| HS Code | 275320 |
| Chemical Formula | (C2H4O)n |
| Water Solubility | soluble_in_water |
| Biocompatibility | high |
| Biodegradability | enzymatically_biodegradable |
| Mechanical Strength | moderate_and_tunable |
| Printability | excellent_for_extrusion_based_bioprinting |
| Viscosity | adjustable_with_concentration_and_molecular_weight |
| Crosslinking Capacity | chemical_and_physical_crosslinking_possible |
| Thermal Stability | decomposes_above_200_degrees_celsius |
| Hydrophilicity | highly_hydrophilic |
| Functional Groups | hydroxyl_groups_available_for_modification |
| Degradation Byproducts | non_toxic |
| Gelation Temperature | concentration_dependent |
| Porosity Control | adjustable_via_process_parameters |
As an accredited Polyvinyl Alcohol (PVA) for 3D Bioprinting Scaffolds factory, we enforce strict quality protocols—every batch undergoes rigorous testing to ensure consistent efficacy and safety standards.
| Packing | Sealed glass jar containing 50 g of sterile Polyvinyl Alcohol (PVA) powder, packaged under inert gas for 3D bioprinting scaffold applications. |
| Container Loading (20′ FCL) | Polyvinyl Alcohol (PVA) for 3D bioprinting scaffolds loaded in 20′ FCL, securely packed, moisture-protected, and containerized for safe transport. |
| Shipping | Polyvinyl Alcohol (PVA) for 3D bioprinting is shipped in sealed, moisture-proof containers to prevent hydration and clumping. Transport at ambient temperature in dry conditions, away from direct sunlight and excessive humidity. Non-hazardous, but handle gently to preserve powder integrity and sterility for scaffold applications. |
| Storage | Store Polyvinyl Alcohol (PVA) in a tightly sealed, airtight container in a cool, dry environment. Protect from moisture and humidity, as PVA is hygroscopic and water-soluble. Avoid exposure to dust and direct sunlight. Maintain room temperature (20–25°C) and keep away from heat sources. |
| Shelf Life | Store unopened in a cool, dry place. Typical shelf life is 2 years from manufacture for optimal bioprinting performance. |
| Freeze-Thaw Cycles | Tensile Modulus (MPa) | Elongation at Break (%) | Crystallinity Index (DSC, %) | Equilibrium Water Content (wt%) |
|---|---|---|---|---|
| 1 | 0.8 ± 0.2 | 220 ± 35 | 12–18 | 91 ± 2 |
| 3 | 2.4 ± 0.4 | 145 ± 20 | 28–34 | 85 ± 3 |
| 5 | 5.1 ± 0.7 | 85 ± 15 | 42–48 | 78 ± 3 |
| Clinical Application | Core Standard | Test Endpoint | Acceptance Criterion |
|---|---|---|---|
| Cartilage Plug | ISO 10993-5:2009 | L929 elution cytotoxicity | Viability index >70% |
| Sacrificial Vascular Channel | ISO 10993-18:2020 | Residual TOC/acetate | < 20 ppm in final rinsate |
| Corneal Stroma Replacement | ISO 11979-5:2020 | Light transmittance, haze | Transmittance >85%, haze < 4% |
| Tumor Spheroid Model | ASTM F3206-17 | Bioink cell compatibility | Viability >85% at 24 h |
| Nerve Conduit | ISO 10993-6:2020 | Local tissue response (intramuscular implant, 28 d) | No adverse tissue necrosis |
| Osteochondral Plug | ASTM F2883-11 | Gradient compressive modulus | Interface delamination peak force >15 N |
Competitive Polyvinyl Alcohol (PVA) for 3D Bioprinting Scaffolds prices that fit your budget—flexible terms and customized quotes for every order.
