecture maximizing cell capacity, mass transport and tissue ingrowth. We then identified calciumephosphate ceramic (e.g. hydroxyapatite, HA) as a potent biomaterial promoting MSC survival post-transplantation. In composite HAePLGA 3D scaffolds, optimization of HA crystal size and concentration as well as modification of the microtopology and biochemical surface properties with biomineralized
Fig. 1. Multidisciplinary approach to gene therapy of hemophilia B. a) Stem cell biology and engineering: isolated MSCs are characterized by their multipotent lineage potential and expression of cell surface markers. Following hFIX-retrovirus cloning, virus production and MSCs transduction, maintenance of MSC properties is confirmed and hFIX production and functionality is quantified. Right panel: crystal structure of FIX Gla domain compexed with Ca2þ is shown (PDB ID: 1J35). b) Material science and engineering: biomaterial scaffolds are engineered from the nanoscale to the macroscale to ensure appropriate MSC differentiation, self-renewal and hFIX production in vitro and in vivo. Left panel: SEM image of an HA crystal embedded in PLGA polymer; overlays show chemical structures of HA and PLGA. Scale bar ¼ 1 mm. Middle panel: SEM image of a biomineralized collagen fiber and coating on HAePLGA scaffolds; overlay shows FTIR spectrum confirming the presence of biomineralized collagen. Scale bar ¼ 50 mm. Right panel: 3D architecture and macroporosity of scaffolds, as used in this study, measured by confocal microscopy. c) Translational study: scaffolds are prepared using rapid prototyping, sterilized and coated with biomineralized collagen I. The scaffolds are then seeded with hFIX þ MSCs which are allowed to adhere, proliferate and infiltrate the scaffold. The scaffolds are transplanted by microsurgery in R333Q hemophilia B mice and hFIX activity measured weekly using an aPTT assay specially designed to detect hFIX activity in mouse plasma. Hemophilic phenotype is monitored over time and MSC fate and persistence is assessed. Some images were adapted and modified with permission from Coutu, D. L. et al. J. Cell Biochem. 108, 537e546 (2009) [4].
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collagen was necessary to obtain long-term MSC engraftment, selfrenewal and differentiation, along with sustained hFIX delivery in hemophilic mice. We thus present a platform for gene therapy of inherited or acquired protein deficiencies using a combination of autologous cells and clinically available biomaterials. The simple fabrication process, ease of use and long shelf-life of our scaffolds make them readily translatable to the clinic and useful for clinical and pre-clinical applications. Our study also provides insightful data on the behavior, fate and requirements of MSCs on biomaterials in vivo. 2. Materials and methods 2.1. Generation of hFIX retroviral vector plasmid and virus The cDNA of human FIX from pCMVFIX (ATCC) was excised by restriction digest and inserted into the multiple cloning site of a retroviral vector pIRES-GFP. This vector was then digested to remove the IRES-GFP. Sequencing of the plasmid by Bio S&T Inc. (Montreal, Quebec) confirmed correct orientation and sequence. For virus production, hFIX retroviral construct was co-transfected with pVSV-G plasmid into 293GP2 packaging cells (ATCC) in the presence of Polyfect Reagent (Qiagen). Virus was collected for 3 days, filtered and frozen until used.
2.2. Cell culture MSCs from 4 to 8 weeks old R333Q male mice were isolated by whole bone crush and plastic adherence, cultured in DMEM with 10% fetal bovine serum (Wisent) and antibiotics and allowed to expand for 2e4 passages to obtain a homogenous population before retrovector transduction. Multipotentiality was determined as described [25,26]. Vitamin K1 (7 mg/mL) was added to stimulate g-carboxylation. For retroviral transduction, medium was replaced with virus containing conditioned medium from packaging cells diluted 1:1 in growth medium, in the presence of polybrene (2 mg/mL).
2.3. Animals All animal work was approved by the Animal Research Ethics Committee of the Lady Davis Institute for Medical Research. R333Q hemophilia B mice [27] were a kind gift from Dr Darrel Stafford. Only males were used for transplantation. For collagen gel injection, mice were injected subcutaneously on the left flank. For surgical implantation, mice were anesthetized using isoflurane, a subcutaneous pocket was created by blunt dissection on the left flank where the implants were placed and the wound closed with surgical staples. When necessary, animals were sacrificed by CO2 inhalation and implants retrieved and processed for histology. For blood sampling, uncoated capillary tubes were used to draw blood from the saphenous vein. Samples were immediately placed in 3.9% sodium citrate (blood: citrate ratio of 10:1) then centrifuged. Citrated plasma was stored frozen until tested.
