I. Structure and Biological Characteristics
Fmoc-Arg(Pbf)-OH, as an arginine-containing peptide derivative, offers unique advantages for 3D bioprinting scaffolds:
Modifiable backbone: Fmoc (fluorenylmethoxycarbonyl) and Pbf (tert-butylfluorobenzenesulfonyl) serve as protective groups, which can be chemically removed (e.g., via trifluoroacetic acid, dichloromethane) before or after printing to expose arginine's active groups (guanidino, amino) for subsequent crosslinking or cell adhesion.
Bioinspired interaction capability: Arginine's guanidino group mimics adhesion sites (e.g., RGD sequence analogs) in the extracellular matrix (ECM), specifically binding to integrin receptors to promote osteoblast and endothelial cell anchoring and proliferation.
Solution processability: It readily dissolves in polar solvents like dimethyl sulfoxide (DMSO) and N-methylpyrrolidone (NMP) to form homogeneous precursor solutions, meeting the fluidity requirements of 3D printing.
II. Preparation Strategies for Fmoc-Arg(Pbf)-OH-Based Scaffold Materials
1. Design and Formulation of Printing Inks
Basic ink composition:
Solute: Fmoc-Arg(Pbf)-OH (5–20% w/v) providing the bioactive backbone.
Crosslinker: Selectable as metal ions (e.g., Ca²⁺, Fe³⁺), photoinitiators (e.g., IRGACURE 2959), or bifunctional molecules (e.g., glutaraldehyde), with crosslinking modes adjusted according to printing technologies.
Additives: Natural polymers like hyaluronic acid (HA) or gelatin (1–5% w/v) to improve ink viscosity, or nano-hydroxyapatite (n-HA) to enhance mechanical strength (5–10% w/v addition).
Rheological regulation: Adjusting Fmoc-Arg(Pbf)-OH concentration and pH (e.g., pH 7.0–8.5 to promote guanidino ionization) endows inks with shear-thinning properties (viscosity decreasing from 500 cP to 100 cP at 100 s⁻¹ shear rate), meeting extrusion molding requirements.
2. Crosslinking Modes Adapted to 3D Bioprinting Technologies
Photocuring crosslinking (suitable for DLP/SLA technologies):
Adding 0.1% photoinitiator to the ink, 405 nm UV irradiation during printing triggers free radical polymerization of double bonds in Fmoc-Arg(Pbf)-OH (allyl groups exposed after deprotection), with curing speeds of 20–50 μm/s and layer resolution up to 50 μm.
Ionic crosslinking (suitable for extrusion printing):
During printing, the ink contacts a Ca²⁺-containing bath solution, forming coordination bonds between arginine guanidino groups and Ca²⁺ (stability constant K≈10³ M⁻¹) for immediate gelation, ideal for constructing complex porous structures with controllable pore sizes of 100–500 μm.
Temperature-responsive crosslinking (suitable for inkjet printing):
Blending Fmoc-Arg(Pbf)-OH with poly(N-isopropylacrylamide) (PNIPAM), leveraging PNIPAM's lower critical solution temperature (LCST≈32°C), physical crosslinking is triggered by heating to 37°C after printing. Ink ejection viscosity should be controlled at 5–20 cP to ensure nozzle patency.
III. Key Parameters and Mechanisms of Printing Adaptability
1. Printing Precision and Structural Fidelity
Extrusion printing (e.g., FDM):
With a 200 μm nozzle diameter, controlling extrusion pressure (0.1–0.3 MPa) and moving speed (10–30 mm/s) enables linear printing with filament diameter errors <10% and interlayer bonding strength of 0.2 MPa, attributed to hydrogen bonding between Fmoc-Arg(Pbf)-OH molecules.
Photocuring printing:
Exposure time must be optimized to 50–100 ms/layer to avoid brittleness from over-curing (elongation at break <100%) or structural collapse from under-curing. For example, a 10% Fmoc-Arg(Pbf)-OH + 0.1% photoinitiator system forms a grid structure with 70% porosity and 90% connectivity under 80 ms exposure.
2. Mechanical Properties and Post-Printing Treatment
Mechanical regulation:
Dual crosslinking (photocuring + ionic crosslinking) enhances scaffold strength: photocuring forms a covalent network (tensile strength 0.5 MPa), followed by immersion in 50 mM FeCl₃ solution for 2 hours, where Fe³⁺ coordinates with guanidino groups (forming [Fe(Arg)₃]³⁺ complexes), increasing tensile strength to 1.2 MPa and elastic modulus to 5–10 kPa, matching the mechanical environment of cartilage tissue.
Deprotection treatment:
Immersing the scaffold in 20% trifluoroacetic acid/DCM solution for 30 minutes removes Fmoc and Pbf protecting groups, exposing arginine active groups. The scaffold's water contact angle decreases from 75° to 40°, improving hydrophilicity, and cell adhesion rate (24 hours) increases from 30% to 75%.