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Polyvinyl alcohol (PVA) for 3D bioprinting scaffolds is a water-soluble synthetic polymer derived from the hydrolysis of polyvinyl acetate, available in a graded spectrum of molecular weights and hydrolysis levels that dictate its processing window in extrusion-based additive manufacturing. Commercial grades such as Mowiol 10-98 (Mw ~61,000 Da, hydrolysis 98.0–98.8 mol%), Mowiol 20-98 (Mw ~125,000 Da), and Kuraray Poval 217 (Mw ~44,000 Da, hydrolysis 87–89 mol%) represent the range routinely evaluated for tissue-engineering constructs. The degree of hydrolysis exerts first-order control over aqueous solubility and crystallinity: partially hydrolyzed grades (86–89 mol%) retain residual acetate groups that disrupt hydrogen bonding, reducing the critical solution temperature and permitting ambient-temperature dissolution, while fully hydrolyzed grades (≥ 98 mol%) require heating to 85–95 °C for complete solubilization and yield stiffer, more crystalline hydrogels after physical crosslinking. A 10% w/v solution of a medium-molecular-weight PVA (viscosity-average Mw 61,000–70,000) produces a shear-thinning fluid with zero-shear viscosity in the range 1.5–3.0 Pa·s at 25 °C, a prerequisite for consistent microfilament deposition through nozzles of 150–250 µm internal diameter. The processing window is narrow: a concentration deviation of ±0.5 wt% can shift the flow behaviour index n from the target 0.35–0.45 to an excessively elastic regime that produces die swell, or to a low-viscosity state in which deposited strands spread beyond the ±10 µm lateral tolerance required for multi-layer architectures.
| PVA Grade | Mw (Da) | Hydrolysis (mol%) | 4% aq. viscosity (mPa·s, 20 °C) | Typical Bioprinting Application |
|---|---|---|---|---|
| Mowiol 4-98 | 27,000 | 98.0–98.8 | 4.0–5.0 | Low-viscosity carrier for cell-laden inks |
| Kuraray Poval 217 | 44,000 | 87–89 | 20.0–24.0 | Sacrificial fugitive ink |
| Mowiol 10-98 | 61,000 | 98.0–98.8 | 10.0–14.0 | Structural bioink for hard-tissue scaffolds |
| Mowiol 20-98 | 125,000 | 98.0–98.8 | 30.0–40.0 | High-stiffness constructs (osteochondral) |
Extrusion fidelity in PVA bioinks correlates with the linear viscoelastic region measured by small-amplitude oscillatory shear on a controlled-stress rheometer (Anton Paar MCR 302, 25 mm parallel plate, 1.0 mm gap, 25 °C). For a 10% w/v Mowiol 10-98 solution, the storage modulus G′ surpasses the loss modulus G″ at a crossover frequency of 0.8–1.2 rad·s⁻¹, indicating the onset of solid-like behaviour critical for shape retention after nozzle exit. The phase angle δ at 10 rad·s⁻¹ must lie between 15° and 25°; if δ exceeds 28°, the extruded filament sags under its own weight, causing lateral spreading that erases inter-filament porosity. Below 12°, the high elastic component increases the first normal stress difference, manifesting as annular die swell that expands the filament diameter by 15–25% beyond the nozzle gauge. These boundary conditions define a printable gel stiffness corresponding to a complex modulus |G*| of 2.0–4.5 kPa at the printing frequency, itself a function of print speed. On a pneumatic extrusion printer (Cellink BIO X, 3 mL syringe, 27G conical nozzle, 15 kPa pressure), a speed of 12 mm·s⁻¹ converts to a wall shear rate of approximately 150 s⁻¹, well into the shear-thinning plateau where n ≈ 0.38. When the same ink is processed at 8 mm·s⁻¹, the reduced shear rate elevates the apparent viscosity by 30%, requiring a pressure increase to 21 kPa to maintain volumetric flow—a compensation that is nonlinear because the ink’s thixotropic recovery time of 2.5 s (from step-stress experiments) is comparable to the inter-layer deposition interval, creating a memory effect that complicates process control. Published data for the exact pressure–viscosity transfer function across multiple commercial bioprinters is limited; bench-top studies consistently show that within-run viscosity drift of ±3% results in detectable filament width variation exceeding 5 µm, which is the acceptable tolerance for 100 µm stacked strands.