2.4. Flow cytometry Cells were harvested, washed with PBS and fixed with paraformaldehyde. Cells were then incubated with the following primary antibodies: anti-mouse CD45, CD31, CD44, CD73, CD105, CD34 or isotypic controls. Appropriate fluorescent secondary antibodies were then applied and analysis performed on a FACS Calibur instrument.
2.5. Western blotting hFIX þ MSC were grown up to 80% confluence, washed extensively with PBS and serum free DMEM was added. Conditioned medium (CM) was harvested after 48 h. After 50-fold concentration by ultrafiltration (MWCO 10 kD), samples were separated by SDS-PAGE, transferred on PVDF membranes and probed with anti-hFIX (Haematologic Technologies Inc) or anti-Gla (American Diagnostica) antibodies and appropriate secondary antibody, then detected using ECL.
2.6. hFIX chromogenic assay To activate hFIX, 1 U/mL hFXIa was added to concentrated conditioned medium from wild type or hFIX þ MSC in the presence of 5 mM CaCl2 and incubated 45 min at 37 C. hFIXa activity was determined using the chromogenic substrate 229 (American Diagnostica). Buffer used was 100 mM Tris, 100 mM imidazole, 100 mM NaCl, 0.1% BSA, 5 mM CaCl2 and 33% ethylene glycol. Samples were compared to a standard curve prepared by serial dilution of purified hFIXa (Haematologic Technologies) in growth medium. Activation of hFIX was confirmed by western blot.
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2.7. hFIX-specific ELISA Assay was performed using the anti-hFIX paired antibodies kit (Cedarlane Laboratories) and performed under the manufacturer’s instructions on concentrated serum free conditioned medium. A standard curve was prepared with purified hFIX (Haematologic Technologies Inc.). 2.8. Scaffolds fabrication and cell seeding Scaffolds were prepared as described [28] with modifications, using 8515DLHighIV PLGA (Lakeshore Biomaterials) with or without the incorporation of hydroxyapatite ceramic particles (Berkeley Advanced Biomaterials). The plotting material was prepared by dissolving PLGA in methyl-ethyl-ketone and mixing with hydroxyapatite particles when necessary. Optimal concentration of the polymer in solvent was determined based on viscosity constraints of the 3D plotter and optimal syringe deposition. Three-dimensional (3D) scaffolds were fabricated by rapid prototyping using a Bioplotter (Envisiontec, Gladbeck, Germany) and computer-assisted design (CAD). Briefly, the plotting material was transferred to the plotting cartridge and was dispensed layer by layer, forming a 0 /90 strand structure. Different topologies were fabricated (strand size and distance between strands) as specified in the text. Plotting needles had 250 and 400 mm inner diameter for PLGA and HAePLGA scaffolds, respectively. Bricks of 20 20 1.1 mm were fabricated. Discs of 6 mm diameter were punched out and vacuum-dried to allow solvent evaporation. Scaffolds were sterilized in 70% ethanol, washed in PBS and coated with 1% collagen type I (BD Bioscience) or biomineralized collagen I [29,30]. For cell seeding, 30 mL of cell suspension (approximately 250 000 cells) was loaded onto the scaffolds by capillarity. After 2 h incubation in a humidified incubator, growth medium was added. The cells cultured for 2 weeks prior to their implantation with frequent medium changes (or additionally for 3 weeks in differentiation medium). 2.9. Scanning electron microscopy SEM was done using a Hitachi S-4700 Scanning Electron Microscope. Scaffolds were thoroughly washed with dH2O and ethanol, then lyophilized and coated using gold/palladium. For cell-seeded scaffolds, a fixation step was included before washing (2.5% glutaraldehyde in 0.1 M sodium cacodylate at pH 7.4 for 2 h). 2.10. hFIX-specific aPTT assay on mouse plasma We used a hFIX-specific aPTT assay modified from Arruda et al. [31]. Briefly, a standard curve was prepared by serially diluting hFIX in citrated plasma from R333Q mice. Samples and standards were diluted 1:10 in imidazole buffer. Samples, FIX-deficient pooled human plasma, aPTT activator reagent and CaCl2 were mixed in equal volumes, calcium being added after 3 min incubation at 37 C. Clotting time was measured using a KC4 coagulometer. 2.11. Histology/immunohistochemistry Implants were retrieved, fixed with formalin and processed for regular histology. Implants were decalcified using 10% EDTA between fixation and processing. Alternatively for undecalcified histology, formalin fixed implants were processed for methylmethacrylate embedding. For immunohistochemistry, we used antigen retrieval solution, avidin/biotin blocking kit and DAB reagent from Vector. Anti-mouse osteocalcin (AbDSerotec), appropriate biotinylated secondary antibodies and streptavidin-HRP (Jackson ImmunoResearch) were used. 2.12. Microscopy Photomicrographs were taken using a Leica DM LB2 microscope equipped with a Leica DFC480 camera and acquired using the Leica Application Suite software. Alternatively, mounted slides were scanned using a NanoZoomer (Hamamatsu) and analyzed using the NanoZoomer Digital Pathology (NDP View) software. 