A degradable magnesium-alloy or PLLA-polymer screw/plate stabilizes a fracture, then dissolves as bone heals — no removal surgery
Fracture fixation hardware exists to hold broken bone fragments in a stable, aligned position while biology does the actual healing. For decades that hardware was titanium or stainless steel — strong, reliable, and permanent. Bioresorbable fixation keeps the strength where it is needed at implantation, but engineers it to disappear once the bone no longer needs it.
A displaced or unstable fracture will not heal predictably if the fragments keep moving relative to each other. Screws, plates, pins, and intramedullary rods restore alignment and hold the fragments still enough — either through absolute stability (rigid compression, allowing direct/primary bone healing) or relative stability (controlled micromotion, allowing callus-mediated/secondary bone healing).
The implant's job in the first days to weeks after surgery is almost entirely mechanical: resist bending, torsion, and shear at the fracture line so that the biological healing cascade — hematoma organization, inflammatory cell recruitment, callus formation — can proceed undisturbed.
Titanium and stainless steel plates are excellent at their mechanical job, but their permanence creates three well-documented downstream problems:
• Stress shielding: metal is 10–20× stiffer than cortical bone. A rigid plate carries so much of the load that the underlying bone is mechanically "unloaded," and by Wolff's Law, unloaded bone resorbs — leading to localized osteopenia that can persist for the life of the implant.
• Second removal surgery: hardware removal is common for symptomatic implants (pain, prominence, infection risk) or simply patient preference, adding a second anesthetic, incision, recovery period, and cost — for a device whose job is already finished.
• Pediatric growth interference: in a growing skeleton, permanent implants near a growth plate (physis) can tether or restrict normal bone lengthening, sometimes causing angular deformity — a strong driver for planned, often premature, hardware removal in children.
Studies of rigid titanium plating report up to 10–15% loss of cortical bone density directly beneath the plate within 12 months, driven almost entirely by stress shielding — bone that is never mechanically challenged never gets the biological signal to stay strong.
Two material families dominate clinical bioresorbable fixation:
• Magnesium alloys (e.g., WE43, Mg-Ca-Zn systems): a structural metal with elastic modulus (~45 GPa) far closer to cortical bone (~20 GPa) than titanium (~110 GPa), reducing stress shielding even before degradation begins. Corrodes electrochemically in physiological fluid.
• Bioresorbable polymers (PLLA, PLGA, PDLLA): synthetic polyesters that hydrolyze in water into lactic/glycolic acid, metabolites the body already processes via the Krebs cycle. Lower initial strength than metal but tunable degradation via molecular weight and crystallinity.
Both are selected and engineered so their initial mechanical strength is sufficient to stabilize the fracture — then calibrated to fade on a schedule that roughly mirrors the bone's own healing timeline.
In the first weeks after fixation, newly forming tissue at the fracture site contributes almost nothing mechanically. This is the phase where a bioresorbable implant must behave most like a permanent one: retaining nearly all of its as-implanted strength so the biology of secondary bone healing can proceed undisturbed.
Most fractures fixed with relative stability heal by secondary (callus-mediated) bone healing, a staged biological cascade:
1. Hematoma & inflammation (days 1–7): a fibrin-rich clot forms at the fracture gap; platelets and inflammatory cells release cytokines (TNF-α, IL-1, IL-6) that recruit mesenchymal stem cells.
2. Soft (fibrocartilaginous) callus (weeks 2–3): mesenchymal cells differentiate into chondrocytes and fibroblasts, forming a soft, collagen- and cartilage-rich bridge across the fracture — mechanically weak but stabilizing.
3. Hard callus / endochondral ossification (weeks 4–12): the cartilaginous callus calcifies and is progressively replaced by woven bone via endochondral ossification, dramatically increasing local stiffness.
4. Remodeling (months to years): woven bone is remodeled into mature lamellar bone along the lines of mechanical stress.
During the hematoma and soft-callus phases, the fracture site has negligible intrinsic stiffness — essentially all resistance to bending and torsion comes from the implant. If a bioresorbable implant were to lose significant strength during this window, the result can be excessive interfragmentary motion, disrupted callus formation, delayed union, nonunion, or outright fixation failure.
This is why bioresorbable implants are engineered with a deliberate "lag phase": a period of near-constant strength retention before degradation accelerates, timed to outlast the mechanically vulnerable early-healing weeks.
