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Chapter 2
Material and methods
2.1 Cardiac patches - Methods of fabrication
To perform our scaffolds we used two different types of techniques.
2.1.1 Freeze drying
Three-dimensional degradable porous polymeric structures with high porosities (93– 98%) and well interconnected pore networks have been prepared by freeze-drying polymer solutions. Freeze-drying is widely used to convert solutions of unstable materials into solid form by removing the solvent.
In freeze-drying, the solution is filled into the vials and placed into the temperature-controlled shelves of the freeze-dryer. The shelf temperature is lowered to around − 40°C, usually in several steps, which results in the transformation of most of the solution into “ice”. Once the material has completely solidified, the vacuum pump is turned on to evacuate the product chamber. Usually the chamber pressure is maintained between 50 and 200 mTorr such that the chamber pressure is well below the vapor pressure of ice at the target product temperature, resulting in high
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sublimation rate.
This step is referred to as primary drying, wherein water is removed by sublimation of ice.
In the last step, referred to as secondary drying, most of the unfrozen water is removed by desorption at elevated shelf temperature with the chamber still under vacuum. The residual water content at the end of secondary drying is about 1%.
The freeze-drying process is a problem in coupled heat and mass transfer, and heat and mass transfer issues must be recognized to achieve process control and optimization. Every single step of the freeze-drying process presents a unique process development and scale-up challenge. The ice nucleation temperature during the freezing step determines the size and morphology of the ice crystals. The ice crystals are removed by sublimation during primary drying leaving behind a porous cake, which is frequently a “template” of the ice crystal structure. [30]
2.1.2 Fibrous scaffolds
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ejection through a capillary tip or needle. The jet coming from the needle draws towards a collector due to an electric field ranging from 10 to 30 kV. Evaporation of the solvent from the jet after leaving the needle results in fiber deposition on the collector. To obtain continuous fibers, the method requires using solutions containing relatively high polymer concentrations of usually around 10‐15 wt%. The diameter of the fibers is within the range of nanometers to microns. Varying the process parameters, e.g. strength of the electric field, distance between needle–collector, polymer concentration, allows tuning of the fiber. A major advantage of electrospinning is the high flexibility and fiber resolution of the obtained scaffold. Additionally, alignment of the electrospun fibers is enabled to induce cell and tissue alignment. A drawback of electrospinning is the risk of breaking fibers during fabrication, which might lead to inferior quality of the scaffold. [31]
2.2 Biomaterials for cardiac tissue engineering
Due to the variation in mechanical properties required in ‘soft’ versus ‘hard’ TE applications, the constructs for these two sub‐categories generally use different
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classes of biomaterials. For soft TE applications, e.g. skeletal muscle or cardiovascular substitutes, generally a wide variety of polymers are applied. Frequently used biomaterials originate from a wide range of natural as well as synthetic sources. Apart from single polymers, scaffolds are also commonly fabricated from co‐polymers of two or more polymers (not listed) to improve the overall characteristics; co‐polymers generally have an average of the mechanical properties of the incorporated single polymers. [31]
2.2.1 Poly-ε-Caprolactone (PCL)
A polymer selected of the polyester family is poly(ε‐caprolactone (PCL), a semi‐ crystalline rubbery polymer with a very low Tg of around ‐60 °C [21]. Generally PCL degrades by bulk hydrolysis like PLA, although also enzymatic degradation can occur under certain conditions. Degradation is significantly slower compared to PLA due to limited fluid inflow as result of the close packed macromolecules; in vivo degradation time extents to over 2 years [1, 29]. Therewith, PCL is mainly suitable for long‐term implants.