2.13. Confocal microscopy Paraffin sections were stained as described [32] using antigen retrieval solution from Vector, anti-hFIX and anti-Ki67 (Dako) antibodies, appropriate isotypic controls (not shown) and Alexa Fluor 488 or 647 conjugated secondary antibodies and DAPI (Molecular Probes). Confocal imaging was done on a WaveFX/Leica spinning disk confocal microscope. Image acquisition, 3D reconstruction and deconvolution were done using Volocity or Imaris softwares, when necessary. 2.14. Collagen gel preparation Rat tail collagen type I (10 mg/mL) was from BD Biosciences. 1 M199 buffer was supplemented with 0.75% NaHCO2 and 0.01 M HEPES. The injections consisted of 75% collagen gel (with or without 40 mg/mL HA ceramic particles, see below) and 25%
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cell suspension with the following final concentrations: 4.9 mg/mL collagen, 1 mM DTT and 0.24 U/mL tissue non-specific transglutaminase (Sigma). PBS diluted collagen was mixed with complete M199 buffer (60:40 ratio). The pH was adjusted using 0.1 M NaOH. Cells, DTT and transglutaminase were then added. Kept liquid on ice, the gel is loaded into a syringe and injected subcutaneously. 2.15. MSCs implantation on HA particles Was performed as described [33,34] with modifications. MSC suspended in growth medium were incubated with 40 mg HA/106 cells (particle size 250 mm diameter, Plasma Biotal) with slow rotation at 37 C for 2 h. The cell/ceramic suspension was centrifuged and mixed in the collagen I gel described above prior to implantation. Up to 107 cells were transplanted in each animal. 2.16. Fourier transform infra red spectroscopy (FTIR)-microscopy Samples were washed with dH2O and ethanol and then lyophilized. Measurements were done using a Bruker Tensor 27 equipped with a Hyperion 2000 microscope on 35 different points. 2.17. Statistical analyses For aPTT assays, dot plots showing individual mice data overlaid with mean SEM is shown. Two-way ANOVA were performed using GraphPad Prism5. For histograms, data from one of at least three experiments is shown. We did not pool data from different experiments because hFIX production levels varied slightly from one transduced MSC lot to another. All tests were performed at least in triplicates.
3. Results 3.1. MSC characterization, gene-engineering and functionality We used the murine R333Q model of hemophilia B because these mice do not develop inhibitory antibodies against hFIX after gene or protein therapy [35]. These FIX knockout mice express, under the murine FIX promoter, an inactive hFIX carrying the R333Q mutation frequently observed in humans. They are thus human cross-reactive material positive but have a hemophilic phenotype, recapitulating the human condition. As an initial step, MSC were isolated from R333Q mice and transduced with a hFIXretrovirus. They were shown to retain in vitro multilineage differentiation potential into adipocytes, osteocytes and chondrocytes (Fig. 2a). They also presented a typical C57Bl/6 MSC immunophenotype (including CD34) for cell surface markers (Fig. 2b). In our hands, CD105 is rapidly lost with culture expansion of murine MSC. hFIX secretion was quantified by ELISA and found to be maintained throughout multilineage differentiation (Fig. 2c). MSC-derived hFIX functionality was assessed by western blot analysis, confirming g-carboxylation of glutamic acid residues (Glu) to g-carboxyglutamic acid (Gla; Fig. 2d). Functionality was also demonstrated using a chromogenic assay, where MSC-derived hFIX was activated by hFXIa and then quantified using the chromogenic substrate 229. Western blot analysis confirmed the activation (cleavage) of hFIX to hFIXa which proteolytically processed 229, releasing free pNA (Fig. 2e). 3.2. Macro- to micro-scale optimization of 3D porous scaffolds We previously demonstrated that synthetic polymers (such as PLLA/PLGA)-based 3D porous scaffolds created using a combination of rapid prototyping (RP, or solid free forming) and porogen leaching can be optimized for maximal cell capacity as well as mass and oxygen transport in vitro [28]. While RP allows fine control over structure, pore interconnectivity and cell seeding efficiency, microporosity influences cell adhesion, behavior and mass transport. Fig. 1b (right panel) shows a disc-shaped scaffold with a simple 0 /90 architecture similar to those used previously and here. Our initial data demonstrated that these scaffolds supported MSCs adhesion, proliferation and protein production in vitro.