Physical gelation via cyclic freeze-thaw processing is the most widely adopted post-printing crosslinking strategy for PVA scaffolds intended for long-term culture. The mechanism—cryoconcentration of polymer chains between growing ice crystallites—forces interchain hydrogen bonding and the formation of semicrystalline junction zones with characteristic melting endotherms at 60–75 °C in differential scanning calorimetry (TA Instruments Q2000, heating rate 10 °C·min⁻¹). In a typical protocol, printed constructs are placed in a programmable freezer (Thermo Scientific TSC Series) and subjected to 1–5 cycles of freezing at -20 °C for 8 h and thawing at 4 °C for 4 h. The cooling rate during the freezing ramp must be controlled to 1.0 ± 0.2 °C·min⁻¹; faster cooling (>5 °C·min⁻¹) produces dendritic ice of sub-10 µm thickness that yields smaller pores (5–15 µm) after lyophilisation, whereas the target pore size for osteoblast infiltration is typically 100–300 µm. After 3 cycles, the crystallinity index by wide-angle X-ray diffraction reaches 32–35%, corresponding to a hydrated tensile modulus of 5.2 ± 0.8 MPa and elongation at break of 320 ± 40% (dogbone specimens tested per ASTM D638-14, Instron 5943 with 5 N load cell, PBS immersion at 37 °C). A property cliff-edge appears beyond 4 cycles: the elongation at break plummets to below 80% while the modulus reaches a plateau near 7 MPa, indicating that further crystallite growth reduces the amorphous chain-segment mobility without additional stiffening—a regime that produces brittle failure under the cyclic compressive strains (5–10% at 1 Hz) expected in articular cartilage applications. Chemical crosslinking with glutaraldehyde (0.25% v/v in acidic methanol, 2 h, followed by extensive washing in glycine buffer) elevates the modulus to 12.4 MPa but at the cost of residual aldehyde groups requiring verification by HPLC below the 0.5 ppm threshold set by ISO 10993-5 cytotoxicity testing on L929 fibroblasts.Sterilising PVA scaffolds without compromising dimensional fidelity remains a process bottleneck. Autoclaving (121 °C, 15 min) dissolves constructs that have undergone fewer than 3 freeze-thaw cycles because the amorphous regions remain soluble at elevated temperatures; scaffolds crosslinked with 3–4 cycles exhibit 15–20% shape distortion (measured by the change in strut intersection angles) and a 40% reduction in compressive modulus due to partial melting of secondary crystallites. Ethylene oxide (EtO) sterilisation per ISO 11135:2014 with a 4 h exposure at 55 °C and 30% relative humidity avoids thermal damage, but residual EtO must be reduced to ≤4 µg·g⁻¹ through forced-air degassing for a minimum of 48 h, validated by gas chromatography according to ISO 10993-7:2008. Gamma irradiation at a dose of 25 kGy produces chain scission that reduces Mw for Mowiol 10-98 from 61,000 to approximately 37,000 Da (GPC with PEO standards), thereby halving the zero-shear viscosity and rendering the pre-printed calibration obsolete. Electron-beam irradiation at 15 kGy generates a similar degree of main-chain rupture, though penetration depth limits applicability to constructs thinner than 5 mm. Aseptic preparation—dissolving PVA powder in sterile water-for-injection (WFI) and sterile-filtering the solution through a 0.2 µm PVDF membrane—is feasible only for grades with Mw below 30,000 Da because the solution viscosity of 10% w/v Mowiol 4-98 remains below 100 mPa·s, whereas higher-Mw solutions clog filtration membranes at operating pressures below 3 bar.
Mass loss of PVA scaffolds under flow conditions is governed by the crosslink density of the crystallite network rather than bulk hydrolysis of the carbon-carbon backbone. Static immersion in phosphate-buffered saline (PBS, pH 7.4, 37 °C) reveals a biphasic profile: an initial burst release of uncrosslinked chains accounts for 10–15% mass loss within the first 24 h, followed by a pseudo-linear rate of 0.3–0.5% per day over 28 days for scaffolds subjected to 3 freeze-thaw cycles. When the same scaffolds are mounted in a perfusion bioreactor (flow rate 0.5 mL·min⁻¹, corresponding to a wall shear stress of 5 × 10⁻³ Pa), the linear-phase mass loss rate doubles to 0.8–1.0% per day due to enhanced erosion at the pore walls, as evidenced by scanning electron microscopy showing rounding of strut edges after 14 days. For chemically crosslinked scaffolds (glutaraldehyde 0.25% v/v), mass loss over 60 days remains below 5%, but the swelling ratio increases from an initial 280% to 410% by day 30, altering the effective pore size for nutrient transport. ASTM F2902-16 provides a framework for assessing the degradation rate of absorbable implants; although PVA is not enzymatically degradable in mammals, the standard’s gravimetric method is routinely adapted with the note that terminal elimination requires chain oxidation and renal clearance, which limits PVA homopolymer scaffolds to non-resorbable applications unless blended with gelatin or collagen at 20–40 wt%.