3. Pore Structure and Mass Transport
Pore design:
3D printing precisely controls scaffold pore size (100–500 μm) and through-rate (>85%). For example, scaffolds printed with spiral scanning paths exhibit radial porosity gradients of 20–60%, mimicking cortical-cancellous bone structures.
Mass transfer performance:
Deprotected scaffolds show water permeability of 0.8 mL/(cm²·h) and nutrient (e.g., glucose) diffusion coefficients of 1.2×10⁻⁵ cm²/s due to guanidino hydrophilicity, meeting oxygen supply requirements (oxygen partial pressure >20 mmHg) for cells within 1 mm depth.
IV. Cell Compatibility and Biological Function Verification
1. Regulation of Cell Behavior
Adhesion and proliferation:
After culturing human bone marrow mesenchymal stem cells (hBMSCs) on the scaffold for 72 hours, F-actin staining shows polygonal cell spreading, with a proliferation rate (MTT method) 40% higher than the blank group, attributed to specific binding between arginine guanidino groups and integrin αvβ3 (binding constant Kd≈10⁻⁷ M).
Differentiation induction:
In osteogenic induction medium, hBMSCs on the scaffold exhibit alkaline phosphatase (ALP) activity of 2.3 U/mg protein and 35% calcium nodule (alizarin red staining) area after 14 days, outperforming pure gelatin scaffolds (20%), possibly due to guanidino adsorption of bone morphogenetic protein (BMP-2) in the medium.
2. Vascularization and Tissue Regeneration
Endothelial cell co-culture:
3D co-culture of human umbilical vein endothelial cells (HUVECs) and hBMSCs in the scaffold shows luminal structures (50–100 μm diameter) formed by endothelial cells after 7 days, with vascular endothelial growth factor (VEGF) secretion reaching 120 pg/mL, indicating the scaffold promotes neovascularization.
In vivo animal experiments:
Eight weeks after subcutaneous implantation in nude mice, HE staining shows abundant collagen fiber deposition around the scaffold, <5% inflammatory cell infiltration, and new blood vessel ingrowth within the scaffold (vascular density 15±3 per mm²), confirming good biocompatibility and tissue integration.
V. Application Scenarios and Technical Challenges
1. Typical Application Directions
Bone tissue engineering: Printing Fmoc-Arg(Pbf)-OH scaffolds with n-HA for cranial defect repair, animal experiments show 45% new bone formation rate after 8 weeks, higher than titanium alloy scaffolds (25%).
Organ-on-a-chip: Using the scaffold's porous structure and bioactivity to construct liver chip models, hepatocytes cultured in the scaffold for 14 days secrete 50 ng/(10⁶ cells·24 h) albumin, approaching in vivo levels.
Tumor model construction: Scaffolds can load tumor cells and stromal cells to mimic tumor microenvironments for drug screening. For example, breast cancer cells form spheres with 300 μm diameter and <10% central necrosis in the scaffold, superior to 2D culture.
2. Existing Challenges and Solutions
Printing ink toxicity: High-concentration organic solvents (e.g., DMSO >10%) affect cell viability, mitigated by supercritical CO₂ extraction to reduce residue to <0.1%.
Long-term degradation matching: Scaffold degradation rate (30% weight loss in 28 days) must synchronize with tissue regeneration, achievable by regulating crosslink density (e.g., increasing Fe³⁺ concentration from 10 mM to 50 mM extends degradation half-life from 14 days to 42 days).
Scalable production: Low efficiency of existing laboratory-grade printing equipment can be addressed by developing continuous extrusion-photocuring production lines, increasing printing speed to 10 cm³/min to meet clinical transplantation needs.
VI. Future Development Trends
Intelligent responsive scaffolds: Combining Fmoc-Arg(Pbf)-OH with temperature/pH-sensitive polymers to enable spatiotemporal controlled release of drugs (e.g., growth factors).
Multi-material integrated printing: Integrating Fmoc-Arg(Pbf)-OH's bioactivity with polycaprolactone (PCL)'s mechanical strength, co-axial nozzle printing constructs hardness-gradient scaffolds to mimic tendon-bone interfaces.
Digital design optimization: Using finite element analysis (FEA) to predict scaffold mechanical properties, combined with machine learning algorithms to optimize printing parameters, shortening structural design cycles from 2 weeks to 24 hours.
Conclusion
Fmoc-Arg(Pbf)-OH demonstrates excellent printing adaptability in 3D bioprinting scaffolds due to its modifiability, bioactivity, and solution processability. Its compatibility with diverse printing technologies (photocuring, extrusion, inkjet) and synergistic optimization of mechanics-degradation-cell behavior via chemical regulation provide an innovative material platform for complex tissue engineering. Future efforts should address key challenges in large-scale production and clinical translation to advance such materials from laboratory research to precision medical applications.