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2.2.2 Poly(glycerol sebacate) (PGS)
Poly(glycerol sebacate) (PGS) is a biodegradable polymer increasingly used in a variety of biomedical applications. This polyester is prepared by polycondensation of glycerol and sebacic acid. PGS exhibits biocompatibility and biodegradability, both highly relevant properties in biomedical applications. The mechanical properties and degradation kinetics of PGS can be tailored to match the requirements of intended applications by controlling curing time, curing temperature, reactants concentration and the degree of acrylation in acrylated PGS. Because of the flexible and elastomeric nature of PGS, its biomedical applications have mainly targeted soft tissue replacement and the engineering of soft tissues, such as cardiac muscle, blood, nerve, cartilage and retina. However, applications of PGS are being expanded to include drug delivery, tissue adhesive and hard tissue (i.e., bone) regeneration. The polymer synthesis was carried out in two steps: (1) pre polycondensation step and (2) crosslinking. For the polycondensation process, equimolar mixtures (1 M) of glycerol and sebacic acid were reacted at 120 ◦C under argon for 24 h before the pressure was reduced from 1 Torr to 40 mTorr over 5 h. For the crosslinking step the
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prepolycondensed polymer (prepolymer) was further kept at 40 mTorr and 120 ◦C for 48 h. PGS is a partially semicrystalline polymer and therefore its thermal properties depend on the temperature relative to the glass to rubber transition temperature Tg of the amorphous phase and the melting temperature Tm of the crystalline phase. It has two crystallization temperatures at −52.14 ◦C and −18.50 ◦C, and two melting temperatures at 5.23 ◦C and 37.62 ◦C. No glass transition temperature was observed above −80 ◦C, which was the lower detection limit of the instrument used in the study. DSC results indicate that the polymer is totally amorphous at 37 ◦C. [32]
2.2.3 Polyvinyl alcohol (PVA)
Polyvinyl alcohol (PVOH, PVA, or PVAl) is a water-soluble synthetic polymer. PVA is nontoxic. It is not biodegradable polymer.
It has the idealized formula [CH2CH(OH)]n. PVA is an atactic material that exhibits crystallinity. PVA instead is prepared by first polymerizing vinyl acetate, and the resulting polyvinylacetate is converted to the PVA. The conversion of the polyesters is usually conducted by base-catalysed transesterification with ethanol and the
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properties of the polymer depend on the amount of residual ester groups. Polyvinyl alcohol has excellent film forming, emulsifying and adhesive properties. It is also resistant to oil, grease and solvents. It has high tensile strength and flexibility, as well as high oxygen and aroma barrier properties.
PVA has a melting point of 230 °C and 180–190°C (356-374 degrees Fahrenheit) for the fully hydrolysed and partially hydrolysed grades, respectively. It decomposes rapidly above 200 °C as it can undergo pyrolysis at high temperatures. [33]
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2.2.4 Polymer blends
In order to improve the properties and extend the application fields of the biodegradable polymers, polymer blending is an effective and convenient way. Miscibility of polymer blends can be divided into three types: namely, completely miscible and partially miscible and completely immiscible.
In terms of crystallizability of the components binary polymer blends can be classified into three types: namely, amorphous/amorphous, amorphous/crystalline and crystalline/crystalline polymer blends.
The structure and properties of biodegradable polymer based blends were influenced significantly by the miscibility, phase behaviour and crystallization conditions. [34].
2.2.5 Polymer blends for biomedical applications
The requirements for a material to be used for tissue engineering purposes are biocompatibility, and biodegradability, as the scaffold should degrade with time and be replaced with newly regenerated tissues.
Various synthetic biopolymers have been used to satisfy different clinical requirements.
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Polyethylene terephthalate: PET consists of polymerized units of the monomer
ethylene terephthalate, with repeating C10H8O4 units. PET has a large use in the bioengineering field. It’s used to make sutures, supports for biosensors, different types of cardiovascular prosthesis (valves, cardiac vessels), different systems for drug delivery like IUD PROGESTASERT® or NORPLANT® sytems.
Poly(lactic-co-glycolic) acid: PLGA is is a copolymer which is used in a host of Food
and Drug Administration (FDA) approved therapeutic devices, owing to its biodegradability and biocompatibility. PLGA is synthesized by means of random ring-opening co-polymerization of two different monomers, the cyclic dimers (1,4-dioxane-2,5-diones) of glycolic acid and lactic acid.
PLGA has been successful as a biodegradable polymer because it undergoes hydrolysis in the body to produce the original monomers, lactic acid and glycolic acid. These two monomers under normal physiological conditions, are by-products of various metabolic pathways in the body. Since the body effectively deals with the two monomers, there is minimal systemic toxicity associated with using PLGA for drug delivery or biomaterial applications. Also, the possibility to tailor the polymer
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degradation time by altering the ratio of the monomers used during synthesis has made PLGA a common choice in the production of a variety of biomedical devices, such as, grafts, sutures, implants, prosthetic devices, surgical sealant films, micro and nanoparticles.
For our works we used a polymer blend made by PGS and PCL ratio 2:1. The blends were obtained using different techniques such as foaming by freeze drying and electrospinning.