To further optimize tissue ingrowth by MSCs on scaffolds, we here modified scaffold macroporosity by varying strand diameter (150e300 mm), distance between strands (150e300 mm) and layer number (0.6e1.3 mm total thickness). We also varied microporosity using porogen leaching, where NaCl (0e30%) was incorporated into the polymer and then leached in water. Scanning electron microscopy (SEM) demonstrated the drastic effect of micro- and macroporosity modification on MSC ingrowth. Scaffold topologies of 200 250 (strand diameter and distance between strands respectively, in micrometers) and 15e30% porogen provided the best MSC proliferation and tissue ingrowth after two weeks in culture (Fig. 3a). This was surprising since we had previously demonstrated that a scaffold topology of 150 150 with 30% porogen provided the highest total surface area and maximal effective surface area. However, the current results indicate that larger pores facilitate MSC invasion of pore space in 3D scaffolds. We confirmed by ELISA that MSCs seeded on PLGA scaffolds maintained hFIX production in vitro throughout mesodermal differentiation (Fig. 3b). We next implanted these optimized constructs in hemophilic mice. We followed systemic hFIX delivery by weekly aPTT assays and retrieved the scaffolds after 8 weeks for histological analysis. Alcian blue stain (for cartilage-specific proteoglycans) revealed that the implants contained mostly fibrous tissue, with evidence of mesenchymal condensations staining weakly with alcian blue and extensive vascularisation (Fig. 3c). Toluidine blue/von Kossa staining (for osteoid matrix and mineralized phosphate, respectively) showed a complete absence of mineralization and osteoid deposition (Supplemental Fig. 3, left panel). Systemic hFIX activity reached therapeutic levels for a short two week period before returning to baseline. These results identified the optimal 3D architecture and macroporosity of scaffolds for in vitro MSC proliferation and tissue ingrowth. However, the PLGA substrate clearly did not allow long-term in vivo MSC persistence (self-renewal) and longterm systemic hFIX delivery at therapeutic levels. 3.3. Calciumephosphate ceramics for MSC delivery Calciumephosphate ceramics (such as hydroxyapatite, HA) have a long history of clinical use and have been demonstrated by many groups to sustain long-term engraftment of MSC in vivo [33,34,36,37]. However, ceramics are not easily processed into complex 3D structures. We first tested the ability of isotropically transplanted 175e250 mm diameter HA ceramic particles to promote MSC self-renewal and protein delivery in vivo. We used a transglutaminase cross-linked collagen I hydrogel to hold the particles together (see Supplemental Fig. 1). Fig. 4a illustrates the sequence of bone development by murine MSC on HA. At 2 weeks post-implantation, the implants are very cellular (with non-hematopoietic, presumably donor-derived cells) with evidence of de novo ECM deposition. At week 4, the implants show a decreased cellularity and predominance of fibrous tissue. At week 6, an increase in cell content is observed, coincident with vascular invasion. Within 7e8 weeks, bone tissue is formed within the implant, with an outer periosteum-like capsule that eventually mineralizes and contains hFIX þ cells (not shown). A hematopoietic marrow-like cavity is observed and the implants are densely irrigated by large, sinusoidlike blood vessels. They also contain extravascular erythrocytes suggestive of erythropoiesis (Supplemental Fig. 2). At 10 weeks post-implantation, the presence of hFIX þ MSC closely attached to the carrier particles can still be detected by immnohistochemistry (Fig. 4b). HA-seeded MSC produced detectable systemic hFIX for up to 12 weeks, although the levels remained below the 2% therapeutic threshold (Fig. 4c). The low total surface area of the ceramic particles (low cell capacity) likely explains these results. Moreover, 40 mg of particles is required for every 106 cells [38], making scaling up
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Fig. 2. MSCs characterization, gene modification and hFIX production. a) MSCs from R333Q hemophilic mice displayed typical mesodermal plasticity and could be differentiated into osteocytes, adipocytes and chondrocytes (left to right: alizarin red S, oil red O and alcian blue stains) after retroviral transduction. b) Flow cytometry immunophenotyping of R333Q MSCs reveals expression of murine MSC markers CD34, CD44 and CD73 with the absence of hematopoietic/endothelial markers CD45 and CD31. c) MSCs maintained hFIX secretion after their differentiation into mesodermal lineages as measured using a hFIX-specific ELISA. (UD: undifferentiated). d) Western blot analysis of wild type (WT) or hFIX þ MSCs conditioned medium demonstrates the presence of g-carboxyglutamic acid on MSC-derived hFIX. e) Upper panel: MSC-derived hFIX functionality was measured using a hFIX-specific in vitro chromogenic assay. hFIX in MSC conditioned medium was activated with hFXIa and its activity measured with the hFIX-specific chromogenic substrate 229. Lower panel: hFIX activation was confirmed by western blot. (FL: full length, HC: heavy chain).