Both magnesium alloys and PLLA polymers can be engineered to retain roughly 90% or more of their initial bending strength through the first 6–8 weeks post-implantation — the period identified as mechanically critical across most long-bone and small-bone fracture models.
The near-full-strength lag phase is not accidental — it is designed into the material at multiple levels:
• Magnesium: high-purity alloying (removing Fe, Ni, Cu impurities that create galvanic micro-corrosion sites) and protective conversion coatings (MgF₂, plasma electrolytic oxidation) slow the initial corrosion rate.
• PLLA/PLGA: higher molecular weight and higher crystallinity slow water penetration and initial ester-bond hydrolysis; copolymer ratio (lactide:glycolide) and processing (extrusion vs. self-reinforcement) further tune the onset of strength loss.
The goal in both cases is the same: hold the line mechanically until the bone itself starts taking over.
Once the early-healing lag phase ends, the implant begins losing material and strength in earnest. The chemistry differs completely between the two material classes — hydrolytic chain scission for polymers, electrochemical corrosion for magnesium — but both convert a solid mechanical device into byproducts the body can process.
Polylactide implants degrade by hydrolysis of their ester backbone: water diffuses into the polymer bulk and cleaves ester bonds, chopping long polymer chains into progressively shorter fragments. This is largely a bulk-erosion process — degradation occurs throughout the implant's volume, not just at the surface — and it is autocatalytic: acidic degradation products (lactic acid oligomers) trapped inside the bulk locally lower pH and accelerate further hydrolysis.
Mechanical strength falls well before mass is lost, because strength depends on molecular weight (chain entanglement), not on how much material is physically still present. Semicrystalline regions resist water penetration longer than amorphous regions, giving a biphasic degradation curve. Eventually short oligomers and monomeric lactic acid are small enough to be phagocytosed or diffuse away, entering normal metabolism via the Krebs cycle to CO₂ and water.
Magnesium corrodes in chloride-rich physiological fluid through a straightforward electrochemical reaction:
Mg → Mg²⁺ + 2e⁻ (anodic dissolution) 2H₂O + 2e⁻ → H₂↑ + 2OH⁻ (cathodic reaction)
The net effect: the metal dissolves into biologically tolerable Mg²⁺ ions (magnesium is an essential nutrient — the body already contains ~25 g), while hydrogen gas and local alkalinization are produced as byproducts.
Hydrogen gas evolution is the signature engineering challenge for magnesium fixation: if gas is produced faster than surrounding tissue can diffuse and clear it, it can accumulate as visible subcutaneous gas cavities on imaging — usually harmless and self-resolving over weeks, but a source of patient concern and occasional local tissue pressure effects, especially in low-vascularity or subcutaneous sites.
Alloying with calcium, zinc, and rare-earth elements (yttrium, neodymium — as in WE43) refines grain structure and forms a more protective surface film, substantially slowing both corrosion rate and gas evolution compared to pure magnesium.
A corroding magnesium implant releases roughly 1 mL of hydrogen gas for every ~1 mg of magnesium consumed. Alloy design, surface coatings, and implant geometry are all tuned specifically to keep this evolution rate below what surrounding tissue can passively absorb and clear.
The central design challenge of bioresorbable fixation is kinetic matching: the implant's strength-retention curve should roughly mirror the inverse of the bone's own stiffness-gain curve. Degrade too fast, and the implant fails before the callus is mechanically competent — risking refracture or nonunion. Degrade too slowly, and the device re-creates the very stress-shielding problem it was designed to solve, while still eventually requiring the body to process a large, still-substantial mass of material.
Because healing speed varies enormously by patient (pediatric vs. elderly, diaphyseal vs. metaphyseal, comorbidities like diabetes or osteoporosis), material and even alloy/copolymer selection is increasingly individualized to the expected healing trajectory.
As the implant erodes and the callus mineralizes, the mechanical roles gradually invert: the device that once carried nearly all the load becomes a diminishing contributor, while the healing bone becomes the primary load-bearing structure — precisely the stimulus mature bone needs to remodel to full strength.
Bone is a living, mechanically responsive tissue: Wolff's Law describes how bone continuously remodels in response to the loads it experiences, reinforcing itself along lines of habitual stress and resorbing where load is chronically absent. A permanent, rigid implant that never releases its grip on the load path denies the healing bone the mechanical stimulus it needs to fully remodel and regain native strength — the mechanistic root of stress shielding.