In contrast to gelatin methacryloyl (GelMA), PVA does not require a UV photoinitiator—typically Irgacure 2959 at 0.05–0.1% w/v—that generates free radicals causing a documented 15–20% drop in encapsulated chondrocyte viability (live/dead assay, 24 h post-printing). GelMA’s crosslinking kinetics depend on UV intensity uniformity; even a ±2 mW·cm⁻² variation across the build platform results in differential stiffness that warps multi-layer constructs. Alginate bioinks crosslinked by immersion in 100 mM CaCl₂ lose structural integrity within 7–10 days in culture medium because monovalent cations in the medium exchange with the divalent calcium, whereas PVA physical crystallites remain stable for >8 weeks. Poly(ethylene glycol) diacrylate (PEGDA) offers fast photopolymerisation and adjustable modulus from 10 kPa to 1 MPa by altering molecular weight between crosslinks, but its inherent bioinertness necessitates RGD peptide grafting (e.g., 1 mM RGD-SH) to promote cell attachment, while partially hydrolysed PVA presents hydroxyl groups that adsorbs fibronectin from serum-containing media. A direct comparison of these materials is summarised below.| Property | PVA (Freeze-Thaw) | GelMA (UV) | Alginate (CaCl₂) | PEGDA (UV) |
|---|---|---|---|---|
| Crosslinking stimulus | Thermal cycling (phys.) | UV 365 nm, 5–10 mW·cm⁻² | Ionic diffusion | UV 365 nm, 5–10 mW·cm⁻² |
| Pre-crosslinking viscosity (Pa·s) | 1.5–3.0 (10%) | 0.05–0.5 (5%) | 0.3–2.0 (2–4%) | 0.01–0.1 (10%) |
| Post-crosslinking modulus (kPa) | 100–7,000 | 5–100 | 10–250 | 10–1,000 |
| Degradation time (weeks) | > 8 (stable) | 4–12 (enzymatic) | 1–4 (ion exchange) | 8–26 (hydrolytic) |
| Cell viability post-process | > 90% (if 3 cycles) | 70–85% | 85–95% | 80–90% |
| Key standard references | ISO 10993-5, ASTM F2902 | ISO 10993-5 | ISO 10993-5 | ISO 10993-5 |
Scaling PVA scaffold fabrication from benchtop to multi-head production lines (e.g., an 8-nozzle custom gantry system with independent pressure regulators) exposes the sensitivity of the process to lot-to-lot variation in raw PVA powder. Hydrolysis degree and molecular weight are controlled by the manufacturer to a stated tolerance of ±0.5 mol% and ±5,000 Da, respectively, yet these specifications translate to a zero-shear viscosity range spanning 1.8–2.8 Pa·s for a 10% w/v solution of Mowiol 10-98 across five separate production lots, as measured by Brookfield DV-II+ Pro viscometer (spindle CP-40, 25 °C, 10 s⁻¹). On a pneumatic dispenser with a 200 µm nozzle, the required pressure to achieve a volumetric flow rate of 0.8 µL·s⁻¹ shifts from 18 kPa to 27 kPa between the low- and high-viscosity lots. Unless the pressure setpoint is adjusted lot-specifically, the filament diameter varies by ±8 µm, exceeding the 5 µm tolerance band that preserves the designed inter-filament pore size of 100 µm. Closed-loop control using an inline pressure sensor (Honeywell 26PC series, 0–30 psi) and real-time machine vision filament width feedback has been demonstrated in academic pilot lines, but industrial adoption remains limited. In the absence of active compensation, pre-formulation blending of multiple lots to a target viscosity of 2.2 ± 0.1 Pa·s is recommended, which adds a mixing step of 12 h at 500 rpm under vacuum (-0.8 bar) for homogenisation and degassing—a non-trivial millroom operation that should be factored into overall manufacturing costs under ISO 13485 quality management planning.