impractical due to the large size of the implants, which also limits mass transport. These results show that HA particles promote long-term engraftment and hFIX delivery by MSCs and that osteogenesis is
required for both processes. However, isotropically transplanted particles lack the 3D architecture maximizing cell capacity and hFIX delivery in vivo. We thus used rapid prototyping to create composite HAePLGA macroporous 3D scaffolds.
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Fig. 3. Macro- and microporosity optimization of 3D porous scaffolds. a) PLGA scaffolds were created using solid free forming processing. Macrotopology was modified by varying strand diameter and distance between strands. Microtopology was modified using porogen leaching. Scaffolds seeding and tissue ingrowth by MSCs in vitro was assessed by SEM two weeks after seeding. Images demonstrate that both macro- and micro-porosity influence the capacity of MSCs to adhere, proliferate and invade scaffold pores. Best tissue ingrowth was observed with the 200 250 (30%) topology, where 200 is strand diameter, 250 is distance between strands, and 30% the porogen concentration. Scale bar ¼ 100 mm b) MSCs maintained hFIX secretion after their differentiation into mesodermal lineages on porous PLGA scaffolds, as measured using a hFIX-specific ELISA. c) Histological analysis of retrieved implants 5 weeks post-implantation reveals poor chondrocytic differentiation (alcian blue stain shown). The implants contained few mesenchymal condensations (arrowheads) and large blood vessels (arrow). Scale bar ¼ 100 mm. C: carrier. d) MSCs were seeded on the best scaffold topology and differentiated into osteoblast, adipocytes or chondrocytes before transplantation. hFIX delivery was monitored by weekly aPTT assays. Only the chondrocytic group (shown) reached hFIX levels above therapeutic threshold, very transiently.
3.4. Nano- to micro-scale optimization of 3D scaffolds To create HAePLGA composite scaffolds, we first needed to optimize the composition, concentrations, and formulation of both components within the scaffolds. HA physico-chemical properties (crystal size and perfection for instance) influence its osteoconductivity [39]. Moreover, crystal size also affected our capacity to process HAePLGA porous scaffolds using rapid prototyping. For instance, 200 mm particles were poorly miscible in PLGA and were not tested further. To avoid delamination problems, we also had to increase both strand diameter and distance between strands to 300 mm, reducing the effective total surface area by about half (not shown). Nevertheless, we successfully created composite scaffolds using two types of HA, nano-HA and micro-HA (100 nm and 5 mm particle diameter, respectively). Our findings are summarized in Supplemental Figs. 3 and 4. Briefly, we found that nano-HA was more osteoconductive than micro-HA at the same concentration (40%). However, both constructs did not support long-term hFIX delivery by MSCs in vivo. At higher concentration, occlusive aggregation of the nano-HA particles occurred within the printing
syringe whereas we were able to reach 65% micro-HA concentration. Our data shows that micro-HA (65%) is more effective than nano-HA (40%) in supporting osteogenesis and hFIX delivery by MSCs. Cell adhesion to biomaterials such as PLGA and HA is thought to require the presence of binding proteins to bridge between cells and material. Coating with serum, collagens or other extracellular matrix proteins is typically used. Because we observed that osteogenesis was necessary for long-term MSC engraftment and hFIX delivery, we sought to direct MSC adhesion and fate decision by modifying the biochemical surface properties and microtopology of the scaffolds. More specifically, we directly compared two different coatings: collagen type I (the most abundant collagen in bones), and biomineralized collagen I [29,30]. To assess the coating efficiency and modification of microtopology we used a combination of SEM and Fourier Transform Infra Red (FTIR)-spectroscopy and microscopy (Fig. 5 and Supplemental Fig. 5). Uncoated PLGA scaffolds show a very smooth surface texture under SEM whereas HAePLGA composite scaffolds presented a rougher surface (Fig. 5a), with evidence of HA particles exposed
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Fig. 4. MSCs persistence, fate and hFIX production on HA/TCP ceramic particles. hFIX þ MSCs were seeded on 200e500 mm HA/TCP ceramic particles and injected subcutaneously in R333Q mice. Bone formation was assessed by histological analysis of retrieved implants and hFIX systemic delivery by weekly aPTT assays. a) Sequence of bone development by MSCs on HA/TCP in vivo. Donor-derived MSCs are highly present in the implants at 2 weeks post-transplant but are gradually replaced with host-derived hematopoietic marrow at 6 weeks, concomitant with development of vasculature (arrowheads). At 7 weeks, the bone-like implant contains a readily identifiable marrow cavity (*). b) hFIX þ MSCs closely attached to the ceramic particles are still detectable by IHC after 10 weeks. c) Weekly aPTT assays reveal that although hFIX persists in the plasma of the transplanted animals for up to 12 weeks, the levels produced are below the 2% normal activity (100 ng/mL) therapeutic threshold (dashed line). Scale bars ¼ 100 mm.