Bioresorbable fixation is explicitly designed to avoid this trap: rather than shielding bone indefinitely, it hands mechanical responsibility back on a schedule.
Conceptually, implant and bone share the mechanical load at the fracture site as two parallel structural elements. Early on, the implant carries nearly 100% of the load because the callus contributes almost none. As weeks pass, two curves move in opposite directions: implant strength declines (hydrolysis/corrosion) while callus stiffness rises (mineralization, endochondral ossification, remodeling). Somewhere in the mid-healing window, the curves cross — the majority of load transfers to bone — and by the time the implant is fully resorbed, bone alone is fully load-bearing.
A well-matched material produces a smooth handoff without ever exposing the healing site to a sudden, unsupported overload.
For a magnesium-alloy implant in an average healer, load-bearing majority typically shifts to the callus somewhere between weeks 8 and 16 — broadly coincident with the transition from soft callus to mineralized hard callus in the underlying biology.
Kinetic mismatch in either direction has real clinical consequences:
• Too-fast degradation: the implant loses mechanical integrity before the callus is competent to compensate, risking loss of reduction, hardware fracture, delayed union, or refracture.
• Too-slow degradation: prolonged stress shielding persists well past the point bone could be remodeling independently — and if resorption stalls or is incomplete, the implant may still eventually require attention, undermining the primary rationale for choosing it.
This is why material choice (magnesium vs. polymer), alloy/copolymer formulation, and implant geometry are matched to the expected patient healing rate — a fast-healing pediatric fracture and a slow-healing osteoporotic fracture are, mechanically, very different design problems.
The final stage is the entire point of the technology: the implant disappears completely, metabolized or excreted through normal physiological pathways, leaving behind bone that has fully healed and remodeled under its own mechanical stimulus — with no second surgery ever required to remove hardware.
By complete resorption, the mechanism differs by material but the outcome converges: nothing foreign is left behind.
• PLLA/PLGA: once hydrolysis fragments the polymer down to short oligomers and monomeric lactic/glycolic acid, these small molecules are cleared by local phagocytosis or diffusion into circulation, entering the same metabolic pathways the body uses for lactate from exercise — ultimately oxidized to carbon dioxide and water.
• Magnesium alloys: continued corrosion converts the remaining metal to soluble Mg²⁺ ions, which join the body's existing magnesium pool (~25 g, mostly in bone and muscle) and are excreted renally like any other physiological magnesium. In parallel, the space the implant occupied is progressively invaded and replaced by new bone in a process resembling creeping substitution seen with bone graft incorporation.
Bioresorbable fixation has found its clearest clinical footing where the drawbacks of permanent metal are most acute and mechanical demands are modest:
• Pediatric fracture fixation: avoids both a routine second removal surgery and the risk of growth-plate tethering from hardware left in a still-growing skeleton.
• Craniomaxillofacial (CMF) surgery: skull and facial fixation benefits from avoiding palpable hardware, growth restriction in a growing skull, and metal-artifact interference on future CT/MRI imaging.
• Small bone and hand/foot fixation: lower mechanical demands than major long-bone fixation make these an excellent fit for current-generation polymer and magnesium screw systems.
• Sports medicine: bioabsorbable interference screws for ligament reconstruction (e.g., ACL) have the longest clinical track record of any resorbable fixation device.
Bioresorbable fixation is not yet a universal replacement for permanent metal. Large, high-load diaphyseal long-bone fixation still often demands strength retention longer and higher than current magnesium alloys or polymers reliably provide. Predicting individual degradation timelines precisely enough for surgical planning remains imperfect. And regulatory pathways, while maturing, still require extensive case-by-case validation.
Active development areas include higher-strength magnesium–rare-earth alloys, patient-specific 3D-printed geometries, degradation-rate-tuned surface coatings, and computational models that predict an individual patient's crossover point from implant-dominant to bone-dominant load-bearing — pushing bioresorbable fixation toward broader use across the skeleton.
Eliminating routine hardware-removal surgery is estimated to avoid roughly 15–20% of pediatric orthopedic reoperations — sparing patients a second anesthetic, incision, infection risk, and recovery period for a device whose mechanical job was already finished.