2.3 Cardiac patches
The patients with heart disease often experience improvement in quality of life following clinical treatment these therapies do not directly repair damaged myocardium. Furthermore, because the tissue is never directly restored to its prior health, individuals may never regain their original cardiovascular function and may experience other debilitating cardiac conditions as time progresses. In order to alleviate the long-term consequences of cardiovascular disease researchers and clinicians are seeking viable cell sources and novel cell delivery platforms that allow for replacement of diseased tissue and engraftment of new cardiomyocytes from a
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readily available in vitro source. Tissue patches offer an alternative cellular engraftment mechanism, posing additional benefits and challenges to cardiac regeneration therapy. Patches may be used to repair congenital defects and also to cover infarcted areas, delivering cellular and pharmacological therapies. Materials currently used for patching cardiac defects include nondegradable synthetic polymers and autologous and bovine pericardium.
Biomimetic materials can play an important role in multiple aspects of cardiac tissue regeneration. As we have discussed, the task of repairing the heart is multifaceted and includes several types of cardiac tissues, numerous pathologies, and multiple potential approaches. In each of the common cardiac regeneration strategies, biomimetic materials can function as key contributors to achieving overall success. To obtain sufficient numbers of cells for cardiac regeneration studies, biomimetic materials can be used to direct the in vitro expansion and differentiation of stem cells into specialized cardiac cells, including cardiomyocytes. To enhance cardiac function in studies injecting cells directly into the heart, biomimetic materials can be used to support the cells during injection, target them to desired locations, and aid in cell
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survival or to be used like a support to repair a damaged tissue after cellularization. [35]
2.4 Methods of fabrication
To prepare the scaffold we used two types of method of fabrication and different polymers and their blends.
2.4.1 PGSPCL foam
The first type of scaffold was made by a blend of PGS (prepolymer 24 h) and PCL. Different concentrations of this blend were investigated, respectively 30% - 15% - 6% - 3% - 1%. The ratio PGS:PCL was 2:1 and solvent was DMC.
The freezing method was a mix of liquid nitrogen (LN) freezing and freezing by freezer at -20°C to obtain different sizes of pores due to different types of freezing. The first step of method was dissolve the polymers into DMC, then freeze the Teflon® mold into a liquid nitrogen to facilitate the freezing with liquid nitrogen and prevent a cracks due to thermal shock. The mold was freeze for 2 minutes for each side into a 200 mL of LN. After this we put the solution into the frozen mold, when the solution was frozen on the surface (after few second) we put all the mold into a
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LN for 1 minute. The last step of freezing is to put the mold into freezer at -20°C over night.
After freeze we put the mold with the solution into a FD over night and we obtain a foam of PGSPCL. Is not necessary, in this case, crosslink the PGS because the PCL gives to a foam the good mechanics stability.
2.4.2PGS foam made from PVA template
To obtain the second type of foam we used a different technique that involves the use of PVA template.
The first step to obtain the scaffold was made a PVA foam by FD. Different concentrations (w/v) of PVA was investigated to obtain the best porosity respectively 8% - 4% - 2% - 1%.
The PVA was dissolved in water, then the solution was put in a Teflon® mold and frozen by freezer at -4°C for 1 week.
After this the frozen solution was put in a FD to obtain a PVA foam. The foam of PVA is used as a template for a PGS solution. Different concentrations (w/v) of PGS was
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investigated, respectively 30% - 20% - 10% and also different volume of PGS solution respectively 200µL, 100µL, 50µL for each foam of PVA.
The PGS was dissolved in DCM then was put drop to drop into a PVA foam to cover the entire surface of PVA foam. The volume of PGS solution used was 50µL for each PVA foam. After the evaporation of DCM the foam was put in the furnace a 120°C for 48h under vacuum to crosslink the PGS. The last step of procedure is leaching out the PVA by washing with water. The bath water was changed every 30 minutes for two days.
We chose as first parameter of discrimination the hardness of scaffold after leaching, after this we carried out analytics test to choose the best scaffold.
2.4.3 Electrospinning
To obtain a different type of scaffold we used a different technique as electrospinning. Our target was to obtain an orientation of fibers to guide the attachment of the cells. In function of this we used different strategies but always the same polymer blends. The blends that we used was PGS:PCL ratio 2:1 dissolved in DCM:Meth 7:3 at concentration of 15% w/v.
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The solution was spun against different types of support with a field of 15kV, a distance of 20 cm form collector and needle and 40 mL/h as a flowrate.
To guide the fibers in a specific directions we used different supports:
Metal grid with 100µm pores size
Plastic grid with 400µm pores size
Stainless steel sieve with 100 µm pores size
Silica support with different geometry printed by laser
The solution was just spun against the support and the fibers have followed more or less a given alignment.