on the surface (presence of calcium was confirmed using X-ray fluorescence, not shown). Upon coating with collagen or biomineralized collagen, the microtopology and texture of the scaffolds were greatly modified (Fig. 5b). In the former, we can observe large collagen fibrils and bundles of various sizes covering the surface. In the later, both the protein-rich collagen coating and large fibrils are completely mineralized. Higher magnification SEM images (Fig. 5c) demonstrate the presence of free-floating HA crystals (upper panel) and collagen fibrils showing evidence of clubbing and bulging, reflecting nucleation of spherulitic HA within gaps of the fibrils [40] (lower panel). We performed FTIR spectroscopy to confirm the composition of the coating and to measure changes in surface chemistry (Fig. 5d). The characteristic n3 vibration mode of phosphate within HA (1075 cm1) was clearly visible in all HAePLGA scaffolds. The sharp amine band near 1750 cm1 and increased hydroxyl band confirmed efficient collagen coating. The n1 phosphate vibration mode (960 cm1) was further increased by biomineralization, suggesting a higher crystalline HA at the strand surface. Other dominant bands between 800 and 1110 cm1 likely reflect various amorphous and protonated phosphates [41]. FTIR-microscopy
confirmed efficient coating of the whole scaffold on at least three layers (Supplemental Fig. 5). 3.5. MSC self-renewal, fate and hFIX delivery on fully optimized 3D scaffolds As with previous experiments, we tested our optimized scaffolds by seeding them with hFIX þ MSCs and transplanting them subcutaneously in R333Q hemophilic mice. We monitored plasmatic hFIX levels by weekly aPTT assays and MSC fate by histology and immunofluorescence confocal microscopy on retrieved implants (Fig. 6). Toluidine blue/von Kossa staining of implants retrieved at 15 weeks post-transplantation demonstrates the extent of osteogenesis in optimized scaffolds. We observed increased osteogenesis, scaffold remodelling, mineralization and cellular invasion in the biomineralized collagen-coated scaffolds than with any other scaffold used (compare Fig. 6a with Supplemental Figs. 3, 4 and 6). To assess MSC self-renewal within the scaffolds, we used immunofluorescence confocal microscopy and were able to demonstrate the persistence of hFIX þ cells co-expressing the proliferation marker Ki67 15 weeks after transplantation (Fig. 6b). These cells were almost
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Fig. 5. Nano- and micro-topology optimization of 3D porous scaffolds. a) SEM images of two 3D porous scaffolds used in this study, show that PLGA scaffolds have a very smooth surface whereas the surface texture of HAePLGA scaffold is rougher (we can also see HA crystals, which appear brighter on the SEM) protruding at the surface of the strands. b) Modification of scaffolds surface properties (chemistry and topology) as assessed by SEM. Scaffolds were coated with collagen type I or biomineralized collagen type I. Left: After coating with collagen, the proteinaceous layer completely covers the scaffold strands and large collagen fibrils are observed. Right: After subsequent biomineralization of the collagen coating, the microtopology is greatly affected and collagen fibrils of all sizes appear mineralized. Boxed area is shown at higher magnification in c). c) High power SEM images of the biomineralized coating of HAePLGA scaffolds reveals the presence of free HA crystals (arrowhead) attached to the thick layer of biomineralized collagen I. The small collagen fibers show bulging and clubbing, suggesting the presence of spherulitic apatite within fibril gaps. d) FTIR spectra of the various 3D scaffolds used in this study confirm coating efficiency and composition. The apatite-specific n3 vibration mode of phosphate (1075 cm1, double arrow) can be seen in all HAePLGA scaffolds. The amine band is clearly visible after collagen I coating. After collagen biomineralization, the n1 vibration mode of phosphate (960 cm1, single arrow) is increased.
exclusively localized closely attached to the scaffold strands. Interestingly, as seen with calciumephosphate ceramic particles we also observed the creation of a hematopoietic marrow-like cavity with the scaffold pores, as evidenced by abundant extravascular erythrocytes (Fig. 6c). The formation of this myelosupportive environment was only observed with these two transplantation vehicles. We were also able to detect long-term hFIX production in hemophilic mice at therapeutic levels in most animals, although we did observe variability in responses to the treatment (Fig. 6d). hFIX levels remained detectable, albeit with reduced levels after 12 weeks, for the duration of the experiment (15 weeks). Interestingly, plasmatic hFIX levels oscillated with a period of 10e20 days. It is
worth mentioning that murine osteoblasts have an estimated halflife of 10e20 days [36,42]. Thus, the oscillating levels of hFIX observed may well represent normal cell turnover and replacement, again suggesting the persistence of self-renewing MSCs within the implants. 4. Discussion In this study, we developed a 3D porous scaffold with a hierarchical design made with clinically-approved biomaterials that, combined with autologous mesenchymal stromal cells, could be used off-the-shelf for personalized cell-based gene therapy of
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Fig. 6. In vivo performance of MSC-seeded, fully optimized 3D porous scaffolds for hFIX delivery in hemophilic mice. Histological and immunofluorescence analysis of retrieved implants at 15 weeks post-transplant. a) von Kossa/toluidine blue stain reveals important remodelling and mineralization (black staining) of the scaffolds (left panel; scale bar ¼ 200 mm). The presence of multinucleated osteoclasts in the mineralized and partially degraded carrier material (c) is shown (*), as well as the deposition of osteoid-like matrix on the scaffold strands (deep blue staining and arrowheads; right panel). b) MSCs persistence, survival and self-renewal within the implants was assessed by immunofluorescence and confocal microscopy. Image shows co-localization of hFIX and the proliferation marker Ki67 in cells closely attached to the scaffold strands (arrowheads). Dashed line represents the tissue/scaffold interface. Scale bar ¼ 50 mm c) Hematoxylin/eosin staining shows the presence of hematopoietic tissue within scaffold pores, as evidenced by cell morphology and the presence of extravascular erythrocytes (arrowheads). Scale bar ¼ 50 mm d) Weekly aPTT assays on the plasma of transplanted animals demonstrates reversal of the hemophilic phenotype of most animals for up to12 weeks, with detectable hFIX persisting for the duration of the experiment (15 weeks).
inherited and acquired protein deficiencies. We used rapid prototyping because it can be performed with commercially available technologies and offers great control over pore size, orientation, interconnectivity and total surface area in the scaffolds. When combined with porogen leaching, it also allows maximal cell capacity and mass transport through the implant [28]. As described by others [43e46], our study confirms that substrate macro- and micro-porosity as well as nano- and micro-texture greatly influences cell behavior in vitro and in vivo. Furthermore, we show that only calciumephosphate ceramics provide the necessary signals promoting long-term MSCs survival and self-renewal. These could be further increased by modification of the surface chemistry and nano-topology of the scaffold, and by modifying particle size and concentration of the ceramic. The complete reversal of the hemophilic phenotype in mice demonstrated in this study strongly suggests that this approach has clinical relevance. The fact that we were able to cure the mice for over 12 weeks is a considerable improvement over previous attempts using similar strategies [17e21]. Although some animals
eventually lost therapeutic hFIX activity over time, additional rounds of transplantation could easily be contemplated. Since R333Q mice are not known to mount immune responses to hFIX in gene and protein therapy, potential explanations for loss of hFIX activity include promoter silencing (we used retroviral gene transfer and LTRs are subject to inactivating methylation in vivo). Furthermore, as seen with hematopoietic stem cells it is not unlikely that subpopulations of MSCs exist that possess either short- or long-term “reconstituting” (self-renewal) potential. These subpopulations have however not yet been described for MSCs. Nevertheless, human MSCs of good quality are much easier to isolate and expand ex vivo than murine MSCs (especially in the C57Bl/6 background) [47,48] suggesting that our strategy would also be successful in humans. Finally, it should be mentioned that FIX is a very abundant protein (with systemic levels of nearly 5 mg/mL). Thus, reaching the relatively low 2% therapeutic threshold still requires 100 ng/mL systemic FIX, a huge protein load. In comparison, therapeutic efficacy for hemophilia A would only require 4 ng/mL systemic factor VIII.
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5. Conclusion In summary, we here present a clinically relevant platform for mesenchymal stem cell-based gene therapy of protein deficiencies such as hemophilia B. Our data illustrates how long-term selfrenewal and differentiation of MSCs can be manipulated through hierarchical scaffold design from the nanoscale to the macroscale. Our use of commercially available technologies and clinically accepted materials combined with autologous cells make our approach easily translatable to the clinic. Our data also provides key insights into the physico-chemical environment necessary for longterm MSCs survival and self-renewal in vivo. Acknowledgements The authors would like to thank Yoon Kow Young and the WEBC Imaging Facility for assistance with confocal microscopy; the McGill Bone Center for MMA embedding and sectioning; Micheline Fortin and the IRIC Histology Platform for assistance with histology and digitalization of slides; Christian Degrandpré for scaffold fabrication; Dr Ian B. Copland and Dr Janet E. Henderson for useful discussions and technical help; the Canadian Stem Cell Network and Fonds de Recherche en Santé du Québec for funding. Appendix. Supplementary data Supplementary data associated with this article can be found, in the online version, at doi:10.1016/j.biomaterials.2010.08.094. Appendix Figures with essential color discrimination. Figs. 1e4 and 6 in this article are difficult to interpret in black and white. The full color images can be found in the online version, at doi:10.1016/j. biomaterials.2010.08.094. References [1] Edelstein ML, Abedi MR, Wixon J. Gene therapy clinical trials worldwide to 2007ean update. J Gene Med 2007;9:833e42. [2] Chuah MK, Collen D, VandenDriessche T. Preclinical and clinical gene therapy for haemophilia. Haemophilia 2004;10(Suppl. 4):119e25. [3] Murphy SL, High KA. Gene therapy for haemophilia. Br J Haematol 2008;140:479e87. [4] Coutu DL, Yousefi AM, Galipeau J. Three-dimensional porous scaffolds at the crossroads of tissue engineering and cell-based gene therapy. J Cell Biochem 2009;108:537e46. [5] Copland IB, Jolicoeur EM, Gillis MA, Cuerquis J, Eliopoulos N, Annabi B, et al. Coupling erythropoietin secretion to mesenchymal stromal cells enhances their regenerative properties. Cardiovasc Res 2008;79:405e15. [6] Kucic T, Copland IB, Cuerquis J, Coutu DL, Chalifour LE, Gagnon RF, et al. Mesenchymal stromal cells genetically engineered to overexpress IGF-I enhance cell-based gene therapy of renal failure-induced anemia. Am J Physiol Renal Physiol 2008;295:F488e96. [7] Eliopoulos N, Lejeune L, Martineau D, Galipeau J. Human-compatible collagen matrix for prolonged and reversible systemic delivery of erythropoietin in mice from gene-modified marrow stromal cells. Mol Ther 2004;10:741e8. [8] Stagg J, Lejeune L, Paquin A, Galipeau J. Marrow stromal cells for interleukin-2 delivery in cancer immunotherapy. Hum Gene Ther 2004; 15:597e608. [9] Eliopoulos N, Al Khaldi A, Crosato M, Lachapelle K, Galipeau J. A neovascularized organoid derived from retrovirally engineered bone marrow stroma leads to prolonged in vivo systemic delivery of erythropoietin in nonmyeloablated, immunocompetent mice. Gene Ther 2003;10:478e89. [10] Prockop DJ. “Stemness” does not explain the repair of many tissues by mesenchymal stem/multipotent stromal cells (MSCs). Clin Pharmacol Ther 2007;82:241e3. [11] Prockop DJ, Gregory CA, Spees JL. One strategy for cell and gene therapy: harnessing the power of adult stem cells to repair tissues. Proc Natl Acad Sci U S A 2003;100(Suppl. 1):11917e23. [12] Stagg J, Galipeau J. Immune plasticity of bone marrow-derived mesenchymal stromal cells. Handb Exp Pharmacol 2007;180:45e66.
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